BIOCERAMICS AND THE HUMAN BODY

Proceedings of the International Congress on Bioceramics and the Human Body held in Faenza, Italy, 2-5 April, 1991, organised by the IRTEC-CNR Institute in collaboration with Agenzia Polo Ceramico.

INTERNATIONAL SCIENTIFIC COMMI1TEE

A. RA VAGLIOLI (IRTEC-CNR, Faenza-I), the President of the Congress

G. BERGER (Academy of Science, Berlin--D) W. BONFIELD (Biomaterials Dept., Queen Mary & Westfield College-IRC, University of London-UK) E.Y. CHAO (Mayo Clinic, Rochester-USA) P. CRISTEL (European Society for Biomaterials, Paris-F) K. DE GROOT (University of Leiden-NL) G. DE MARIA (Universitl"La Sapienza" di Roma-I) C. DOYLE (How Medica International, Staines-UK) P. DUCHEYNE (Pennsylvania University, Philadelphia-USA) U. GROSS (European Society for Biomaterials, Berlin-D) G. HASTINGS (Biomaterials Dept., Queen Mary & Westfieid College-IRC, University of London-UK) G. HEIMKE (Clemson University-USA) S. HULBERT (American Society of Biomaterials Terre Harte-USA) A. KRAJEWSKI (IRTEC-CNR, Faenza-I) P.G. MARCHETTI (Istituti Ortopedici Rizzoli, Bologna-I) G. MURATORI (Gruppo Italiano Studi Implantari, Bologna-I) A. MORONI (Universitl di Bologna-I) H. OONISHI (Japanese Society for Biomaterials, Osaka-J) F.J. SCHOEN (American Society for Biomaterials, Boston-USA) P. TRANQUILLI LEALI (Universitl Cattolica, Roma-I) D. WILLIAMS (University of Liverpool-UK) H. WOLF (University of Berlin-D)

ORGANIZING COMMl1TEE AND REPRESENTATIVES FROM SCIENTIFIC SOCIETIES

G.N. BABINI (IRTEC-CNR, Faenza-I) W. BONFIELD (Biomaterials Dept., Queen Mary & Westfield College, IRC, University of London, London-UK) L. CINI (Societl Ceramica Ita1iana, Bologna-I) S. CONTOLI (Ospedale Civile, Faenza-I) P. CHRISTEL (European Society of Biomaterials, Paris-F) S. CHAO (Mayo-Clinic, Rochester-USA) S. GIANNINI (Istituti Ortopedici Rizzoli, Bologna-I) P. GIUSTI (Societl Italiana Biomateriali, Siena-I) P.G. MARCHETTI (Istituti Ortopedici Rizzoli, Bologna-I) F. MAROTTI (Soc. Ital. Biomeccanica Ortopedia Traumatologia, Modena-I) A. RA V AGLIOLI (IRTEC-CNR, Faenza-I) C. ROVELLI (Agenzia Polo Ceramico, Faenza-I) R. SILVESTRINI (CNR Biomaterials Program, Roma-I) M. SPECTOR (Society for Biomaterials, Atlanta-USA) T. SAKURAI (Japanese Society for Biomaterials, Tokyo--J) P.S. WALKER (European Society of Biomechanics, London-UK)

BIOCERAMICS AND THE HUMAN BODY

Edited by

A. RA V AGLIOLI

and

A. KRAJEWSKI Research Institute tor Ceramic Technology, National Research Council, Faenza, Italy

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International Congress on Bioceramics and the Human Body (1991: Faenza, Italy) Bioceramics and the human body. I. Title 11. Ravaglioli, A. III. Krajewski, A. 610.28

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v

PREFACE

For a long time now ceramic materials have been produced for application in the chemical, steel, and glass industries in a variety of conditions, e.g. in the presence of high temperatures, reducing atmospheres, or corrosive liquids.

With the advent of the nulcear reactor certain ceramics have been required to resist high radiation fluxes, extreme temperature gradients, or degradation in the presence of corrosive liquids while remaining dimensionally and mechanically stable, without developing fissures or any other kind of incon­venience over many years.

The space age has turned to ceramics for a substantial contribution, for example in the case of missile and NASA shuttle coatings.

Nowadays even man's life may depend on the performance of certain ceramic materials used as components of surgical implants designed for application inside a living organism-an environment obviously very different from that of a nuclear reactor or outer space but equally (perhaps more) hostile.

These are in brief outline the materials discussed at the 'Bioceramics and the Human Body' Congress organized in Faenza with the involvement of 190 representatives from 21 countries.

The Americans and the East Germans could not attend the Congress, the former owing to the Gulf War, then in a fast-developing stage, the latter for the well-known economic and relocation problems arising from unification. Their absence was certainly unfortunate, also in consideration of the fact that we collaborate with them in research on bioceramics. The Congress was promoted and organized by the IRTEC-CNR Institute in collaboration with the Agenzia Polo Ceramico. The latter put at our disposal their entire personnei, whom we now wish to thank for their competence in coordinating a meeting of such a specialized nature.

Further collaboration came from various societies and institutes such as the European, the American, and the Japanese Societies for Biomaterials, the Istituti Ortopedici Rizzoli, and others. The aim of the Congress was to determine the state of affairs of ceramic science and technology with special reference to the capacity of ceramic materials for osteointegration, from the points of view of basic research and application. Special attention was addressed to the manufacturing of ceramics and their functional role in surgical applications.

vi

What the Congress achieved was to put forward proposals and compare the different successful or (so far very few) unsuccessful postoperative experiences carried out adopting ceramic prostheses in the shape of either bulk prototypes or coatings applied through highly sophisticated techniques (applications range from the maxillo-facial area to joint extremities).

In future perspective, the Congress identified the targets of a further improvement of the already very good interaction between tissues and ceramics (whether inert or bioactive) and of the introduction of new materials such as SiC, ShN4, TiN, TaN, etc. In addition, the tendency emerged to give priority to working for the definition 0/ new standards regulating raw materials and finished products.

A. RAVAGLIOLI

A. KRAJEWSKI

vii

CONTENTS

Preface............................................... ........................................ v

Introductory Report ..................................................................... 1 A. Ravaglioli and A. Krajewski

1. Surgery of the prosthetic implantations

Fifteen years' experiences with alumina-ceramic total hip-joint endo-protheses: a c1inical, histological and tribological analysis..................... 17

H. Plenk, Ir., M. Böhler, A. Walter, K. Knahr and M. Salzer Bio-functional adaptive behavior to ceramic implants .......................... 26

A. Toni, A. Sudanese, S. Stea, S. Squarzoni, P.P. Montina, A. Bueno Lozano, F. Calista, A. Pizzoferrato and A. Giunti

A review on the aseptic total hip replacement failures ......................... 35 G. Monticelli, L. Romanini, O. Moreschini

Solving of prosthetic problems through bioceramies ............................ 46 F. Zarotti

Hand MP joint implant arthroplasty: the state of the art. Advantages and disadvantages of the most common implant arthroplasties .................... 49

A. Caroli and S. Zanasi Alumina total joint replacement of the first metatarso phalangeal joint: a biomechanical study of the design ................................................... 62

S. Giannini, A. Moroni, A. Krajewski, A. Ravaglioli, R. Martinetti, C. Farina

Titanium hydroxylapatite coated metacarpo-phalangeal and interphalan-geal implant ................................................................................ 67

E.C. Marinoni, G. Venini and A. Ravaglioli Hydroxyapatite coated plate in the surgical treatment of the forearm non union with bone loss ..................................................................... 73

G.F. Zinghi, A. Moroni, L. Specchia, P. Bungaro, G. Gualdrini, G. Rollo, C. Sabato

A study on HA-coated titanium dental implants, part I: stress analysis of dental implant ............................................................................. 78

liyong Chen, liming Zhou, Xingdong Zhang, Deri Wan, Shaoan Wang and Anyu Chen

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A study on HA-coated titanium dental implants, part 11: coating properties in vivo, implant design and clinical evaluation ....................................... 89

liyong Chen, liming Zhou, Xingdong Zhang, Deri Wan, Shaoan Wang and Anyu Chem

Rehabilitation of radical mastoidectomy cavities with calcium phosphate ceramics ..................................................................................... 101

M. DeI Bo and A. Zaghis Clinical results of IMZ dental implants ............................................. 107

P. Passi Direct composite-ceramic restorations: a ciinical study ........................ 113

C. Prati, G. Montanari, E. Toschi and A. Savino Articulation of ceramic surfaces against polyethylene .......................... 118

R.M. Streicher, M. Semlitsch and R. Schön Hip anatomical uncemented ceramic arthroplasty (AN.C.A.): results at a 3-year follow-up ........................................................................... 124

L. Specchia,A. Moroni, L. Ponziani, G. Rollo, S. Pavone, V. Vendemia

2. Bioceramics: properties ud their technology of production

Fluorapatite and hydroxyapatite heat-treated coatings for dental implants 130 H. W. Denissen, H.M. de Nieuport, W. Kalk, H. G. Schaeken and V. van den Hooft

Porous titanium implants with and without hydroxyapatite coating ........ 141 A. Moroni, V. Caja, E. Egger, F. Gottsauner Wolf, L. Trinchese, G. Rollo and E. Y. Chao

Surface reactivity and biocompatibility of bulk glass and glass coatings ... 148 B. Locardi

Plasma spray systems for the deposition of materials for biomedical applications ................................................................................. 156

A. Salito, G. Barbezat, H. Filmer, I. Hochstrasser, A.R. Nicoll and F. Trotta

Bioceramics for maxillofacial applications ......................................... 166 I.G.C. Wolke, C.P.A.T. Klein and K. de Groot

Nucleation and growth of dicalcium phosphate dihydrate on titanium alloy substrates ................................................................................... 181

P. Royer, M. Freche, C. Rey Relationships between bulk and surface structure and biomaterial bio-compatibility ............................................................................... 189

A. Bertoluzza, M.A. Morelli and A. Tinti Experimental study on the properties of hydroxyapatite coated implants. 195

C. Gabbi, P. Borghetti, N. Antolotti, S. Pitteri Physico-chemical characterization of hydroxyapatite of unknown manufac-ture ........................................................................................... 203

A. Ravaglioli, A. Krajewski, A. Piancastelli and R. Martinetti

ix

Application of ceramic composites as implants: results and problems ..... 206 S.M. Barinov and Yu. V. Baschenko

Radioactivity measurements of zirconia powders ................................ 211 G. Capannesi, A.F. Sedda, C. Piconi, F. Greco

Microstructural analysis of bioceramic materials................................. 217 B. Bousjield

Yttria and calcia partially stabilized zirconia for biomedical applications. 223 P. Fassina, N. Zaghini, A. Bukat, C. Piconi, F. Greco and S. Piantelli

Mechanical properties of plasma sprayed ceramic coatings on orthopaedic implants ..................................................................................... 230

J. Rieu, H. Carrerot, J.-L. Aurelle, A. Rambert and G. Bousquet Plasma-spray coating of titanium supports with various ceramics: a study at the interface................................................................................ 236

A. Krajewski, A. Ravaglioli, V. Biasini, A. Martinetti, A. Piancastelli, S. Sturlese, S. Fioravanti, N. Antolotti, C. Mangano and F. Trotta

Microstructure and microanalysis of bioglasses and glass-ceramics from the MgO-CaO-P20S-Si02 system with Zr02............................................ 244

J. Ma. Rincon and P. Callejas The colour of aluminium oxide ceramic implants .............................. .. 250

G. Willmann Dental ceramics and composite resins as restorative materials............... 256

C. Prati, E. Toschi, C. Nucci, R. Mongiorgi, A. Savino Reinforced silver glass-ionomer cement and light-cured H.E.M.A. glass­ionomer cement under silver-amalgam restorations: a microleakage study 260

C. Nucci, E. Toschi, R. Mongiorgi and C. Prati Tensile bond strength of dental porcelain to dental composite resins ...... 265

R. Mongiorgi, C. Prati, E. Toschi and G. Bertocchi Glass-ionomer cements as base for composite restorations ................... 270

E. Toschi, R. Mongiorgi, C. Prati, G. Valdre and C. Nucci

3. Biological characterization and effects on bioceramics

Systemic control of tissue and cell reactions relating to ceramic implants 275 U. Gross, C. Müller-Mai and C. Voigt

In vitro cytocompatibility and tissue reaction to ceramics ..................... 285 A. Pizzoferrato, E. Cenni, G. Ciapetti, S. Savarino and S. Stea

Bioceramics in orthopaedic surgery: know how, status and preIiminary results ........................................................................................ 295

S. Giannini, A. Moroni, G. Coppola, L. Ponziani, A. Ravaglioli, A. Krajewski, A. Venturini, M. Pigato and D. ZaJfe

Osseointegration of hydroxyapatite-coated and uncoated bulk alumina implants in the femur of Göttingen minipigs: mechanical testing of bonding strength .......................................................................... 302

J. Orth, S. Macedo, A. Wilke and P. Griss

x

Thromboresistance of Ti6Al4V, coated with a thin film of turbostratic carbon, for cardiovascular applications ............................................. 308

M.A. Gatti, E. Monari, M. Dondi, G. Noera, G. Fattori, F. Vallana, S. Rinaldi, and E. Pasquino

TCP-impurities in HA-granules and crystalinity changes in plasma­ftamesprayed HA-coatings detected by spectroscopical methods and their consequences............................................................................... 317

M. Weinländer, H. Plenk, Jr., F. Adar and R. Holmes Longterm stability of TiN .............................................................. 321

M. Schaldach and A. Bolz In vitro toxicity of fine particles of hydroxyapatite .............................. 334

E.J. Evans and E.M.H. Clarke-Smith Differences in behaviour of cultured fetal rat osteoblasts upon bioglass and nonreactive glasses ................................................................. 340

W. C.A. Vrouwenvelder, C. G. Groot and K. de Groot Complement activation by ceramics ................................................. 345

I. Dion, A. Baquey, C. Baquey, T. Mesana, M. Pourtein, B. Candelon and J.-R. Monties

Sister chromatid exchanges (SCEs) and proliferation rate index (PRI): the application of cytogenetic methods in biocompatibility field . . . . . . . . . . . . . 353

M. Cannas, S. Biasiol, A. Masse, A. Ruggeri and R. Strocchi Surface coating of PECVD a-SiC:H to improve biocompatibility ........... 360

A. Bolz and M. Schaldach Bone tissue response to hydroxyapatite-coated and uncoated titanium wire-MESHS in an in.fected site: results of an animal experiment .......... 366

A. Wilke, J. Orth,' M. Kraft and P. Griss Biocompatibility of siliconcarbide and siliconnitride ceramics: results of an animal experiment.................................................................... 372

J. Orth, M. Ludwig, W. Piening, A. Wilke and P. Griss Effect of Ti02 ceramic precursors on human lymphocyte mitogenesis ..... 378

S. Piantelli, G. Maccauro, P. Fassina and A. Bukat Experimental study of two corals used as bone implant in the sheep ...... 383

P. Jammet, F. Bonnei, P. Baldet, F. Souyris and M. Huguet Interfacial study of some inert and active ceramics implanted in bone .... 388

D. ZaJfe, S. Giannini, A. Moroni, A. Krajewski and A. Ravaglioli Surface charge of the bioglass treated by a physiological solution .......... 396

S. Szarska In vivo study of a new active glass for bone repair: short term results .... 402

A.M. Gatti, D. ZaJfe and O. Anderson A study of hydroxyapatite ceramics and its osteogenesis ...................... 408

Zhang Xingdong, Zhou Pin, Zhang Jianguo, Chen Weiqun and Wu Chuong

Biological apatite as a material for artificial bone: a preliminary investi-gation on its possibility .................................................................. 417

K. Hirota

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4. Analytical techniques and standards

The design and manufacture of joint prostheses and stress distribution... 422 P. Dalla Pria

Biomechanical principles of the surgical treatment of the long bones non unions ........................................................................................ 432

G.F. Zinghi, L. Specchia, A. Moroni, G. Galli, G. Rollo and C. Sabato

Quality control and soviet standards for bioceramics ........................... 438 V.A. Dubok and L.L. Suhih

Physico-chemical techniques to characterize structure and composition of interfaces involving bioceramics: potentiality and limits ....................... 444

F. Garbassi and E. Occhiello Problems concerning the industrial production of alumina ceramic components for hip joint prostheses .... . ....... . ........ . ......... . ........ .. ....... 454

E. Dörre Bulk and thin film carbon materials for biomedical applications: quality control criteria and procedures ....................................................... 461

F. Vallana, P. Arru and M. Santi Development of a human bone marrow ceIl culture to test the cytocom­patibility of bulk hydroxyapatite materials......................................... 471

A. Wilke, f. Orth, P. Griss, V. Nehls and D. Drenchkhahn Finite element analysis of a ceramic hip-joint head and its failure mode due to a crack in the material......................................................... 477

E. Ravagli and E. Maggiore Research, planning and design of ear ossicle prototypes based on A12ü 3 ,

hydroxyapatite and Zrü2 ••••••..••••••••••••••••..••••••.•..••••••••.•.•••••••••••••••• 486 P. Laudadio, L. Presutti, A. Ravaglioli and R. Martinetti

Preliminary tests to determine the inftuence of sterilization and storage on compressive strength of hydroxyapatite cylinders ........................... 491

B. Gasser, W. Müller and R. Mathys, fr. A study of the methodology for treatment of titanium substrates to be coated with hydroxyapatite ............................................................ 497

A. Krajewski, A. Ravaglioli, R. Martinetti and C. Mangano Stress analyses of puIl-off tests for strength measurements of coatings .... 504

U. Soltesz, E. Baudendistel and R. Schäfer CeIl adhesion strength to bioceramics and its mathematical model......... 510

T. Tateishi and T. Ushida

Index of contributors ..................................................................... 517

1

SKELETAL I~PLANTS: FROM ~TALS, TO POLYMERS, TO CERAMICS

Dr. Antonio Ravaglioli and Dr. Adriano Krajewski (IRTEC-CNR, National Council of Research, Faenza, Italy)

INTRODUCTION

Skeletal implants, as used in orthopaedic or traumatic surge­ry, must firstly be "functional", i.e. able to successfully restore in an active or a passive way the physiological functions of the tissue, organ, or body part surgically replaced. Secondly, they must be very "durable", i.e. able to sustain all external or internal loads without fracturing; they must be able to transfer these loads to surrounding bone and/or soft tissues and to give rise to permanent i~plant fixation without causing any bone fracture; and they must have sufficient wear resistance for articulating surfaces.

All these prerequisites apply to endoprostheses for bone and/or joint replacement, to dental implants, to artificial tendons and ligaments, and to bone fracture fixation devices.

In order to improve the products for such surgical applications -- in terms of nlanning, design, and manufacture of skeletal im~lants -- it is necessary in the first place to accurately select the materials, paying special attention to implant design and geometry. The properties of a material are in fact decisive to carry out certain snecific designs expected to ensure imnlant durability, that is, a lifetime as long as possible.

Secondly, good manufacturing and a proper quality control of the end nroduct are essential. This is particularly imnort­ant in the case of articulating surfaces of joint endopros­theses, which are expected to provide low-friction character­istics over a long period of time. Surgical records prove that the cause of excessive wear ( abrasion) is not always an incorrect positioning of impl~~ts durin~ surgery, but also an inadequate design, surface finish, or material composition.

This points to the decisive importance of evaluating surgical records and storing all data from clinical and functional tests on patients to achieve maximum efficacy and

2

reliability of such a complex and expensive type of surgical intervention. Once correlation of such data on im plant per­formance is obtained, detection of the causes of failure over time should become possible on a statistical basis. Unfortuna­tely, most of the information currently available is about short-term interactions between tissues and skeletal implants, while there are only a few studies reporting on long-term performance, i.e. over more than 10 years. The reasons for failure were therefore primarily found -- much too hastily in mechanical factors such as stress from fatigue.

Today, skeletal implants in general and joint endopros­theses in particular can be manufactured in compact or porous form from three classes of mechanically and chemically dif­ferent materials: 1) metals and alloys; 2) polymers; 3) inert and bioactice ceramics, the so-called "bioceramics".

The aim of this conference is to assess the state of the art of "bioceramics", of which only some have the mechanical strength required, but all exhibit a "very good surface be­haviour". This means not just excellent tribological propor­tions for artieulating surfaees, but also specific surfaee eharaeteristics which favour the fixation of implants to bone and tissue.

If their meehanieal properties do not allow heavy-load bearing, inert and bioaetive ceramies ean be used as eoatings, for example on metal implants to provide fixation to bone.

Metals and polymers were onee the materials most used for skeletal implants. The greatest problems eneountered were in the first plaee mechanical durability and seeondly liabi­lity to corrosion , shown by metals and their alloys, not to ~ention the degradation of polymers in biological environment~

This would result in toxie reaetions to the constituents released from an implant, in sensitivity to metal ions such as chromium or nickel, or -- very seldom -- in careinogenic responses.

Whereas these problems seem to have been solved satis­faetorily today with the use of titanium and its alloys, the polymers most utilized in joint replaeement -- namely ultra­-high molecular weight (UHMW) pOlyethylene and PMMA bone eement -- still suffer ei ther from mechanical wea.kness, con­dueive to abrasive wear or fatigue, or from unfavourable reactions to wear particles or released eonstituents, with the inevitable result of loosening and failure of the skelet­al implant.

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2. METAL IMPLANTS

Studies -- sometimes specific ones -- and in the first place experience have confirmed that "surgical grade" metallic materials are confined to: 316 L stainless steel, Co-Cr-Mo alloys, and titanium or Ti-6Al-4V alloy. Some special inert metals are sometimes used, in particular tantalum andniobium.

Co-Ni-Cr alloys nresent better mechanical properties than Co-Cr-Mo ones, but their high Ni content is conducive to wear problems.

All the mechanical properties described play a key role both for implant design and for obtaining products capable of adaptation to the nature and the characteristics of bone tissue.

A comparison, for example, of the elastic moduli of a variety of implant materials will provide values of less than 2 x 10~ MPa for high-density pOlyethylene 4and traditional bone cement (PMMA), of between 2 - 4 x 10 MPa for cortical 4 bone, of around 18 x 104 WPa for 316 L steel, and of 22 x 10 MPa for Co-er-Mo alloys.

Fatigue resistance ranges from 200 MPa for 316 L steel to about 900 MPa for titanium alloys (Ti-6Al-4V), and the elastic limit from 180 MPa for 316 L steel to 750 MPa for Ti-6Al-4V.

In orthopaedy, supermetals were recently introduced which -- though similar in composition to those previously used -- can guarantee improved fatigue resistance thanks to special forging methods consisting in high-pressure compress­ion of the metals and melting of fine metal powders (hot isostatic pressing, or HIP).

Among other things, these nrocedures tend to decrease grain dimensions, thus minimizing the defects of a material.

3. POLYMER IMPLANTS

Of all polymers, the best known and the most utilized for both bone and joint imnlants is ultra-high-molecular weight poly­ethylene (UHMW), which, among other things, is also the most reliable for application in joint implants in combination with other met als or ceramics.

This material, classified as "thermoplastic", is capable of altering its mechanical pronerties with varying tempera­tures.

It may surely be true that today the performance of UHMW regarding degradation in vive and ageing is good in most cases, but problems are certainly still posed by the high radiation dose that can change the physical behaviour of the

4

material. Cold flow and surface wear are factors to be taken into

account, because HDP particles may give rise to biological reactions at the bone/prosthesis and bone/cement interfaces.

An important feature is a correct design of the HDP­-based implant in order to avoid any malfunction that might derive from the not always excellent characteristics of poly­mers in terms of inertia, especially when they are subjected to tensile and/or shearing stresses.

It is possible to have recourse to a metal as a rein­forcer to reduce, in particular, cold flow of the UHMW. But in this event there will be a decrease in the HDP thickness which will favour cold flow and wear. An interesting appli­cation is reinforcement with fibres for improved mechanical properties (resistance to bending and to crack formation). But this solution has been proved to be unreliable in respect of wear and fracture propagation, with the additional risk of finding carbon-fibre residues in surrounding tissues.

A number of other polymers or copolymers, such as poly­tetraphthalate and polyacetyl, have been abandoned because of a high wear rate and negative biological reactions towards wear particles. So HDP is the only one left, particularly for application in prosthetic joints, but it needs further im­provement.

To sum up, polymers can be usefully employed in joints in association with other materials. But their low resistance to cold flow and their proneness to damage as a result of breakage from bending indicate the necessity of imnrovement.

We are therefore very much in need of new pOlymers presenting better mechanical properties and capable of better biological responses.

4. CERAMIC IMPLANTS

It is a well-kno~n fact that ceramics constitute a wide class of sintered materials characterized by intimate interconnect­ion of crystal grains joined to each other and made up of medium- to highly-oxidized metallic elements bonded covalent­ly to medium- to highly-reduced non-metallic elements.

Ceramics exhibit a high modulus of elasticity combined with low ductility, very hi~h resistance to abrasion, high compression strength, low bending strength -- all qualities imparted by the covalent nature of the chemical bonds in­volved.

However, they include materials -- such as the ZrO ~- of relatively high bending strength and elastic modulus. T~ey are

5

often chemically inert and can therefore withstand chemical aggression.

The commonest of biostable cera~ics is high-nurity A1203' which has a Young's modulus of about 400 GPa and a bending strength of about 700 ~Pa.

The most distinctive feature of ceramics is their adapt­ability to tissue due to their surface wettability resulting from their high surface tension (energy). This means that an aqueous solution -- e.g. nhysiolop;ical fluids -- wetting a ceramic will always subtend a small contact angle in its respect (45 0 for A1203). This factor exerts a very favourable influence both on cell adhesion to the surface of an implant­ed prosthesis and on the lubrication of joints, certainly a much more favourable influence than with contact angles of steels (72-87°) and pOlymers (80 0 ). As far as joints are concerned, this property ensures lower friction and wear (CHAO, E.Y.S., Orthopaedic Ceramic Implants, Vol. 5-1, 1978, p. 1).

Again in relation to joints, of importance is the ex­tent of friction developing between two materials in sliding motion against one another under a load, for example in the case of a ball-acetabulum joint. Clinical experience has proved that with this kind of coupling ceramic-ceramic com­binations ensure better wear resistance than do metal-metal or metal-polyethylene ones, and a still better resistance if comnared with ceramic-polyethylene combinations. This is due to the small size of ceramic grains, which permits to obtain finished products of extra-smooth surface.

Among the advantages of ceramics, in addition to high resistance to corrosion, is their exceptional compatibility with tissue.

Bioceramics, though, present a number of drawbacks which so far have not always allowed them to be used in orthopaedics, particularly in environments liable to high shearing, torsion­al, or tensile stress.

One reason is that their nature of typically covalent compounds makes them brittle. Another reason is their low resistance to fracture propagation, because they have -­compared with bone -- a high elastic modulus.

At any rate, the best results are obtained by the plasma­-spraying technique. It is beyond question, however, that remarkable chemical inertia and great hardness, and conse­quently low wear under chemical attack by physiological fluids, warrant the use of ceramics as corrosion-resistant coatings on the surface of metal prosthetic implants.

Besides hydroxyapatite (traditionally studied), the bio­ceramies sector includes bioactive glasses. These are glassy

6

compositions (generally silica-based) capable of interesting, favourable interactions with bone tissue. They are ordinarily brittler than ceramies and more easily prone to chemical attack, but they can alter -- one way or another -- their own surface characteristics in time to become acceptable by bony tissue newly developed around them.

Bioactive glasses are now studied prevalently as sub­stances potentially suitable for coating other support mate­rials (metals or ceramies) to make them bioactive.

5. HOW CERAMICS INTERACT WITH BONE TISSUE

In order that a prosthesis may be able to restore the biofun­ctionality of the joint or the organ portion replaced, it is necessary that the surface interactions between the material constituting the prosthesis and the host tissue should give rise to a host response (by the living system towards the material) to enable biointegration of the material in question with surrounding tissue.

Bioceramics are now in great demand for biomedical appli­cation precisely for their very good biological performance (biocompatibility). Barring some cases of utilization either for their ability to be absorbed or with the function of fillers (granules, aggregates, etc), ceramics are gene rally used to carry out certain processes of bioadhesion to host tissue. The nature of such bioadhesion will vary depending on the type of ceramic adopted.

A classification into two classes is possible: inert ceramies, producing a minimal interfacial response not result­ing in tissue bonding or rejection, and bioactive ceramics, producing an interfacial response resulting in tissue bonding.

The behaviour of bioinert ceramics at the interface of an implant is of the kind shown in Figures I and 2', corres­ponding to Al203 and Zr02 respectively, where between the surface of the ceramic and that of the neoformation bone in­grown over the ceramic there is a gap of minimum thickness equal to 30 pm. The gap is gene rally filled with fibrous tissue. Contiguity is thereby established between the two phases, namely bone and ceramic.

Of all bioactive ceramics the most important is hydroxy­apatite, which can determine a bioattachment (i.e. fastening of cells and/or tissue to the surface of a material, includ­ing mechanical interlocking) of the type shown in Figure 3, ensuring substantial continuity between bone tissue- and the implanted material. Not all bioactive materials, though, behave in this way. A feature common to all bioactive cera-

7

mies is the eapaeity for biostimulation to promote osteoin­duetion, that is, the proeess whereby osteogenesis is indueed, with formation and development of bone. Another proeess in­volved is that of osteoconduction, whereby bone is directed so as to eonform to the surfaee of a material.

Bioaetive glasses are eounted among the bioaetive sub­stanees for their eapacity for biostimulation, osteoconduct­ion, and osteoinduetion. Figures 4 and 5 show examples of ingrown and well-developed bone, with osteons and vessels close to the bioactive glass -- an oecurrence normally not taking place with inert ceramies. Bioactive glasses are gene­rally less capable of osteointegration (the combination of new bone with a bioactive material) than is hydroxyapatite, and their behaviour is approximately intermediate between that of totally inert substances and that of hydroxyapatite. In bioactive glasses there can be noted astate of biodegradation all around them to a depth of between 50 and 300 pm depending on the type of bioactive glass.

As a rule, the nature of the bond established with tissue depends on the adsorption of protein moleeules onto the sur­face of an implanted material. In the case of hydroxyapatite there also occurs -- in places -- direct deposition of mine­ral hydroxyapatite performed by cells.

Biostability (the ability of a material to resist changes in a biologieal environment) is maximum with bioinert ceram­ies, optimal with hydroxyapatite, and sufficient with bioact­ive glasses.

6. NEW HORIZONS'OF RESEARCH'AND TECHNOLOGY

Until recently the implantation -- by internal surgery -- of prosthetie substitutes was considered from the point of view of the interaction by an implanted material towards tissue but never specifically the other way round, i.e. by tissue towards a material.

There has always been a tendency to transform the surface of implantable materials into an inert one, or anyway to sub­ject it to passivation in order to avoid visible phenomena of tissue corrosion or inflammation. Our conviction, particular­lyon account of the availability of bioactive materials, is that the time has now come to investigate the biochemical mechanisms that tissues apply to cause the extraneous object to adapt to their own requirements.

To identify such interaction mechanisms the best thing is to carry out experiments both in vitro and in vivo on the basis of preliminarily defined and supposedly reliable models.

8

In vitro tests make wide use of cell proliferation techniques adopting various cell lineages. An interesting proeedure eurrently utilizes cell lineages coming from human bone marrow cells, useful for their cornnatibility with blood.

Studies on rnicroporous hydroxyapatite have shown tissue to react in the presence of ultra-fine porosity by allowing sorne organie and inorganic ion species -- and probably en­zyme ones too -- to filter through the norosity to a depth of about a hundred microns (KRAJEWSKI, A., RAVAGLIOLI, A., MONGIORGI, R., and MORONI, A., r!!ineralization and Calcium Fixation within a Porous Apatitic Ceramic Material after Implantation in the Femur of Rabbits, in J. Biomed. ~at. Res., 1988, gg, 445-57).

Hydroxyapatite is a substance well accepted by an orga­nism, even though its behaviour is still to a certain extent open to question. While in fact compact hydroxyapatitic cera­mies appear finally to become incorpoFated and enclosed with­in bone tissue, porous apatites -- especially low-density (65%) powders -- are either partially or completely metabo­lized by organie tissues, which break them down into their constituents which may in turn be utilized for bone recons­truction.

This is the reason why porous grains of this -- or of triealcium phosphate, more resorbable to solve problems of loeal bone reconstruction in plants.

substance are employed dental im-

There are, though, limits to powder dimensions, beeause over-fine powders (5-10 pm) of all materials (including hydroxyapatite) have been observed to induce toxie effects causing inhibition of mitosis or even cell death.

This is because micrograins become pasted on proteins of the cel1 membrane as a result both of surface-tension pheno­mena and of the radius of curvature, so that the rnicrograins finally cannot come unstuck from the membrane. The inference is that the toxic phenomena connected with ultra-fine powder dimensions have no relation at all with the chemical toxicity of the composite constituting a given powder, even though it may often be possible to find eases of chemical toxicity added to the toxicity conneeted with powder miero-dimensions.

Henee the need for aceurate dimensional screening of hydroxyapatite powders (or any other type of powder) and for control over the formation -- if any -- of powder debris in joint implant deviees.

The procedure with bioactive glasses consists in veri­fying how ion-yielding velocity is connected with a capacity for optimum interaction, which represents the ideal condition for glass-tissue adhesion. Neither a too fast nor a too slow

9

ion release can in fact bring ab out optimum adhesion. There is, however, a form of ion-release regulation (seemingly associated with the mechanism of glass recrystallization) which partially alters the composition of the remaining glassy portion and thereby also modifies the chemico-physical proper­ties of it in terms of solubilization.

A bioactive glass interacts with surrounding tissue by developing a colloidal-silica layer which becomes negatively charged and grows richer in calcium phosphates. Electrostatic charging and the ensuing interaction are connected with the nature and the concentration of the ions released from the ~lass and scattering within the col10idal-silica 1ayer, thus acquiring an ability to estab1ish bonds of varying stability with protein substances (collagen and ce1l membranes inc1u­ded). Such interaction is precise1y the cause of the extens­ive proliferation observed in osteoblast cu1tures in glasses. But in this sector there is still 1itt1e knowledge about the mechanics of this kind of interaction at the interface.

An increasing number of studies have pointed to some new ceramic substances as potential candidates for the role of biom~teria1s.

In this respect it must be remembered that the interact­ion between protein macromo1ecules and the surface of a bio­material is re1ated to the e1ectronic state of the surface, which in turn is inf1uenced by the nature of the substance constituting the surface in question and by the impurities that occur there (whether by design or by accident).

At the roots of the adaptabi1ity or non-adaptability of each material candidate to biomedica1 use are the bio10gica1 effects that correspond to the ability of the material in question either to bind protein substances to its own surface or to cause them to be denatured or repu1sed.

Substances never considered before such as diamond, graphite, nitrides, carbides, borides, si1icides, and certain oxides have furnished encouraging resu1ts when subjected to analytic tests aimed to assess their suitabi1ity for the vascular sector. Carbon and silicon nitrides do not apnear to assure good osteointegration, and they have been monitored to give rise to inf1ammatory reactions.

A good bio1ogica1 and cyto1ogica1 compatibi1ity is dis­p1ayed by Zr02 stabi1ized with both Y203 and CaO, though in femoral heads the Zr02 (similarly to Si3N4 ) proves less abrasion resistant than A1 203 • The latter therefore appears to be the one with the lowest wear and friction coefficient. Among the blames imputed to Zr02 is the fact that it is avail­able on the market in a form containing impurities which make it radioactive. To be sure, in most cases this is untrue,

10

while in the few rema1n1ng cases radioactivity is extremely low (comparable to that emitted by the human body).

Some special ceramics are utilized also for thin coat­ing of metal substrates, particularly of titanium or titanium alloys. Particular attention was placed on comparing the performances of TiN and TaN composites. These composites, originally designed to coat metal parts in friction areas (mainly at joints), can find additional applications thanks to their good compatibility with blood. Histological tests have in fact confirmed their applicability on surfaces as semiconductive and dielectric layers for an improved resist­ance to corrosion and a better biological compatibility with proteins, especially blood proteins.

Often, specific dopants are added, for example F and Sr to hydroxyapatite or bioactive glass. This is a procedure by now universally applied, so that one could even talk of com­posite materials. The additives are made up of substances chosen among those already known as stimulators of bone in­growth but at the same time also susceptible of interesting applications, for example as coatings on other materials.

The types of equipment currently adopted both by tradi­tional techniques and by techniques of analytical evaluation (e.g. Raman, IR, EPR, Auger) are almost exclusively those enabling to identify surface transformations.

7. BASIC QUESTIONS ON HOW TO ACQUIRE FURTHER KNOWLEDGE

Oan the introduction of a substance (e.g. a glass ceramic) result in activation and/or snecialization of specified cells? And what role do such cells play in the synthesis of the various enzymes that may attack or mineralize the surrounding surface of an implanted material?

Given that our studies on bioactive glasses have engen­dered the doubt that granulated bioactive glasses may give rise to osteocyte activity, is osteocyte intervention and activation triggered by a chemical or a physiological environ­ment? For example, a material may release ions capable of altering the chemical environment in terms of pH', pNa, pK, pOa, etc with the result that cell receptors may be induced to intervene and/or activate cells: in this case, can the material -- with its presence, its surface roughness, its porosity, its ability to activate electrostatic charges on the surface, its wettability towards proteins -- interfere with and stimulate tissue, causing it in its turn to induce activa­tion (by signals to be identified) of specific cells?

Definitely, what is really needed to carry out a study of

11

the interactions between cells and a material in order to identify the transformations taking place on the surface of the material?

7. BASIC QUESTIONS ON THE CRITERIA FOR INDUSTRIAL PRODUCTION

Given the great complexity of the points at issue and the need for interdisciplinary research, it is easy to understand the reasons for the concern -- and sometimes even perplexity -- of all of uso We are in fact all aware of the difficulty of succeeding in making our "realizations" practicable and so­cially adequate.

We must therefore find an answer to the following funda­mental questions concerning production of the materials need­ed: 1. What should be the orientation of manufacturers of ceramic

objects for biomedical use? 2. Which of the existing industrial structures are the most

competent and the best suited for this kind of production? It is in fact beyond quest ion that manufacturers of tradi­tional ceramies and ceramic tiles find great difficulties in solving the specific problems associated with the pro­duction of advanced ceramies, e.g. those used as biomate­rials.

3. What place should Italian, French, German, etc products occupy in the Euronean economic context?

4. What strategy should be adopted to create staffs with tech­nical skills suited to this kind of production?

5. In what direction should research move to provide an answer to the preceding questions?

6. Talking of a "biomaterials science", what criteria should be a~plied for best allocation of the different scientific interests within this new science?

Cerarnic products, which are destined to nlay an increas­ingly important part among biomaterials , are faced wi th th·a following main problems: a) reactions at the interface; b) influence on tissues, alsoin the long te'rm; c) technical applicabili ty of certain kinds of information; d) biofunctionality of implants.

9. NORMS AND TERlHNOLOGIES

In perspective, the fundamental thing will be to nrovide the authorities in charge of laying down regulations with all the experimental data available on biochemical interactions, ion transnort and analysis of surfaces, compactness and remodell-

12

ing, mechanical strength, wear, degradation. In the near future it will be essential to start evalua­

ting the approach to be adopted to standardize the methods of clinical, chemical and mechanical, and cytologico-physiologi­cal investigation.

One further objective should be the drafting of a docu­ment proposing new technical norms, both national and inter­national, for standardized control over the quality and the performance of each material or product. Such specific norms should at a later stage be officially codified into a nroper regulatory system.

Another important factor will be an increasingly accura­te standardization of the terminologies adopted by the differ­ent sectors. It may in fact happen that single terms are applied to partially or entirely different concepts depending on the scientific sectors in which the terms in question are used.

At the present moment, nationwide regulations and norms (such as nermany's DIN or USA's ASTM) continue to bring toge­ther scientific experts and federal officials for step by step discussion and selection of the fields of application of biomaterials (e.g.: "surgical implants", "dental implants" or "dental materials", and so on). Other countries, like Austria, Italy, etc, prefer to either simply copy existing reeulations (mainly DIN ones) or draft their own regulations.

There currently exist international bOdies, such as the ISO, engaged in trying to establish regulations applicable to all countries (adhering to the ISO). Obviously, the more are the countries that try to reach an agreement, the more com­plicated and time-consuming is the procedure.

Lanp,uage differences come into play too, for it is ob­vious that different terrninologies make things more compli­cated.

Consensus Conferences, however, as organized in the USA or by David Williams in Europe (Chester lQ86), have been of much help for adontion of uniform definitions ("biocom-patibi­lity", "bioreactivity", etc) in compliance with nationwide or international regulations. The same applies to technical norms in the matter of quality and composition of materials, testing procedures and quality control, etc.

10. MATERIALS FOR BIOMEDICAL APPLICATION: COMMERCIAL ASPECTS

Aseries of technological breakthroughs has given impetus worldwide to research on a number of new biomaterials increas­ingly used as substitutes for fundamental organs.

13

The emergent biomaterials sector is therefore becoming strategically important in Italy too, both for the market prospects it opens Up and for the benefits it can bring na­tionwide through synergy between researchers in the fields of industrial technology and medicine. With 1992 approaching, Italian, French, German, etc industries must be ready to occupy a strong position in the European context and to this purpose they must put competitive biomedical products on the European market.

There is in fact an awfully high potentiality for pro­ducts for biomedical application, and forecasts indicate a yearly growth of 20 to 25% (according to FIND/SVP sources, USA, 1986).

The USA is the place of origin and development of bioma­terials. The Americans, currently the chief producers and exporters, are the undisputed leaders in this market area with a 47% share, followed by Japan with 19% and Germany with 13% (according to clinical sources, 1987).

The more interesting fields of application are, in terms of market potentialities: - The cardiovascular segment, with a sales turnover on the USA

market of USD.550,000,000 in 1983 and a forecast of USD. 1,140,000,000 for 1991.

- The orthopaedic segment, with USD.336,000,000 in 1983 and an expected 570,000,000 for lQ91.

- Other segments (urinary components, eye lenses, etc), with a market potentiality of USD.l,OOO,OOO,OOO (foreseen for 1991).

Another important market segment, emerged recently, is that of products for bone/joints application.

Japan, which generally imports its requirements from the USA, has only recently started a number of R & D projects on bioactive materials for artificial organs and biochips. Japan is nonetheless highly competitive -- and a significant export­er -- in certain sectors of medical equipment.

Europe, on the contrary, is trailing very much behind as a consequence of scant R & D investment. An exception is Germany, at the forefront of research into new materials capable of dete~ining in situ release of drugs or antibio­tics.

A number of interesting research projects have recently been started in the framework of the Eureka programme of investigation of the methodologies for controlling the quali­ty of biomaterials.

The fact is that most of the European market is supplied by U.S. companies, which are always very clever in locating and occupying new markets (such as those of Portugal or

14

Figure 1 - Overall view of an A1203 ceramic implant (a) and a magnification

at the interface (b).

2nvn

Figure 2 - Overall view of a Zr02 ceramic implant (a) and a magnification

at the interface (b).

15

Figure 3 - Overall view of a hydroxyapatite ceramic implant (a) and a magni­

fication at the interface (b).

Figure 4 - Overall view of a Hench I s type 4585 bioglass implant (a) and a

magnification at the interface (b).

16

Spain. Italy gene rally imports the products it needs, especially

from the USA. There are, however, a few exceptions such as heart valves, oxygenators, and dialyzers.

In Italy, though, the exploitation of some existing technological potentialities and the targeting of research on well-defined market areas can result in development of high-technology products commercially viable not just in Italy but also abroad.

The statistics concerning the Italian market are not very recent. They date back to 1983 and indicate a market value assessed at a thousand billion Lire, which -- calculating a 20% growth -- means a foreseeable market of more than two thousand billion for 19Q1.

Current domestic output is negligible, despite a not indifferent technological knowhow at a level of basic re­search. The effects, though, are not appreciable where pro­duction is concerned.

Of some interest is the sector of artificial organs. In conclusion, there is no doubt that the USA lead the

field. But I would like to emphasize the higher potential of the Europeans in terms of inventiveness and engineering skilI.

We Europeans simply must encourage our industry to be better, and sooner or later the USA market (and FDA) will have to accept the superiority of our implant systems.

An example: dental implants such as those produced by the Branemark group (Sweden) or the IMZ (Germany) are still the trend-setters, and the Americans simply copy them or allow them for clinical use. By the way.

Figure 5 - Overall view of an AKRA 15 biological glass implant (a) and a

magnification at the interface (b).

17

15 YEARS EXPERIENCES WITH ALUMINA-CERAKIC TOTAL HIP-JOINT ENDOPROSTHESES. A CLINICAL, HISTOLOGICAL AND TRIBOLOGICAL

ANALYSIS

H. PLENK Jr.', M. BÖHLER' , A. WALTER1 , K. KNAHR4 , and M. SALZER1

1) Bone & Biomater.Res.Lab.,Histolog.-Embryolog.lnstitute, Univ.of Vienna, Schwarz spaniers trasse 17, A-1090 VIENNA.

2) Orthopaedic Hospital Vienna-Gersthof, AUSTRIA. 3) Lab.Biomechan.exp.Orthop.,Orthop.Univ.Clinic, Munich,FRG. 4) Orthopaedic Hospital Vienna-Speising, AUSTRIA.

ABSTRACT

The clinical and tribological performance of 67 ceramic sockets was evaluated, implanted 1976 to 1979 without bone cement in alumina-on-alumlna metal compound total hip­joint replacements. 8 sockets showed radiological signs of loosening, resulting in a survival probability of 83.80f0(±5.4%) at a mean survival time of 144.7 months. Around all stable sockets a 1-2 mm wlde radiolucent seam developed and remained constant In time. The surrounding tissues of stable sockets, removed due to stem failures, showed a replica of the ceramic surface grooves by a pseudosynovia or occasional fibrocartilage, but no bone ingrowth. Small amounts of ceramic debris corresponded weil to low wear rates/year (socket:O.95 pm, head: 1.7 pm), calculated from tribological analysis of one alumina-pairing after 133 months of service. Early loosening appeared fram 8 months on as concentric widening of the radiolucent seam to more than 10 mm, while late loosening was indicated from 25 or 80 months on by tilting into valgus position and migration of the socket. Enormous amounts of ceramic wear particles stored in macrophages were found in surrounding tissues, accompanied by chronic inflammation and bone resorption only in early loosening. Tribological analysis of one loose alumina-pairing showed design-related damage, inadequate ceramic density and thus increased abrasion of the components (socket:27.2 pm, head:40.9 pm/year). In conclusion, properly manufactured alumina-ceramic can meet the expectations for long-Iasting articulating surfaces, provided correct position and stable condition of the implant. Cementfree ceramic socket stabilization is only achieved by a connective tissue interlayer to bone which is vulnerable by accumulation of ceramic wear particles.

INTRODUCTION

Dense, pOlycrystalline high purity alumina-ceramic was in the early '70ies introduced as

a biomaterial for artificial joint replacement because of its technically proven wear

resistance and chemical inertness. Similar to other groups, Salzer et al.(1) reported on

18

layour.,'e experIencea with alumina-ceremic joint encloprostheses, end we were among

the ftrat who coulcl demon.trate histomorphologlcally (2) that undi.turbed bon. healing

end Ingrowth lad to a firm and dwable mechanlcal interlocking with cerarnlc surface

grooyea, as put forward by HuJbert et al.(3).

Meanwhile, time has shown that some of the alumina-ceramic implant systems had only

limited application for tumor patienta (2), whi\e others failed, e.g. for the surface

replacement of the femoral head (4). Howeyer, the acetabular replacement of all these

ceramic arthroplastles showed promiaing results In experimental and clinical trials (see

Figs.1, a-b), and was then wed in combination with ceramic ball-heads on metallic

stems as a total ceramic replacement for osteoarthrltlc hip-joints (4). In this paper, a

retreat to metal- or ceremic-po\yethylene articulation was announced, but these

combinations proYed not totally satislactory elther. Recently the alumina-on-alumina

pairing has again been promoted because of the remarkably low wear (5). In the study

pre.ented here, our cllnical and histomorphological experiences with the alumina-ceramic

socket endoprosthese. will be updated to nearly 15 years after Implantation.

Figure 1 ,a-b: Bone ingrowth(arrows) into surface grooves of a ceramic socket in the dog(a) and in a human tumor patient(b). (Microradiographs from corresponding ground

section., a) 6 months, b) 12 months after implantation.)

19

MATBIAL MD IETHODS

AItogether 81 socket endoprostheaea (hemlapherlcal wIth 3 excentrlc pedlclea end

outaide clrcular grooves, partly wIIh a centr .. hole), made of poIycryat8lllne high denalty

and high purlty a1umlnum oxide-ceramlc (RosenthaI Technik AG, Lauf/P.,FRG) were

Implanted between 1978 and 1979 without using bone cement. combined with ceramlc

ball-h .. cIa (32 mm diameter) on cobaltbued aUoy-stema (n-19 cernented, n-48

uncemented) for total hip-joint replacement in 65 patients (13 males, 52 females, mean

age 83.8 y .. ra).

CHnical (according to Harris'hip-score) and radlological foUow-upa were performed every

3 months postoperatively, and then at yearly recalls. Baseel on the radlologieal flndlngs,

a statistica! survival analysis was caJculated (8).

At re-operation, the implanta were removeel and the periprosthetic tissue. processed for

microscopicaJ evaluation under transmitted and pOlarized Hght (1).

One clinicaUy and radlologically stable and one looaened socket each with the

corresponding ball-heads underwent tribological evaluation, consisting of electronic

sphericity measurementa, SEM-surfaee analysis, and ceramie denalty estimetlon.

RESULTS

Only radiological observations and flndinga made on retrieved implanta and perlimplant

tissue. are reported in the present study. Radiological evaluation criteria were the

radiolueent aeam between socket and bone, selerosis of the acetabular roof, and a

change in the socket position.

In 9 casea, as reported previously (4), fracture of the ceramic ball-head occured 2 to

64 months after implantation and eaused the removal of the clinieally and radlologically

stable sockets. Therefore, these were not eonsidered a fallure of the socket ,in the

present analysis, and fIndIngs from histologieal and tribologieal evaluations wiU not be

reported here.

Up to an observation period of 1 68 months, 59 of the ceramic sockets revealed the

pieture of a stable sockel i.e. In all cases a 1-2 mm radlolucent seam had developed

along the socket-bone interface which did not change in course of time, and the position

of the socketa had not changed too. In 18 cases an increase in the sclerosls of the

acetabular roof was observed between 24 and 48 months after implantation and

underwent no changes thereafter. In 3 cases such a stable socket was removed

because of stem loosening. Histologieal evaluation of the tissue underlying the socket

showed an about 1 mm thiek fibrovascular layer, covered by pseudosynovlal ceUs or In

some areas by fibroeartilage which formed an exact replica of the .urfaee grooves of

the eeramlc socket. Within the pseudosynovia and around vessels groups of

20

macrophages, containing minor amounts of fine, blretringent ceramic particles, were

found (Fig.2,a-b). The fibro-vascular layer was connected to living bone trabeculae,

undergoing active remodelling.

Loosening of the socket became radiologically apparent in 2 cases trom 8 months on

as a concentric widening of the radio-Iucent seam to about 10 mm and more (Fig.3,a­

b), and led to revision surgery after 11 and 18 months post impl., respectively.

Histology revealed enormous amounts of ceramic wear particles, mainly stored in

macrophages, all over the thickened fibrous tissue interlayer. In addition, Iymphocyte and

plasmacell-infiltrates could t.e seen, and the adjacent bone showed osteoclastic

resorption.

In 3 other cases, this loosening process or a beginning change in the position of the

socket were noted at 25 to 30 months after implantation. This second mode of fallure,

amigration and tiHing of the socket into valgus position, was particularly noticeable in

another 3 cases at 80 to 86 months after implantation. A widening of the radiolucent

araa was only visible at the caudal pole of the socket (Fig.4,a-b). Only in two of these

cases the sockets were removed so far, and histology showed agaln excessive

accumulation of ceramic wear particles, but no signs of chronic inflammation as

described above.

Figure 2,a-b: Perivascular macrophages, containing amaI biretringent ceramic particles. (Pseudosynovia around a atable socket, 21 months after Implantation.Toluidin. blue,

b)polarized light, mag. 1 80x.)

21

a b

8mo

Figure 3,a-b: Schematic drawings and radiographs from aceramie socket which showed widening of the radiolucent periimplant seam at 8 months post Impl.(b)

The statistical survival analysis of the ceramic socket endoprostheses which were

radiologically considered either stable or instable (-Ioosened) at a certain point of time

after implantation which ia the actuallast follow-up, is shown in Fig.5. After 168 months

(mean survival time 144.7 months) and altogether 8 radiologie al loosenings, the

expectable success rate for the remaining stable sockets is at 83.8%(5.4%).

The tribo!ogica! anaIysis revealed in the one stable alumina-on-alumina pairing a maximal

sphericity difference of 10.5 pm for the socket (Fig.6 a), and 19 pm for the head, after

1 33 months of service. From these values and the implantation time a wear rate /year

can be calculated (see Table 1).

In the socket (as weil as on the ball-head) the zones of abrasion were even visible by

eye inspection as a loss of the shiny smooth surface finish mainly in areas a10ng the rim

of the socket. SEM inspection shows the high proportion of bearing surfaces in a smooth

zone, and in contrast the rugged surface In the abrasion zone (Ag.6 b).

In the loosened ce ramie components the sphericity loss was much higher (at least 160

pm for the socket, and 200 pm for the head), resulting in increased wear rates (see

22

Table 1). In addition, the socket showed a fracture damage around the central hole (see

Fig.6 cl, and estimation of the ceramic denaity reveaJed a difference of 5.2% to the

theoretical density of dense, high purity aluminum oxide-ceramlc.

a b

3mo SOmo

Flgure 4,a-b: Schematic drawings and radiographs from a ceramic socket whlch showed crenial migration, tllting into valgus posi-tion, and a wide radiolucent area at the caudal

pole at 80 months post impl.(b)

%

S u r y

i y

a I

100

~-----.Q~B--------<[)o.-" 80

60

40

20 l_. __ n~==_~_;._-_S=O=C~k_e_~_s __ ~ __ ·_··_· _n_=_59 SOcket:] o ~------------------------------------------~

o 12 24 36 48 60 72 84 96 108 120 132 144 156 168 180

Months

Figure 5: Survival probability of alumina-ceramic socket endoprosthese. up to 168 montha after implantation.

23

Figure 6,a-c: Sphericity m .. aurementa in 2 horizontal planes (a) and SEM-picturea (b) of a amooth (above) and a rough surface ar.. (below) In the stable ceramic socket. SEM-pietures of the fraeture damage around the centraJ hole of the loosened socket (c).

Table 1: Tribologic:aJ data on 2 alumina-on-aJumina pairings.

HIP-ARTHROPLASTY Stable Loosened

Implantation time 133 mo 135 110

Wear rate pm/year

SOCKET 1.7 27.2 BALL-HEAD 0.95 40.9

Difference , theoret.Density • 5.2 2.2

• Range of 1976 ceramies 1.2% to 3.2%

24

DlSCUSSION AND CONCLUSIONS

TM design of thl. a1umlna-cerarnlc socket endopro.the.l. aeemed to prove

experimentally end cllnlcally .ucceaaful (8): The implantation i. technlcally .Imple, a

stable primary anchorage In correct po.ltion could be achieved In nearJy a11 ca .. s,

secondary stabilizatlon by bone Ingrowth into the surface grooves was observed, at least

in the main load-bearlng erea (Fig.1 , a-b), end nelther radIologicaI nor cllnical finding.

Indicated eny considerable wear of the ceramic.

The present and previous long-term evaluations (4,9), however, showed that only with

respect to wear reslstance dense polycrystaUJne a1umina-ceramic cen meet the

expectations, provided 1) proper manufacturing end design, 2) Implantation in correct

position, end 3) stable conditlon of the implanl(s). As with alumina-ceramic

endoprostheses from other manufacturers, in some c .... excesaive weer and abrasion

was encountered, explainable by inadequate grain size and resuRing lower density of the

material. Also the design of the implant can be fauHy, as In our case the central hole

which should facilitate insertion, but created posaibly too high contact tension end led

to fracture. Fractured ceramic partlcles will inevitably increase the wear of artlculating

surfaces. And so does en Incorrect implant po.ltion, e.g. In valgus position.

Finally, the stable anchorage of the artIculating implents is of utmost importance. While

in this study end other reports on ceramic hip-joint socket implents (5) the cement-free

enchorage can be regarded .. cllnically succeaaful, the radiologie al and the hiatological

observations reported here demonatrate that practlcally no direct bone contacts hold

these implants in place, but only this thin fibrovasculer interlayer with fibrocartilage­

metaplasia in some areas. In addition, revision surgery showed that osteophytic bone

formations at the rim of the socket end in the newly formed joint cap.ule helped to

stabilize the implent. In contrast, the excessive wear accumulating along the implent­

bone interfaces contributed to the loosening process, making this soft tissue interlayer

only thicker, but even less suitable for bearing loads. In the cases with early loosening

of the socket, the resuRing foreign body reaction seemed accompanied by a chronIe

inflammatory response, leading to increased bone resorption and thus to a more

agressive widening of the periimplent space. In the cases with late loosening, the sole

accumulation of excessive wear particles could Mve made the interface more susceptible

to the loading forces, resuRing in migration of the ceramic socket and the typical tiRing

into valgus position.

To avoid this apparent weak point in enchorage, but still making advantage of the

theoretically excellent wear resistance of alumina-ceramic, a titanium-alloy socket of

similar design (Intraplant AG, Zug, Switzerland), but with aceramie inlay (Feldmühle,

Plochingen,FRG) instead of the conventlonal polyethylene one, will soon be introduced.

25

ACKNOWLEDGEMENT

The support of this study by COMESA GmbH., Vlenna,Austrla, ja gratefully acknow\edged.

REFERENCES

1. Salzer, M., Locke, H., Zweyaüller, K., Plenk, H.Jr., Punzet, G., Zeibig, A. and Stärk, N., Further experimental and clinical experience wi th aluminum oxide endoprostheses. J.Biomed.Mater.Res., 1976, 10, 847-856.

2. Plenk, H.Jr.,Salzer, M., Locke, H., Stärk, N., Punzet, G. and Zweymüller, K., Extracortical attachment of bioceralDic endoprostheses to long bones without bone cement. Clin.Orthop.Rel.Res., 1978, 132, 252-265.

3. Hulbert, S.F., Cooke, F.W., Klawitter, J.J., Leonard, R.B., Sauer, B.W., Moyle, D.D. and Skinner, H.B., Attachment of prostheses to the musculo-skeletal system by tissue ingrowth and mechanical interlocking. J.Biomed.Mater.Res.Symp., 1973, 4, 1-23.

4. Salzer, M., Knahr, K. and Plenk, H.Jr., Long-term clinical and histological evaluation of bioceramic total hip ende prostheses. Orthopedics, 1981, 4, 1231-1240.

5. Sedel, L., Kerboull, L., Christei, P., Meunier, A. and Witvoet, J., Alumina-on-alumina hip replacement. Results and survivorship in youngpatients. J.Bone.Jt.Surq., 1990, 72B, 658-663.

6. Kaplan, E.L. and Meier, P., Nonparametrie estimation for in-complete observations. J.Amer.Statist.Assoc., 1958, 53, 457-481.

7. Plenk, H.Jr., The microscopic evaluation of hard tissue implants. In Techniques of BiocompatibilityTestinq Vol.I, ed. D.F.Williams, CRC-Press Inc., Boca Raton,FL, 1986, pp. 35-81.

8. Plenk, H.Jr., Biocompatibility of ceramies in joint prostheses. In Biocompatibility of Orthopedic Implantsr Vol.I., ed. D.F.Williams, Boca Raton,FL, 1982, pp. 269-295.

9. Knahr, K., Böhler, M., Frank, P., Plenk, H.Jr., andSalzer, M., Survival analysis of an uncemented acetabular component in total hip replacement. Arch.Orthop.Trauma.Surq., 1987, 106, 297-300.

26

BIO-FUNCTIONAL ADAPTIVE BEHA VIOR TO CERAMIC IMPLANTS

A. Toni*, A. Sudanese*, S. Stea°, S. SquarzoniA, P.P. Montina*, A. Bueno Lozano*, F. Calista*, A. Pizzoferrato° and A. Giunti*

* Cattedra di Patologia Apparato Locomotore dell'UniversitA di Bologna ° Laboratorio di BiocompatibilitA

A Istituto di Citomorfologia Nonnale e Patologica - CNR Istituto Rizzoli - Bologna, ltaly

INTRODUCTION

The biological interaction between host bone tissue and prosthesis surface material is a debated topic among scientists. In cementless total hip arthroplasty new ceramic coating were recently proposed, aiming to improve hone ingrowth and implant stability through a positive remodelling of bone. However, the long term results of ceramic coating seems to be still undefined. Large clinical results are necessary to verify the effectiveness of ceramic coatings. Bone adaptive remodelling is influenced either by the implant surface chemistry, or its overall design, or by the initial stability achieved at surgery. We studied clinical bone remodelling following the implant of an alumina coated anatomie uncemented total hip arthroplasty we are using since 1985: with this significant experience we intend to present some aspects of the bio-functional adaptive behavior to ceramic implants.

MATERIALS AND METHODS

Since October 1985 to February 1990 we implanted 265 Anatomie Ceramic Arthroplasty [1], herein defined as AN.C.A. model (produced by Cremascoli, Italy). The AN.C.A. prosthesis is made with a cobalt-chrome alumina coated stern, a 32 mm alumina ceramic head and a screw-in socket, made by combining an outer titanium threaded ring to an alumina acetabulum, coated in the domed outer surface by a new Porous Alumina-bioglass composite (PORAL®)[2]. Thirty-eight implants were used for revision surgery ofprevious loosened prostheses and were not included in the present study. Among the remaining 227 implants,only 147 cases with ~12 months follow-up were considered. Cases with septic loosening were also discarded due to the fact that the septic inflammatory reaction impairs the bone tissue adaptive remodelling. To make useful comparisons to better understand the role of ceramic coating on the biofunctional adaptive behavior of hone we considered also 48 cases operated on with the LORD sockets [3] . The AN.C.A. and LORD sockets represent a good clinical comparative model, since both search for primary stability through a screw fixation and, besides the shape (truncated cone for LORD and spherical for AN.C.A.), the main difference is due to the presence of a hone ingrowth porosity (PORAL®) in the AN.C.A. implant to achieve secondary stability.

27

The radiographie evaluation has been perfonned with a Detailed Eye Evaluation Protocol (DEEP), previously described [1]. We classified the prosthetic stabilization referring to three levels: bony,jibrous and unstable, evaluating the incidence in the regions of interest (7 for the femur and 3 for the socket according to Gruen protocol [4]) of lucency or progressively widening val1us between prosthetic surface and bone. We had then to retrieve 2 sterns because of persistent and worsening pain at load, due to stress shielded bone spongious degeneration. These cases were then thoroughly studied with undeca1cified histology, SEM, TEM and EDAX microanalysis of the bone-alumina interface.

RESULTS

Ceramic porous coating Radiographie stability for the AN.C.A. and LORD sockets is reported in Tab. 1. The relevant higher incidence of 100sening and fibrous stabilization for the LORD sockets proves the clinical relevance of the porous ceramic layer coating the AN.C.A. implant

TABLE 1 Radiographie stability

Bony stabilization Fibrous stabilization Unstable

LORD

78% 4% 18%

AN.C.A.

89% 6% 5%

======================================== 48 147

The screw fixation is reported as not liable for socket stability: our results confinn this negative trend for the implants which search for stability only by the screw (LORD), but prove that implants achieving secondary bone ingrowth stability represent a promising solution for cementless socket fixation, as clearly shown by the survival curves comparison in fig.1 (Unstable cases being the negative results), where also our cemented sockets experience is reported. The AN.C.A. sockets behave as the cemented one at similar follow­up; while the LORD sockets survival curve starts to drop at 36 months and reaches the 38% of failure at 76 months follow-up. Hopefully the cementless and ceramic AN.C.A. implants will improve the late failure incidence of the cemented socket, largely due to cement brittleness and polyethylene wear.

Alumina-bone interface Two sterns were retrieved due to pain at load. Radiographie picture showed only marked stress shielding phenomena (severe density loss of cortical and spongy bone), with no evidence of loosening of implants. Nevertheless, forced by the strong need of some improvements c1aimed by the patients, we devised to make revision surgery. At macroscopic examination sterns were both finnly connected to the femoral cancellous bone, without any fiber tissue layer in between. The metaphyseal bone was scarcely consistent from a mechanical point of view, since it was largely defonnable just by a ligth thumb pressure. The cortical bone was thinned, with a "wet cardboard" consistency. By a osteotome the bone was easily disconnected from the ceramic surface in the lateral, anterior and posterior aspects of the proximal bone-implant interface. U sing the proper extractor and a heavy hammer the sterns were then taken out with some difficulties. In the medial rough surface bone remained attached to the implant. Examining then the diaphysis we could clearly see how the stern was totally enveloped by a thin bony layer c10sely contacting the

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III .... 0 H

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-A

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et

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29

alumina surfaee. This layer was earefully removed for histologie examination prior to insert the new stern, whieh in one ease was eemented. The sterns were aetually laying onto a spongeous bony bed, which was probably not sufficient to support the implant at heavier load eycles. The histologie analysis showed very good eontaet between alumina and bone trabeculae or the thin shell above mentioned, without any fiber tissue at the interface, as shown in fig. 2. Nevertheless, the tissue more elosely related with the eeramie surfaee showed to be lighter at ematossilin-eosin staining (see arrows in fig. 2). We then stained the slide according to Goldner method [5]: the lighter area turned out to be a less calcified tissue with respect to surrounding bone tissue (fig. 3). The lower ealcified area was regularly found from the ceramic-bone interface extending 100-150 microns into the surrounding bone. Such finding was present in sampies retrieved from both patients, either in the lateral smooth eeramic surface and in the beaded rough one. The initial hypothesis was that the bone contacting the ceramic layer was still actively remodelling, interpreting such a less calcified band just as immature osteoid tissue. Opposing this theory, a closer examination of the slides showed such layer presenting a weH established osteonic structure (fig. 4).

I I

,

Fig. 2 Paragon stain (25x). See text

This finding modified our previous interpretation, leading us to believe that bone had previously obtained a close contact with the alumina surface, reaching a complete grade of osteonic organization. Something thereafter modified the mineral content. The strong relation of such decalcifying process with the alumina surface support the suspect of a chemical interaction of aluminum with the calcium metabolism, as reponed in patient with renal insufficieney on chronic maintenance dialysis[6]. The SEM analysis showed similar sharp distinetion between the two layers (Fig. 5: see arrows for the deamarcation line), it is interesting to note as the demineralized part still present a lamellar aspect, even if no so compact as the confining normally calcified lamellar bone. Edax microanalysis in the two areas confmned the Goldner staining results referring to the lower calcium levels in the area proximal to the alumina, as shown in fig. 6a for the normal bone and 6b for the decalcified area. These analysis were rispectively performed in area Bande of fig. 5.

30

Fig. 3 Goldner stain fucsine-light green Mx). A : cobalt-chrome bead; B : alumina coating. C : weH calcified tissue. Between B and C is the whitish layer corresponding to the decalcified tissue. The good contact between the decalcified layer and the alumina is clearly shown.

31

Fig. 4 Goldner stain (40x). Bone lamina retrieved in the lateral aspect of the stern. Arrows indicate the transition line between the calcified bone (stained blue in the lower left part) and the decalcified tissue (stained red in the upper part). It is clearly evident how also in the decalcified bone the osteonic structure is preserved.

32

Figure5

33

p Ca

... 0 .. .., ... ...

Fig.6a Fig.6b

DISCUSSION

The bio-functional adaptive behavior to ceramic implant proved to be clinical usefuillimiting the failure rate of cementless screw socket: bone ingrowth secondary stabilization was considered to be the factor which avoided higher loosening rate in AN.C.A. sockets. The compatibility of such porous alumina (PORAL®) was highly statisfactory, since no clinical adverse reaction was ever detected out of 147 implants at an average 36 months follow-up. The 32 MPa shear strength of the ceramic coating proved to be sufficient for this kind of application, since no bead detachment was ever seen. Histology showed a very good contact between alumina and surrounding bone, which tends to envelope with a very thin shell the whole implant. Such shell is made by a continuous 0.5-1 mm thin layer of lamellar bone where the stem is smooth, or it can be made by bony trabeculae areas of contact, which are connected among each others by a very thin layer of bony metaplasya of bone marrow. Nevertheless some concems remain due to the presence of the decalfied bone tissue interfacing the alumina coating. Such finding was detected at 28 and 41 months follow-up and could be interpreted as chemically induced osteomalacia, due to higher aluminum contamination in the bone proximal to alumina surface. Deeper study are on going to obtain specific staining for aluminum to confmn these suggestions.

REFERENCES

1. Toni, A., Sudanese, A., Ciaroni, D., Dallari, D., Greggi, T. and Giunti, A., Anatomical ceramic anhroplasty (AN.C.A.): preliminary experience with a new cementless prosthesis. Chir. Or~ani Moy., 1990,75, 81-97.

34

2. Pizzoferrato, A., Toni, A., Sudanese, A., Ciapetti, G., Tinti, A., and Venturini, A., Multilayered bead eomposite eoating for hip prosthesis; experimental studies and preliminary elinieal results. 1. Biomed. Mal. Res., 1988,22,1181-1202.

3. Lord, G.A., Hardy, I.R. and Kummer F.I., An uneemented total hip replaeement. Experimental study and review of 300 madreporique arthroplasties. Clin. Orthop., 1979, 141,2-16.

4. Gruen, T.A., Me Neiee, G.M. and Amstutz, H.C., "Modes of faHure" of eemented stern-type femoral eomponents: a radiographie analysis of loosening. Clin. OrthQP., 1979, 141, 17-27.

5. Me Manus, I.F.A. and Mowry, R.W., In Staining methods. Histologie and histochemical, A. Hoeber International Reprint, Harper & Row Publishers inc., New lork, 1960, pp.55-64.

6. Malluche, H.H. and Faugere M.C., Stainable aluminum and not aluminum content reflects bone histology in dialyzed patients. Kidney Int., 1986 Nov, 30(5), 717-722

35

A REVIEW ON THE ASEPTIC TOTAL HIP REPLACEMENT

FAllURES

G. MONTICELLI, L. ROMANINI, O. MORESCHINI

Institute of Clinic Orthopaedic,Rome University

"La Sapienza".

In Italy alone, 28.000 total hip replacements

were performed in 1990.

On these, 60% were cemented, and 40% were ce=

mentless. it is calculated that 10% of the total

is due to revision • If we furthermore consider

the 400.000 hip prostheses implanted world-wid

the approximately 40.000 of these could be "re=

visions".

In the past 17 years the rate of implantation

as gone up 00% so it seems likely that these st~

tistics should increase. Of the materials used

for artificial hip implantation, 48% were Cr-Co,

35 steel alloys, 10% titanium alloys and 7% va=

rious materials.

This brief statistical survey demonstrates the

size of an ever growing phenomenon . In fact we

have studied the revisions in order to comprehend

the causes of failure in prosthetic hip implanta=

tion.

36

MATERIALS AND METHODS ----------------------

A clinical survey of ten patients with prosthetic

failure was performed. The pre-operative status of

these ten is divided into: 7 cases of primary arthr!

tis of the hip, 2 cases of hip fracture and 1 case of

hip dysp1asia. Seven (7) imp1ants were cemented and

3 were cementless in six women and 4 men with an ave=

rageage of 69 (min. 59 - max 81). The previous im=

plant had been performed at an average of 7 years

earlier (min. 3 - max 17).

The onset of pain before revision averaged 1.7

years (between 1 - 4).

All the clinicam, radiologicam and labo.examina=

tions favored the hypothesis of an aseptic loosening

of the hip implant. This was confirmed afterwards.

The standard TC 99 bone scan was quite useful in

indicating bone tissue turnover of the acetabu1ar or

diaphysea1 regions. At the same times a marked leuko=

cyte bone scan dismissed any doubts concerning poss!

ble asepsis of the imp1ant.

During the actual revision, biopsies were taken

at the bone - pros thesis interface and the bone-ce=

me nt interface in the medial and lateral aspects of

the acetabulum, and in 4 distinct area of the dia=

physis (medial and lateral of the upper and lower

diaphyssis).

The bipsies were studied with histological cito=

chemical techniques (using v.v. light) and bacte=

riological examination. The following interfaces were

examined:

37

- cement - cement

- vitallium - polyethilene

- vitallium - ceramic

- ceramic - polyethilene

- polyethilene - cement

- steel- cement

- vitallium - cement

- titanium - cement

- titanium - polyethilene

The extemporaneous byopsy material was included in gly=

col -methacrylate and slices at a 2 mm. thickness.

Histological stain was performed with May - Grunwald,

Giemsa and metathylene blus. Cytochemical (functional)

evaluation was studied under ultra-violet light after te=

tracycline marking.

Acid phosphatase and trap levels were measured using

the Naftolo As-Bl phosphoric acid as substrate and the

Fast Garnet GBC as coupling reagent.

The 10 mm. slices for UV light examination were air­

dried and not mounted on slides.

38

RESULTS

The hystological results remain ynvaried in both

the cemented and the cementless models of hip prosth~

ses.

In either situation, bone tissue at the cement on

prosthesis interface showed signs of growth and reab=

sorption. The deposition of new bone is proved by

the presence of osteoid tissue in close contact with

osteoblasts.

Conversely, bone reabsorption is proved by the pre=

sence of Howship lacunal full of active osteoclasts.

The bone-prosthesis interface was infiltrated with

a connective tissue rich in histiocytoid and giant

cells; clearly present because of the prosthesis on

the cement.

Certain portions of this infiltrate started to

give signs of a sinovial-like cell polarization.

In other parts there was cell necrosis. There were

never any signs of infection, even though these

areas were surrounded by reactive vasculastis with

a typically thickened tunica fibrosa and endothelium.

These arteries were often surrounded by histiocy=

toid elements. The giant-cells often looked empty

because of the chemical affinity between the glycol­

methacrylate used to include them and the cement

made of polimethyl methacrylate. As a matter of

fact the two types of cells were less common in the

cementless prostheses; even though small metal-like

39

particles were found in their cytoplasm.

In the cementless prostheses the interface con=

nective tissue of the diaphysis was more compact

and genrally more organized.

The acetabular component generally had a fibro=

cartilage interface. Joint capsule and ligamentous

hypertrophy secondary to the intense mechanical

stress supplied by the prosthesis femoral component

is quite probable.

The femoral interface was consistenly composed

of vascularized connective tissue with sporadical

calcification.

We sporadically found a sinovial-like tissue

with chromic lympocitic infiltrates and vascuolized

macrophages TSAP and TRAP were positive in the hi=

stiocytoid and giant cells but even more so in the

pre+osteoclasts and osteoclasts, making it theref~

re easy to distinguish the latter. The cytoplasmatic

distribution and entity of the enzymatic reaction

in these different types of cells could lead us to

hypothesize somme kind of osteogenetic link between

them. Even though histiocytoid cells do not directly

differentiate into osteoclasts, it is reasonable to

assume that they both derive from a primal cello

There is surely a common activating factor for ma=

crophagic cells and osteoclasts, that explains hi=

stiocytoid activation

of bone cement.

(TRAP) even in the absence

In vitro studies on monocytes indicate

that there is TRAP activity after interleu activa=

tion. Therefore this enzymatic activity is an

aspecific response to local phagocytic stimulation.

We can there-fore assume that there is a relation=

ship between macrophage activation and osteoclast

differentation.

Trap activity can thus be considered aspe=

cific for the above mentioned reasons.

Our histological studies show that the

tissue reaction around and within the prosthetic

interfaces, regardless of their debris, is aspeci=

fic and therefore composed of similar cell populations

which vapy from case to case only in their relative

proportions to each other.

CONCLUSIONS

In the cases we have observed it is our opin-

ion that the duration of the prosthesis is indepedent

of age and type of implant.

It is quite probable that factors stimula=

ting osteoclasts and macrophages play an important

role in producing bone reabsorption around the implant.

On tetraciclyne markers have demonstrated that where

there is reabsorption there is also formation of

osteoide tissue and presence of osteoblasts.

41

This process remains in equilibrium until

nechanical stimuli induce an increase in the former

with consequent implant loosening. Once this pro=

cess begins, weight - bearing stress does the rest

in loosening the prosthesis.

We cen conclude by stating that implant

materials are not responsible for prosthetic failu=

re. It is more likely to be caused by a cellular

activation factor friggered by mechanical micro-

stress.

CAPTIONS

1) Tomographie presentation of biopsy sites,

2) A - Tc99 bone scan shows two diaphyseal "hot

spots",

B -marked leukocyte bone scan of the same case

(as above) to be considered negative.

3) Numerous histiocytoid cell with vascuolized cy=

toplasm in contact with bone tissue. The bone

surface shows no howship lacunal (May-Grunwald

Giemsa) .

4) TRAP + giant cells surrounded by cement fragments,

5) Tetracycline markers taken from cementless prosthe=

sis trabecular bone interface.

42

BIBLIOGRAFIA

EFTEKHAR N.S., NERCESSIAN 0.: "Incidence and mechanism of failu re of cemented acetabular component in total hip arthroplasty" Orthop. Clin.North.am 19:557 - 566 1988.

FREEMAN M.A.R., BRADLEY G.W., REVELL P.A.: "Observations upon the interface between bone and polymethylmethacry= late cement". J. Bone Joint. Surg. (Br) 64 -B: '489 - 493, 1982.

GOLDRING S.R.,S6HILLER AoL;, ROELKE M., ROURKE C.M., O'NEILL D.A.,HARRIS W.H.: "The synovial - 1ike membrana at the bone -

cement interface in losse total hip replacements and its proposed role in bone lysis". J.Bone Joint Surg.65A: 575-584,1983.

HARRIS W .H., MC GANN W.A.: "Loosening of the femoral component after use of the medullary-plug cementing techni= que follow-up note with a minimum fiveyear follow­up". J. Bone Joint Surg. 68A: 1064,1986.

KIM W.C., H.C. :

NOTTINGHAM P., LUBEN R.A., FINE&~N G.A.M., AMSTUTZ "Detection of osteoclast-activating factor in membranes removed at revision total hip arthropl~ stie". Division of Orthopaedic Surgery, University of California, Los ANgeles 90024.

LINDER K., CARLSSON A.S.: "The bone cement interface in hip ar= throplasty". Acta Orthop. Scand. 57: 495-500, 1986.

marks S;C;, POPOF S.N./ "Bone cell byology: the regulation of development, structure and function in the skele= ton". Am. J. An 183: 1 - 44, 1988.

RADZUN H.J. KREIPE H., PARWARESH M.R.: "Tartrate-resistant acid phospatase as a marker for the human mononuclear phagocyte system". Hematol Oncol. 1:321-327,1983.

43

Fig.l

A B

Fig.2

44

Fig.3

Fig.4

45

Fig.5

46

SOLVXNG OF PROSTHETXC PROBLEMS THROUGH BXOCERAMXCS

FRANCO ZAROTTI Bologna

In order to fully realize the significance of the role the bioceramic within the human body, one should look back to the situation of some decades aga in relation to the fracture of the femoral neck in elderly and old people. In orthopaedic and traumatological surgery, in fact, subcapitate fracture and head necrosis have always been the main problem. In the late 1940s Judet seemed to have come to a solution when he applied a methacrylate head ball; but the stern proved too short (only cervical) and there was excessive wear, leading to fracture. The acrylate polymer was also applied -- and it showed a good tolerability -- for replacement of bones of the carpus, radius, ulna, humerus head, etc .• But the difficulties grew incredibly when the problems of hip arthrosis was tackled. After conducting intensive research on the tolerability of acrylic cement, Charnley an Englishman, devised a hip arthroprosthesis to remedy coxitis from various kinds of pathogenesis. That seemed the ideal solution, but in time there emerged the first drawbacks: mobility of the cotyle and the stern, bone wear, not to mention cases -- although very rare -- of infection that were very trouble some and hard to remedy. In the meantime Muller, from Bern, had designed a prosthesis which consiste of a metal femoral stem provided with a bigger head for better load distribution. Both the cotyle and the metal stem were cemented. Then the problem of wear of the plastic cotyle was addressed, because continuous friction of the metal head against the cotyle in an hydrocolloidal environment was responsible for slow-developing abrasion and the release of minimal particles which gathered to constitute a viscous, self-digesting fluid.

47

To obviate such constant metal-plastic degradation, McKee and Farrar developed a metal-metal head-cotyle joint. Cases of seizure were sometimes observed, though. To prevent cotyloid mobility, Ring designed a cotyle to be fixed to the pelvis by a long axial screw. The stern remained cemented. Other systems were devised, but they had their Achelles heels too. The main drawback was still represented by both cotyloid and stern mobilization, and various kinds of etiopathogenesis were observed (as a result of the type of technique adopter, an incorrect positioning, slight infections, intolerance, wear, etc.) • At any rate it is beyond question that cement, on account of its intrinsic and necessary parameters (even when they are chemically reduced), is thermochemically harmful to bone. Plastic wear of the head-cotyle components is conductive to loosenings and favours the formation of marginal bone mush. There exists in fact no plastic resin, however very-high-density it may be, able to withstand prolonged liquid imbibition. For twenty years now the prosthesis has been modified by abandoning the use of cement, barring exceptional ca ses where cement is still applied on the basis of appropriate indications. Wi th regard to the cotyle, anchoring to the pelvis used to be achieved with the help of a finned, self-threading structure that enabled fixation of the metal wi thin bone, wi th the possibili ty of osteoblast invasion. As for the stern, a variety of shapes, volumes, and lengths were devised as weIl as a variety of surface conformations (threaded, striated longitudinally, embossed, harpoon-shaped, madreporic, ect. ). But even in these cases, intolerances were sometimes observed. Apart from the always possible -- though rare -- occurrence of slight infections, certain cases of reactivity were interpreted as the result of wear of the thin interposed cotyle made from plastic resin. For this reason many of the prostheses produced were so designed as to be "interchangeable". Many of the above problems appear now to have been solved by the introduction of well-tolerated, resistant, accurately screened and controlled bioceramics. These favourable results were also made possible by the in-depth research carried out in a number of German laboratories which exploited Rosenthal's knowhow and experience in this field.

An important contribution came from the expertise of Faenza's bioceramicists, who realized biotolerated hydroxyapati te and

48

bioactive glass as well as a number of ceramic-metal applications. The wide-ranging experimentation carried out by Wagner prompted orthopaedic surgeons to opt for this kind of technique, more reliable although economically more expensive. The advantages that have emerged can be summed up as follows: 1) Elimination of cement and consequently also of the reactions

it causes. 2) Elimination of the plastic resin cotyle interposed between

metals. The sliding of the two ceramic-ceramic surfaces on one another is improved wi th the exclusion of friction and of debris formation from erosion. Another, not negligible benefit for the cotyle is the possibili ty of porous adhesion and of osteoblast penetration for durable stability. Similar advantages, in terms of improved adhesion to sorrounding femoral bone, are obtained by coating the metal stem. Bioceramies surfaces have been applied also in other joint parts of the human body, and numerous further applications are envisaged.

49

HAND MP JOINT IMPLANT ARTBROPLASTY: TBE STATE OF TBE ART. VANTAGES AHD DISADVANTAGES OF TBE MOST COMMON IMPLANT

ARTBROPLASTIES.

A. CAROLI, S. ZANASI HAND SURGERY UNIT, UNlVERSITY OF MODENA,

POLICLINICO, LARGO DEL POZZO 71, 41100 MODENA. ITALY

ABSTRACT

The Authors on the ground of their personnel experience and of most re cent literature data, review and analyze indications, advantages, disadvantages and limits of the most used metallic hinged prosthesis, cemented and semiconstrained prosthesis,and plastic interposition devices for MP joint restoration: in particular cemented and semiconstrained implants, even if show an acceptable clinical performance, generally assuring a painfree, stable functional range of motion, present the great problem of durability and salvageability linked to the rigidity of the implant and to the use of cement. Bone absorption and matherial breakdown often negate at long term follow-up many of the early good resul ts obtained, so that these implants are preferably used in older patients. The method of Swanson silicone arthroplasty since 20 years assures long documented maintainance of functional results, not high incidence of complications and essential salvageability. However this method, either in our or in other Authors experience has shown a not very good clinical performance due to the limited average range of motion in flex/ext. and to a not infrequent tendence to rotation and subdislocation which are only minimally limited by the grommets use's artifice. A promising step towards an ideal prosthesis which should be painfree, mobile, stable, durable, salvageable and of technical facility could be obtained with alumina or hydrossiapatitis implant matherial, fixed by tissue ingrowth system, or by a bioceramic resurfacing of the actually "well-functioning" devices to avoid problems of bone intolerance.

INTRODUCTION

Classification of hand implants can matherials, (rigid-metals, semirigid

be high

done according to density polymers,

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flexible elastic polymers or elastomers),or to the basic type (flexible hinge, interlocking hinge with/without fixation pin, ball and socket implant, interposition arthroplasty, condilar replacement implant), or to the method of fixation (mechanical interlocking with a stem, collar sleevelike cap or grooving on implant, cementing, direct chemical bonding I ike the organic fixation of bioceramic with host tissue, ingrowth of tissue into the implant, encapsulation around a smooth surface silicone implant by capsuloligamentous reconstruction, fasteners as pins, screws, staples, nails and so on). On the ground of these features we can appreciate that since the early 1960's hand implants have evolved from three basic design concepts: metallic hinged prosthesis (Brannon and Klein, 1960; Flatt, 1961), cemented and semiconstrained prosthesis (steffee, 1981; Schultz, 1982; Biomeric, 1978), plastic interposition in form of silicone rubber marketed as silastic (Swanson, 1966), polypropilene (Nicolle and Calnan, 1972), silastic and dacron (Niebauer, 1967). Our aim is to investigate advantages and disadvantages of these most common prosthesis.

FLATT PROSTHESIS (1961). It's a metallic hinged prosthesis p~rmitting only flex/extension movements ; its fixation 1S obtained by mechanical interlocking with the two stainless steel malleable prongs of each stem. The hinge is secured by a screw after stems fitting. The prosthesis requires at least 4 weeks of immobilization to assure interprogs ingrowth of bone and fibrous tissue able to avoid rotational instability. At long term follow (6-14 years) is shown a relatively weIl preserved functional range of motion despite the elapsed time, even if average ROM is quite poor ranging from 15 (Flatt, 1972) and 24 degrees (Blair, 1984); frequent is pain relief and a maintained considerable improved appareance of the operated hand regards preoperative evaluation. On the contrary due to the rigidity of the implant there is a 87% of cases of periprosthetic prongs and hinge masssive bone reabsorption with implant migration and metacarpal (44%) or first phalanx (69%) cortex perforation. (Fig. 1)

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Seldom is appreciable a bone production around hinge with stiffness. Furthermore due to the geometry of the prong hinge­interface resulting in high stress and matherial fatigue there is a high prosthesis screw or prongs failure rate(47%). Recurrent ulnar deviation is frequent too(64%). In conclusion according to Blair (1984), actually the Flatt metallic hinged prosthesis can only provide a great historie perspective in implant design and a clinical standard with which future investigations may compare both existing prostheses and implant design innovations.

STEPPEE MARK 111 PROSTHESIS (1981). It's an improved model developed at the Mayo Clinic by Linscheid and Beckenbaugh, composed of a polyethylene proximal female component and a metal alloy distal male one, which snap­fit loosely together; through a volar slot, provides radial ulnar stability in flexion (Fig. 2).

Fixation is obtained by cementing and early mobilization is allowed. Evaluation at 4 years fOllow-up shows it's able to maintain a complete pain relief in 94% of patients, with good correction of ulnar deviation and an average ROM of 46 degrees. Even though fracture and dislocation rate of components is low «1%) an asymptomatic loosening of one or both components was seen at bone -cement interface in a fourth of implants. So, because of the potential complications of loosening and stress shielding, linked to the cement use, and prosthetic failure, its employ is generally limited to severe deformities with softer bone and in the older age .

SCHULTZ PROSTHESIS (1982) Is a semiconstrained device made of a metal phalangeal component and a plastic metacarpal one with ball-in socket articulation which fixation is obtained by cementing. Early mobilization is allowed. At 11 years follow-up (Adams, 1990) in rheumatoid patients main limits are a progressive decreasing

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with time of ROM up to the average value of 10 degrees, constant recurrence of ulnar deviation and proximal component neck fracture (39%) with extensive periarticular osteoproduction (22%) and bone- cement interface X-rays lucency (80%) even if no prosthetic loosening was seen.

BIOKERIC PROSTHESIS (1978) Is composed by a highly flexible rubber matherial firmly bound to a central ti tanium core hinge and sterns. A central pin bonded to adjacent bion-elastomer provides a 90 degrees of flex/extension motion. Oue to the fact its fixation is obtained by cementing, no joint immobilization is required (Fig. 3).

At 3 ys. average f. u. in rheumathoid and postraumatic patients (Caroli and Zanasi, 1989) verified that the implant firmly assures an average ROM of 55 degrees without any pain and in presence of a constant excellent joint stability, complete correction of ulnar or rotational deviation and good cosmetic result in spite of a progressive deterioration with time of X-rays results (bone reabsorption around sterns in 75% of cases, subdislocation in 40%, buckling in 24%, bone fracture in 4%, asymptomatic loosening in 25%)

SWANSON FLEXIBLE IKPLANT ARTHROPLASTY (1966) Consists of an intramedullary stemmed one piece implant made from medical grade Hp 100 silicone alastomer, matherial highly resistent to flexion fatigue failure. Midsection of the load­distributing flexible hinge implant maintains vertical stability and acts as both a spacer and a flexible hinge mechanism. In fact it is not areal prosthesis but it's a spacer whose fixation is obtained by encapsulation awith possibility of early mobilization of the treated joint. The rationale of this form of "implant arthroplasty" is basesd on two main concepts: first, the implant is a dynamic spacer acting as internal mould around which a new ligamentous system develops during postoperative mobilization (encapsulation concept); the intramedullary implant stems help in proper joint alignment and prevent displacement. Secondly the smooth stems of implant are included in the encapsulation process and the

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implant glides a few mm . on flex/ext. movements (gliding principle, piston effect): the stress forces around the implant are spread over a broad section so it causes less stress to the surrounding bone.

Follow-ups ranging from 5 to 9 ys . show a relatively well preserved functional range of motion with time (mean value of 40 degrees, with a range of 17-61 degrees in 976 implants considered among 5 cases series; a very reliable improving of pain relief, and a considerable improved appareance of the operated hand. Neverthelesss disadvantages in rheumathoid patients are not to underevaluate : average 22% of implant fracture, bone reabsorbtion around stems ranging from 24% (Vahvanen, 1986) to 65% (Hagert 1975), bone production around the hinge with joint stiffness ranging from 9% (Vahvanen, 1986) to 34% of Blair(1984) and 50% of Hagert (1975), dislocation or encasement ranging from 1% (swanson, 1972) to 23% (Vahvanen, 1986) , siliconitis phenomenon «1%) and a high rate (10-45%) of recurrence of finger deformity. However, viewing the whole picture, the results ought to be considered positive since, in spite of X-Ray deterioration, clinical improvement is maintained over the years, as demonstred by its usual employ in about 600.000 patients in the last 20 years all over the world.

Fig . 4: Swanson prosthesis with its grommets

NICOLLE AHn CALNAN PROSTHESIS (1972) To improve the clinical and radiographie performance of the Swanson model Nicolle and Calnan developed a similar hinged prosthe'sis in polypropylene, deeming this matherial more resistant than silicone, with an hemispherical smooth capsule of silicone rubber, abuting the bone ends and maintaining a constant joint space, covering the hinge in order to avoid hinge fibrous tissue invasion and to reduce machanical trauma of the overlying soft tissues, allowing free gliding of tendons. The prosthesis allows 90 degrees of flex/ ext .

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movement and requires at least 3 weeks of immobilization to assure stable fixation of the interlocked semirigid stems . (Fig.5)

Medium-short term fOllow-up results show a relatively weIl preserved good functional range of motion (38 degrees, Griffith, 1975,- 59 degrees, Nicolle, 1972), a considerable improved appareance of the operated hand, and a constant pain relief. On the other hand ther is a deterioration again in the performance of the prosthesis in the time, generally due to mechanical failure (6% of fractures or dislocations in the stems) with recurrence of the deformity, bone reabsorbtion around the stems and at the resection level, recurrence of ulnar deviation. However, compared with other designs of prostheses, polypropylene is shown to be at medium term follow­up a more durable matherial than either silicone or metal .

NIEBAUER PROSTHESIS (1967) A way to improve the general performance of Swanson model is thought could be implant stable fixation by ingrowth of the fibrous tissue into a dacron mesh coated stem surface which should be assured to the bone with some ties. On the ground of this concept Niebauer developed a siliconedacron hinged prosthesis, permitting 90 degrees of flex/extension motion, in which only the biconcave silicone rubber molded hinge is free from dacron mesh coating . Naturally the prosthesis requires at least 3 weeks of immobilization to assure ingrowth of bone and fibrous tissue.

Even though the implant assures at follow-ups ranging from 6.5 to 11.5 years a relatively weIl preserved functional average ROM -ranging from 29' (Derkash, 1989) to 64' (Goldner, 1977)- with time, a very reliable improving of pain relief and a considerable improved appareance of the operated hand, disadvantages are severe and common: infact due to either granuloma from particles of dacron or silicone, or pressure, or rheumathoid pannus or a comhination of all factors, there is a

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high rate (87%) of bone reabsorbtion around stems, with implant buckling, late recurrence of palmar subluxation (44-56% of joints are instable), late dislocation and loosening. Futhermore besides a 66% of cases of hypertrophie spurring , there is a high implant failure rate, ranging from 6% (Derkash, 1986) to 38% of cases (Beckenbaugh, 1976), and ulnar deviation recurrency (44-80%).

Fig. 6: Niebauer prosthesis

On the ground of our personnel experience with preferred implants we report a medium term comparl.son follow-up study between Biomeric cemented prostheses and not cemented Swanson flexible implants with grommets, in severe rheumathoid deformi ties of MP joints. In this study we have taken in consideration only cases in which Swanson implants were used with grommets, which are thin titanium shields, recently developed by Swanson, to provide an unattacched, smooth and durable interpositional titanium between the silicone implant and bone. Grommets should protect the implant from sharp bone edges in areas where abrasion, wear, cutting are most likely due to palmar subluxation forces exerted on implant during flexion

MATHE RIAL AND METHODS

Between february 1985 and November 1987 we have made 24 MP joint Biomeric cemented implants in 8 patients (4male, 4 female) with mean age of 40.7 ys. (range 28-53) affected by rheumatoid arthritis since average period of 7ys. and at the 4th and 5th Larsen X-rays stage. All patients were reviewed with an average follow-up of 43.2 months (range 34-55).

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Conversely between July 1987 and july 1989 we have used 48 MP joint Swanson implants. 24 were used with grommets (10 pairs at the index, 10 pairs at the middle, 2 pairs at the ring and at the small finger respectively), and 24 without grommets in the same patient in which titanium shields were employed. Patients were 12 (6male and 6 female) with mean age of 55.4 ys. (range 49-69) affected by rheumatoid arthritis since an average period of 12.7 ys., and at the 4th and 5th Larsen X-ray stage. All patients returned for control: average fOllow-up was of 21.1 months (range 14-38)

RESULTS

Clinical and X-rays findings are shown in the following tables:

BIOMERIC

PREOP. 24/24

POSTOP. 2/24

PREOPERATIVE

TABLE 1 PAIN

SWANSON WITH GROMMETS

PREOP. 16/24

POSTOP. 2/24

TABLE 2 TOTAL ACTIVE MOVEMENT

BIOMERIC

SWANSON WITHOUT GROMMETS

PREOP. 24/24

POSTOP. 0/24

POSTOPERATIVE FLEX. EXT. TAM FLEX. EXT. TAM

63 -4 59 68 -9 59

SWANSON WITH GROMMETS

PREOPERATIVE FLEX. EXT.

58 -13 TAM

45

POSTOPERATIVE FLEX. EXT. TAM

42 -12 30

SWANSON WITHOUT GROMMETS

PREOPERATIVE FLEX. EXT.

70 -4 TAM 66

POSTOPERATIVE FLEX. EXT. TAM

36 -8 28

PREOPERATIVE 24/24

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TABLE 3 INSTABILITY

BIOMERIC

TABLE 4

POSTOPERATIVE 0/24

RECURRENT ULNAR DEVIATION >10· <30·

BIOMERIC

PREOP. 16/24

POSTOP. 0/24

SWANSON WITH GROMMETS

PREOP. 10/24

POSTOP. 6(0*)/24

SWANSON WITHOUT GROMMETS

PREOP. 12/24

POSTOP. 4(2*)/24

(*)= Early postoperative result

IMPLANT FRACTURE IMPL. DISLOCATION IMPL. ROTATION INFECTION SILICONITIS

PROX. BONE REABS. DIST. BONE REABS. BONE PROD. AT HINGE CORTICAL PERFORATION IMPL. ENCASEMENT MP VOLAR SUBLUXATION IMPL. LOOSENING

TABLE 5 COMPLICATIONS

BIOMERIC

0/24 0/24 0/24 0/24

SWANSON WITH

GROMMETS

0/24 0/24 2/24 (9%) 0/24 0/24

TABLE 6 X-RAY FINDINGS

BIOMERIC

8/24 16/24

0/24 2/24 2/24 6/24 6/24

SWANSON with

GROMMETS

12/24 16/24

2/24 0/24

14/24 4/24

SWANSON WITHOUT GROMMETS

2/24 (9%) 0/24

10/24 (24%) 0/24 0/24

SWANSON without GROMMETS

14/24 12/24

6/24 0/24

16/24 0/24

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Fig 7: X-ray main complications of Biomeric (left) and Swanson (right) implants

In summary Biomeric cemented prostheses , at average four years follow-up, are generally able to assure complete relief of pain, maintaining a very good average TAM (60°), an excellent stability without any ulnar deviation recurrence or cases of components' fracture. Hand cosmetic improvement is appreciable too. On the other hand besides the relatively more difficult surgical technique, these devices, due to the rigidity of the implant, could present , with time, a possible symptomatic increasing percent of loosening and, due to the excessive bone remotion, possible problems of salvageability in case of failure, so that could be preferable their employ in most severe deformities and in older patients.

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Swanson implants in fact are able to assure a complete relief of pain, with quite good stability due to the encapsulation process, and an improvement either of global hand functionality or hand cosmetic appareance; grommets employ seems relatively able to decrease wear phenomenon, implant fracture and siliconitis occurrence. Furthermore is possible implant removal or substitution without deteriorating an useful function which maintainance is documented since over 20 years. However, disadvantages are not to underevaluate: the average TAM is quite fair (30·), frequent is ulnar deviation relapse and implant rotation. There is always a progressive deterioration of the X-Ray picture with bone reabsorbtion around the two stems and osteoproduction around hinge and implant buckling. with time increases the possibility of siliconitis too. Finally their use is controindicated in case of excessive softness of the cancellous bone (due to a greater risk of occurrence of implant rotation), and excessive contracture of soft tissues with shortening of the fingers because the necessary soft tissues release could not allow a satisfactory soft tissue stabilization.

CONCLUSIONS

The ideal implant should be painfree, mobile, stable, durable, salvageable: up to day no one prosthesis seems able to answer most of these conditions. New interesting solutions involve the research od improved matherials, which are inert, durable and which produce minimal wear particles, and the type of fixation: Fixation of implants by tissue ingrowth is a subject of greater attention. Bioceramies or bioglasses with 200-400 micron sized holes into which bone can grow appears may have promise . The problem remains the immobilization for the period of time

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required to gain adequate bone ingrowth: a too quickly mobilization (less than three weeks) may produce shearing of bone ingrowth so that only fibrous tissue remains; a too long immobilization will result in stiffness. What's about the next future? Biomatherial evaluations, including flexion-extension tolerance, stretching, twisting, extraction tests and histological examinations of affinity for bone demonstrated that earlier bioceramic implants have characteristics superior to those of all previous finger implants. FOllow-up studies, even though too short yet, have been most encouraging with satisfactory functional recovery and no surrounding bone or intrinsic implant problems (Kazuteru Doi, 1984; Minami m., 1988).

REFEREHCES

1. ADAMS B.D., BLAIR W., SHURR D.G., Schultz metacarpophalangeal arthroplasty: a long term fOllow-up study, J. Hand Surg. 15A: 4, 641, 1990.

2. BECKENBAUGH R.D., OOBYNS J.H., LINSCHEID R.L., BRYAN R.S., Review and analysis of silicone -rubber metacarpophalangeal implants, J. Bone joint Surg. 58A: 4, 483, 1976.

3. BLAIR W.F., SHURR D.G., BUCKWALTER J.A., Metacarpophalangeal joint implant arthroplasty with a silastic spacer, J. Bone Joint Surg. 66A: 3, 365, 1984.

4. BLAIR W., SHURR D.G., BUCKWALTER J.A., Metacarpophalangeal joint arthroplasty with a metallic hinged prosthesis, Clin. Ort. Rel. Res. 184: 156, 1984.

5. CAROLI A., ZANASI S., CRISTIANI G., MARCUZZI A., GUERRA M., PANCALDI G., Indicazioni e limiti delle protesi cementate Biomeric nelle gravi distruzioni di natura reumatoide e postraumatica delle articolazioni MF ed IFP. Atti deI congresso congiunto della Societa Italiana di Ricerche in Chirurgia e della Societa Italiana di Fisiopatologia Chirurgica. Monduzzi Ed., Bologna, 1990.

6. DERKASH R.S., NIEBAUER J.J, LANE C.S., Long term follow-up of metacarpophalangeal arthroplasty with silicone dacron prostheses, J. Hand Surg. 11A:, 4, 553, 1986.

7. DOI K., KUWATA N., KAWA I S., Alumina ceramic finger implants: a preliminary biomatherial and clinical evaluation, J. Hand Surg. 9A:5, 740, 1984.

8. FLATT A.E., ELLISON M.R., Restoration of rheumatoid finger joint function. A follow-up note after 14 years of experience with a metallic hinged prosthesis, J. Bone Joint Surg. 54A: 1317, 1972.

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9. GOLDNER J.L., GOULD J.S., URBANIAK J.R., McCOLLUM D.E, Metacarpophalangeal joint arthroplasty with silicone-dacron prosthesis (Niebauer type): six and half years' experience, J. Hand Surg. 2A: 3, 200, 1977.

10. GRIFFITHS R.W., NICOLLE F.V., Three years' experience of metacarpophalangeal joint replacement in the rheumatoid hand, The Hand 7: 3, 275, 1975.

11. HAGERT C.G., ElKEN 0., OHLSSON N.M., ASCHAN W., MORRIN A., Metacarpophalangeal joints implants, Scan. J. Plast. Reconstr. Surg. 9: 145, 1975.

12. MINAMI M., YAMAZAKI J., KATO S.,ISHII S., Alumina ceramic prosthesis arthroplasty of the metacarpophalangeal joint in the rheumatoid hand. A 2-4 year follow-up study, J Arthroplasty 3: 2, 157, 1988.

13. NICOLLE F.V., CALNAN J.S., A new design of finger joint prosthesis for the rheumatoid hand, The Hand 4: 2, 135, 1972.

14. STEFFEE A.D., BECKENBAUGH R.D., LINSCHEID R.L., DOBYNS J.H., The development, technique, and early clinical results of total joint replacement for the metacarpophalangeal joint of the fingers Orthopedics, 4: 2, 175, 1981.

15. SWANSON A.B, Flexible implant resection arthroplasty, The Hand 4: 2, 119, 1972.

16. SWANSON A.B., DE GROOT-SWANSON G., Flexible implant arthroplasty of the digits. Indications, methods and results, Acta orthopaedica Belgica 51: 4, 679, 1985.

17. VAHVANEN V., VILJAKKA T., Silicone rubber implant arthroplasty of the metacarpophalangeal joint in rheumatoid arthritis: a follow-up study of 32 patients, J. Hand Surg. 11A: 3, 333, 1986.

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ALUMINA TOTAL JOINT REPLACEMENT OF THE FIRST METATARSO PHALANGEAL JOINT~ A BIOMECHANICAL STUDY OF THE DESIGN.

S.Giannini, A.Moroni, A. Kraiewski*, A. Ravaglioli*, R. Martinetti*, C. Farina**. Bologna University, Rizzoli Orthopaedic Institute, Bologna,Italy

* IRTEC, CNR, Faenza, Italy ** Finceramica, Faenza, Italy

ABSTRACT

A new prosthesis for the total joint replacement of the first metatarso-phalangeal joint was developed. The prosthesis has a phalangeal and a metatarsal component , both are alumina made. The design is anatomie and was chosen after a study of 20 human cadaveric first metatarso phalangeal joints. The bone prosthesis fixation is uncemented. Three sizes are available.

INTRODUCTION

First metatarso-phalangeal joint degenerative diseases can be treated by several surgical techniques. Among these the Keller and Heuter-Mayo operations are the most popular. We performed the Keller technique for several years obtaining good clinical results. However we decided also to develop a metatarso-phalangeal joint pros thesis to find a remedy for cases of severe reduction of the first metatarso-phalangeal joint range of motion, joint instability and insufficiency of the first metatarsal ray. Nowadays there are several metatarso-phalangeal prostheses available. The most popular model is the Swanson silastic prosthesis. The results of this prosthesis seem to be not completely satisfactory due to wear debris. Thus we investigated a new model. The characteristics of this new model should be: biocompatibility, mechanical strength, low wear, anatomical design and uncemented bone prosthesis fixation.

PROSTHETIC MATERIAL

Like prosthetic material we chose alumina because its biocompatibility, bioinercy, low wear and good mechanical performances. In comparison to the metallic alloys utilized in several total joint prostheses (AISI 316 L, Cr-Co Alloys, Titanium) alumina has the advantage to be bioinert. Thus metallic ion release is avoided. The very low wear of alumina was demonstrated by several authors. (1)

The positive alumina characteristics were confmned in many clinical studies reporting the results of hip total joint prostheses with alumina components. Failures were due mostly to fatigue fractures due to the manufacture technique. (1)

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At the present time the improvements of manufacturing technique allows to avoid these problems. Alumina powder size must be less than 1 micron to obtain both a good mechanical strength and a low superficial roughness (less than 3 micron) of the fmal prosthesis. (1)

Other Authors recently utilized alumina prostheses for the total joint replacements of small joints. Yonenobu, et al., developed an alumina prosthesis to be used in the metatarsal bones lenghtening. (2) Minami, et al., investigated an alumina metacarpo phalangeal total joint prosthesis. (j)

In a previous study we described the sheep femoral bone response to transcortical alumina implants confmning the excellent biocompatibility of this bioceramic material. (4)

PROSTHETIC DESIGN

Twenty cadaveric human frrst metatarso phalangeal joints were analyzed in order to develop an anatomical design of the prosthesis . In each cadaveric joint we calculated the metatarsal head diamaters (A,B,C); the length of the diaphyseal metatarsal canal (0); the diameters of the diaphyseal metatarsal canal (E,F); the distance between the metatarsal basis and the center of the metatarsal head (0); the diameters of the phalangeal basis «H,Hl,L); the length of the proximal phalangeal metaphysis (M); the distance between the more and less prominent points of the phalangeal basis (N); the distance between the phalangeal basis and the end of the phalangeal medullary canal (P); the diameters of the phalangeal medullary canal (Q,R). (Fig.I)

Three main dimensional groups were discovered by these measurements and three sizes of the prosthesis were developed.

Figure I In this figure the measures calculated in the cadaveric first metatarso phalangeal joints are shown.

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BIOMECHANICAL CONSIDERA TIONS

From the biomechanical point of view fIrst of all we emphasize that our metatarsal prosthesis component was designed to be implanted after a metaphyseal osteotomy 15° anteriorly and distallyoblique to the longitudinal metatarsal axis. In this way the compression stress on the osteotomy surface are prevalent. Moreover with an oblique osteotomy the bone prosthesis contact area is increased. (Fig.2)

Length of the metatarsal prosthesis stern was chosen to achieve a fIrm stability of the bone implant fixation. To avoid rotating displacements the stem has a trilobate section shape. To reduce the possibility of displacements on the frontal plane lateral prominences are present in the metatarsal prosthetic head.

Figure 2 In this fIgure the metatarsal component is shown: this component was designed to be implanted following a metatarsaI osteotomy 15° oblique to the metatarsal axis.

The phalangeaI prosthetic component shape and length were measured to obtain a good fIlling of the phalangeal medullary canal. Prominences are present in this stem to avoid rotating displacements. The bending radius of the basis of the phalangeal prosthetic component is always greater than the bending radius of the head of the metatarsaI component. (Fig.3)

Figure 3 In this fIgure the phaIangeaI component is shown. Notice the prominences of the stem.

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Our prosthesis allows plantar flexion range of motion of 30°, dorsal flexion of 50°, and varus and valgus movements of 10°. (Fig.4-5)

Figure 4 Photograph showing the metatarsal and phalangeal components implanted in a cadaveric specimen.

Figure 5 Dlustration showing the range of motion of the prosthesis.

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REFERENCES

1. Ducheyne, P., J.Biomed.Mat.Res. 1987, 21, 219-336.

2. Yonenobu, K., Tada, K., Takaoka, K., Tsuyuguchi, Y., Ono, K., J.Biomed.Mat.Res. 1986.2Q, 1249-1256.

3. Minami, M., Yamazachi, J., Sadatoshi, K., Seiichi, 1., Joum. of Arthrop. 1988,1. 157-166.

4. Moroni, A., Giannini, S., Zaffe, D., Ravaglioli, A., Kraiewski, A., Venturini, A., Pompili, M., Pezzuto, V., Trinchese, L., in Bioceramics vol 2, editor G.Heimke, Cologne, Germany, 1990.

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TITANIUM HYDROXILAPATITE COATEO METACARPO-PHALANGEAL ANO INTERPHALANGEAL IMPLANT

EDOARDO CARLO MARINONI Associated Professor of Hand Surgery, University of Milan, I Orthopaedic Department S. Gerardo Hospital, Monza, I.

GIUSEPPE VENINI Director of DISPO-Research Laboratories, Cologno Monzese, MI,I

ANTONIO RAVAGLIOLI Research Institute for Ceramics Technology, CNR, Faenza, I.

ABSTRACT

At present, the artificial replacement of the metacarpo­phalangeal or interphalangeal joints is still an unsolved problem. Clinical experiences have releaved the limits of cemented prostheses (Flatt 's prosthesis , Schul tz' s prosthesis and Biomeric) and also of articular spacers (Swanson's implant, Niebauer's prosthesis, Swanson's-gromets implant). So we have planned an uncemented implant, a titanium Dac Blu hydroxiIapatite coated prosthesis. In this paper, we explain the features of the prosthesis , giving an example of its clinical use. The Dac Blu hydroxiI­apatite coating improves the biostability and the bio­compat ibili ty of the implant to the host bone. To confirm this fact, we report histological data on the interface between Dac Blu hydroxiIapatite screws and human bone. Besides, our aim is to extend the casuistry and the follow-up.

INTROOUCTION

Arthroplasty in the joints of the hand is indicated in the presence of pain, stiffness and/or deformity. Arthrodesis is, almost always, the major alternative and should be considered with regard to the potential problems and benefits ant icipated when elect ing arthroplasty. If arthroplasty is elected, it may be performed with bone and soft tissue,with

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interpositional materials or with complete joint replacement. Soft tissue arthroplasties of the metacarpo-phalangeal (MP) joints were desbribed by many Authors (Smith-Petersen, Aufranc, Larson, Riordan, Fowler, Tupper, Vainio). In 1959 Brannon and Klein and later Flatt developed metallic hinged prostheses for MP joints. In 1960 Swanson and Niebauer developed silicon rubber prostheses. Niebauer's design utilized a hinge concept with ingrowth fixation, while Swanson's design were basically spacers to improve the stability of resectional arthroplasties (Green, 1982). Application of these devices was initially enthusiastic, but some problems with breakage, recurrence of deformity and inability to improve grip and pinch have consistently led others to seek new solutions. Some Authors preferred cemented prostheses (Steffe, hinged Mayo prosthesis, Heiple with Biomeric finger joint, et al.). We have planned an uncemented implant, using t itanium as metal substrate. The prosthesis is also available with Dac Blu hydroxiIapatite coating (plasma spraying technique), to improve the osseo­integration of the implant.

MATERIALS AND METHODS

The prosthesis (Fig. 1) consists of 3 titanium alloy elements (Ti 6A14V), assembled together. This alloy is excellent for resisting corrosioni it is biocompatible and presents a high mechanical toughness compared with its lightness. Its thermic dilatation coefficient and elastic modulus are low. In fact titanium is bioinert, which means chemically inert to the host bone. In order to improve the interface with host bone, the implant is also available with Dac Blu hydroxiIapatite coating (Plasma spraying technique). Dac Blu hydroxiIapatite is a high bioactive ceramic and forms a tight chemical bond with living bone. Two implants elements are fixed, because they are screwed in metacarpal or phalangeal endostal canal. Afterwards, the central joint is assembled (Fig. 1). The specific strumentation consists of reamers to prepare the endostal canal (metacarpus or phalanx) and of dirne (spacers) to control the obtained space after articular surfaces section. Several sizes of fixed elements and central joint are available, in order to promote higher or lower stability of the implant according to the clinical case.

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RESULTS

Three months ago, a voluntary subject suffering from ankylosis of proximal interphalangeal of left index was treated with this prosthesis. In postoperative, articular stability and motion range are sat isfactory. Index extension is full, volar flexion from 180 0 to 100 0 • Grip and pitch are good (Fig.2).

DISCUSSION

By conceiving this new implant solve two problems; to obtain an implant and, at the same time, integration of the new resolution. We think that the first problem

we essentially wanted to easy to apply, non cemented to obtain a perfect osseo-

was solved very weIl by screwing the implant elements to the bone; in fact, the endostal canal (of metacarpo and/or phalanx) is prepared by means of a conical cutter, whose size is only just a little smaller than that of the titanium screw (hydroxiI­apatite coated or non-coated), which is self-threading. This resolut ion assures to the whole "bone-implant" an excellent immediate mechanical stability. Titanium Dac Blu hydroxiiapatite coating allows to solve the second problem: the starting mechanical implant stability must change into an achieved biological stability, that is we must have a complete osseointegrat ion of the implant. With histological tests, we checked the best integration between human bone and Dac Blu hydroxilapat ite screw (Marinoni et al. 1987). Our histological studies about undecalcified specimens obtained from human bone and Dac Blu hydroxiiapatite screw abrasion put in evidence an active osteogenesis on the interface between Dac Blu hydroxiiapatite screw and human host bone (Fig. 3). Several Authors (Denissen, 1980; Geesing, 1988; Kawamura, 1987) came to the same conclusion. This result surely proves both Dac Blu hydroxiiapatite biostability and its bioactivity.

CONCLUSIONS

shows a solut ion of hand joints

several surgeons

We think that this already applied implant to some problems, which limited the use implants so far; this is the reason why often choose arthrodesis techniques. Hydroxiiapatite coated implant doesn't need the use of

70

bone cement and so we can eliminate one of the possible defect to the detriment of human bone. The implant is not only a spacer, in fact it is very different from all "Swanson type" implants: we want to obtain a mechanical stability immediately and the joint function as soon as possible. To our mind, indicat ions for above ment ioned implant remain a discussion matter; during this first phase of study and applicat ion we think that every case must be valued individually; "severe osteoporosis", as normally by rheumatic hand, is a contro-indication for the use of this implant. On the contrary, articular stiffness of arthritis and/or post-traumatic origin are surely the preferable indication. Furthermore , the resect ion width on art icular surfaces by implant application are minimal; this means that, if the implant must be removed, it results quite easy to return to arthrodesis for a functional recovery.

Figure 1. Titanium implant (left) and Dac Blu hydroxiIapatite titanium coated implant (right)

71

Figure 2. PIP joint of the index: after and post operative AP radiograph (follow-up: 3 months)

Figure 3. Non decalcified specimen showing new bone formation between the Dac Blu hydroxilapat ite and the host bone. On the right the same picture as the left, in greater enlargement, showing the normal structure of the newly-formed bone of the interface.

72

REFERENCES

1. Denissen;H.W., Tl}e linkage between Apatite Implant Material and Living Bone. Ultramicroscopy 5: 124-128, 1980.

2. Geesink,R.G.T., Groot,K.de, Klein,C.P.A.T., Bone bonding to apatite coated implants. J Bone Joint Surg 70B: 17-22,1988.

3. Green,D.P., Operative hand Surgery, Churchi1l Livingstone, New York, Edimburgh, London and Melbourne, 1982, pp 141-161

4. Kawamura,M., Iwata,H., Sato,K., Miura,T., Chondro-osteo­genetic Response to crude Bone Matrix Proteins Bound to Hydroxiiapatite. C1in. Orthop. 217: 281-292, 1987.

5. Marinoni,E.C., Simonatti,R., Memeo,A., Mosca,L., A Study of the Interface Human Bone and Hydroxi1apatite.A morpho1ogica1 and morphometric investigation. Italian Journal of Ortho­paedics and Traumatology XIII, 527-534, 1987.

73

HYDROXY APATITE COATED PLATE IN THE SURGICAL TREATMENT OF THE FOREARM

NON UNION WITH BONE LOSS

G.F. Zinghi, A. Moroni, L. Specchia, P. Bungaro, G. Gualdrini, G. Rollo, C. Sabato

Rizzoli Orthopaedic Institute, Via Pupilli 1, Bologna 40136, ltaly

ABSTRACT

From 1974 to 1990 23 forearm nonunions with bone loss were operated on at the third Department of Rizzoli Orthopaedic Institute. Length of bone loss ranged from 0.5 to 8 centimeters. The forearm nonunions were operated on of internal fixation with metallic plate; in 8 patients hydroxyapatite (HA) coated plates were implanted. A cortical-cancellous bone graft was interposed between the nonunion segments. A cortical bone graft was implanted opposite to the plate. Union was achieved in 21 cases. Functional results were evaluated according to the prono-supination range ofmotion. We had 10 excellent results. 5 good. 6 fair and 2 poor. The restoration of the shape and length of the forearm long bones is considered indispensable to obtain goodjunctional results. The HA coated plate advantages could be osteoconductive property. bone direct apposition and bonding. no fibrous tissue formation near the plate. not necessary removal.

INTRODUCTION

The foreann nonunions with bone loss modify significantly the anatomo-functional units of

the foreann that are radius, ulna, radio-ulnar joints, interosseous membrane and ligamentus

quadratus.(1) The integrity of these units is very important for the prono-supination range of

motion. Infact when the relationships of shape and length between radius and ulna are

modified a reduction of the prono-supination range of motion occurs. (1)

The aim of the treatment must be both the union and the salvage of the prono-supination

range of motion. (1-2-3-4-5)

74

MATERIALS AND METHODS

From 1974 to 1990 23 forearm nonunions with bone loss were operated on at the third

Department of Rizzoli Orthopaedic Institute. Ninenteen (82.6%) patients were male, 4

(17.4%) female. Average age was 31.5 years, range 14-47 years. Thirteen (56.5%)

nonunions followed both mdius and ulna fra.ctures, 3 (13%) mdius fra.ctures, 5 (21.8%) ulna

fractures, 2 (8.7%) ulna fractures with dislocation ofthe mdial head. All these fra.ctures had

been previously operated on: 16 (69.5%) fra.ctures had been operated on one time, 5 (21.8%)

two times, 2 (8.7%) three or more times. First surgical treatment bad been open reduction and

internal fixation by intramedullary naH in 6 (26%) cases, by plate in 8 (34.8%) cases, by

intramedullary naH in one of the forearm bone and plate in the other one in 7 (30.5%) cases,

by external fixation in 2 (8.7%). In 8 (34.8%) patients nonunion with bone loss was in the

radius, in 15 (65.2%) patients in the ulna. Length of bone loss ranged from 0.5 to 8

centimeters, average 2.74 centimeters. In 1 case signs ofinfection were present at the time of

our treatment. Previous infection was detected in the anamnesis of 2 other patients. Surgical

treatment of the nonunions was undenaken after an average of 19.4 months from trauma

(range 5 months - 6 years). The foreann nonunions were operated on by internal fixation with

metallic plate. A cortical-cancellous bone graft was interposed between the nonunion

segments. A cortical bone graft was implanted opposite to the plate to achieve a very stable

fixation. Fixation was performed with forearm in intermediate prono-supination position. In

the 8 more recent cases an AISI 316 L plate coated·of hydroxyapatite (HA) by air plasma

spray was used (Biocoatings, Flametal, Fornovo Taro,ltaly). In 19 cases homologous bone

grafts and in 4 cases autologous ones were implanted. A 3 months pIaster cast immobilization

period followed surgery. Average follow-up was 3.87 years (range 8 months - 13 years).

Functional results were evaluated according to the prono-supination range of motion. We

evaluated as excellent the cases with prono-supination range ofmotion from 175° to 140°,

good from 140° to 70°, fair less than 70° with possible intermediate prono-supination

position, poor less than 70° with absence of intermediate prono-supination position.

RESULTS

Anatomo-mdiographic results were positive in 21 (91.3%) patients. In 2 cases (8.7%) the

anatomo-mdiographic result was negative for the onset of a deep infection. In these 2 cases

uncoated plates were used. In one of these cases, at a follow-up of 4 years, a bone loss of 5

centimeters was still present. Functional results were excellent in 10 (43.5%) cases, good in 5

(21.8%), fair in 6 (26%) and poor in 2 (8.7%). (Table 1). As post-surgical complications

we observed 1 haematoma, 2 radial nerve palsy and 1 skin necrosis.

Excellent

Good

Fair

Poor

Tot.

Figure 1

75

TABLE 1 Funetional results

HA COATED PLATE

6 (75%)

2 (25%)

8

UNCOATEDPLATE

4 (27%)

3 (20%)

6 (40%)

2 (13%)

15

a) 32 years old male patient affected by nonunion with hone loss of radius. b) the radiographie figure 2 years after surgery shows the union. Funetional result was exeellent.

76

Figure 2 a) 29 years old male patient affected by a severe nonunion with bone loss of the radius. HA coated plate was implanted. b) the radiographie figure 16 months after surgery shows the union. Functional result was good.

DISCUSSION AND CONCLUSIONS

First of all we emphasize the ulna greater rate (65.2%) of non unions. On the contrary in our

previously published series of non union without bone loss of the forearm, the radius

involvement was greater (85%).(1-2) Secondly we emphasize that the foreann non union

with bone loss seems to be due to a previous inadequate surgical treatrnn.ent, in fact in all our

cases a previous inadequate surgical treatment was performed.

On the basis of our experience we reccomend to evaluate accurately the anamnesis

searching signs of infection, if infection is detected the surgical treatment must be very

careful.

Our surgical technique utilizes a cortico-cancellous bone graft, interposed between the

non union segments, to restore the shape and length of the forearm long bones. Any

shortening of the forearm is not tollerated determining radical cosmetic and functional

changes. (1)

77

Tbe opposite hone graft is useful to obtain a very stable fixation. Stable rigid fixation is

difficult to achieve in these cases which are characterized by a large amount of porotic hone. (2)

Comparison between the HA coated and uncoated plate results is not possible because the

very short follow-up and the few cases. Infact an the cases with HA plates are very recent. So

far our impression is very positive. We decided to use HA coated plate in order to reduce the

metal ion release and the fibrous tissue fonnation near the plate. Infact in our opinion fibrous

tissue could affect the hone remodelling of the hone grafts. On the contrary the HA coating

should promote direct hone apposition and honding. This effect could be very usefull in the

healing process of these severe forearm lesions.

In our opinion removal of the HA plate is not necessary, frrst of an because for its thickness

our plate is well tollerated from the clinical and biomechanical point of view and secondly

because we expect that hone honding and osteointegration of the plate can be achieved.

REFERENCES

1. Zinghi, G. F., Montanari, G., and Specchia, L., La pseudartrosi diafisaria d'avambraccio. It. J. Orthop. Traum .. 1981, 7, 205-219.

2. Ruggieri, F., Zinghi, G. F., and Lanfranchi, R., Le pseudoartrosi diafisarie. Aulo Gaggi, Bologna, 1977.

3. Nicoll, E.A., The treatment of gaps in long hones by cancellous insert graft. L..ßQm; Joint Sorg., 1956,38 B, 70-79.

4. Green, S.A., and Dlabal, T.A., Tbe open graft for septic nonunion. Clin. Orthop., 1983, 117, 124-133.

5. De Buren, N., Causes and treatment of nonunion in fractures of the radius and ulna. L Bone Joint Surg. 1962,44 B, 614-622.

78

A STUDY ON HA-COATED TITANWM DENTAL IMPLANTS

PART I: STRESS ANALYSIS OF DENTAL IMPLANT

JIYONG CHEN, JIMING ZHOU, XINGDONG ZHANG, DERI WAN

Institute of Matarial Science and Technololgy, Sichuan University

Chengdu 610064, P.R. China

SHAOAN WANG and ANYU CHEN

College of Stomatology, West China University of Medical Science

Chengdu 610041, P. R. China

ABSTRACT

The stress distributions around pur kinds 0/ endosteal implants were calculated by using three dimensional finite element stress analysis. Lower stress concentration in the cresteal bone region and the most uniprm stress destribution were observed in the cylindrical im­

plant. For cylindrical HA-Ti endostal dental implants. pur possible stages 0/ implant-bone inter jJce were simulated according to our animal experiment and a8sumption. The results 0/ stress analysis jör the pur stages were usejül pr design optimization 0/ implants and climical application.

1. INTRODUCTION

The design of dental implants should consist of three aspects: shape design,

micromorphological and surface design and a wise selection of materials. The mate­

rials to make implants must be possessed of good biomechanical properties and

This project was surported by the National Natural Science Foundation of China and the National Educational Committee of China.

79

biocompatibility[ 1.2]. The bioactive materials usually formed osseous integration at

implant-bone interface. But the bionert materials mostly forr:ned fibro-osseous inte­

gration. The osseous integration in the transmission of occlusal force and the stress

destribution differed from fibro-osseous integration, the shape design depended on

the type of interface integration. The proper surface roughness and porosity of im­

plants would increase the bone-bonding intensity of implants by both mechanical

interlocking and inceeased surface ares. The proper selection of patients and the

suitable implanting technique and prosthodontics were also important for the suc­

cess of implanting.

The long-term success of any endosteal dental implant was dependent on the

design optimization of implant. A correct design should distribrte applied stresses

properly so that the destructive stresses would not be transfered to the supporting

tissue. In this research, the three dimensional finite element stress analysis was used

to compare the stress distributions around four endosteal implants with defferent

shapes. From the results of stress distribution comparison, we chose the cylidrical

implant. which had lower stress concentration in the cresteal bone region and the

most uniform stress distribution, as the implant to be developed mainly in our labora­

tory. We also investigated the change of stress distribution of HA coated Ti implant

in the interface as the implant-bone interface developed. These results laid the foun­

dation of our design of dental implants.

2. CONSTRUCTION AND CALCULATION OF mE MODEL OF mREE

DIMENTIONAL FINITE ELEMENT STRESS ANALYSIS

2.1 ANATOMIC BASIS OF THE MODEL

A sixty-year-old male teeth-free mandible was chosen as the anatomic basis of

constructing a finite element model. A cylindrical Ti implant. which was 18mm long

and 5mm in diameter, was implanted into the first molar of the left mandible. The im­

planting depth was 12mm. A 30mm long block of mandible, in the center of which

the implant was, was cut off. Then the block was cut into 15 slices along the

buccilingual direction. The locations of the slices in mandible and the thickness of

cortical bone and cancellated bone in every slice were measured by micrometer.

2.2 MECHANICAL PARAMETERS OF MATERIALS

Young's modulus and possion ratio of titanium (TA2) and HA were measured by

resonance method. The parameters of cortical bone and cancellated bone were got

80

from reference[31.

TABLE1

Mechanical properties of materials

Young's modulus Possion ratio Material

(MPa) v

titanium (TA2) 111600 0.35

HA 84340 0.24

cancellated bone 490 0.30

cortical bone 4904 0.30

lamina dura 9807 0.32

2.3 SELECTION OF FOUR KINDS OF IMPLANTS

The shapes of four selected implants were the followings:

II m IV

TONGUE

I II rn IV

IMPLANT

c=J CANCELLOUS

lITllIIIIlI CORTICAL BONE

~ LAMINA DURA

Fig. 1 Buccilingual cross sections of four kinds of implants and their

supporting tissues.

81

I) Screw-shaped (a modified Branemark implant).

II) Cylindrical with three screws on the surface of the implant.

ill) Conical.

IV) Cylindrical with smooth surface.

The maximum diameter of the four implants was 5mm; the lenth was 18mm; the

bottoms were semi-sphere. The thickness of the lamina dura layer surrounding each

implantwas assumed to be O. 5mm.

The cylindrical HA coated titanium implants were simulated to investegate the

change in stress distribution as the implant-bone interface developed. The four pos­

sible stages of implant-bone interface. according to our animal experiment, were as­

sumed to be the folowings: (A) neither lamina dura nor fibrous tissue at the

implant-cancellous bone interface. (B) thickness of the lamina dura around the im-

TONGUE CHEEK

Fig. 2 Two kinds of loading models.

Fig. 3 Buccilingual finit element structure of type IV implant and

its supporting tissue.

82

plant was O. 3mm. (C) the lamina dura was O. 5mm. and (D) the lamina dura

was 1. Omm.

490 N Verlicalloading was applied on the center of top of implant along the axis

of implant. 490 N horizontalloading was applied on the crown post of implant along

the buccilingual direction. The two kinds of loading model were showed in figure 2.

2.4 DIVIDING UNIT

The model was divided into many eight node hexahedrons and six node wedged

pentahedrons. The unit numbers and node numbers of the models for the four kinds

of implants were showed in Table 2.

TABLE 2

Unit number amd node number

model unit number node number

I 1002 1101

n 906 1001

m 858 951

IV 858 951

2.5 ASSUMED CONDITIONS

In this research the following conditions were used:

1) All materials in the models were homogeneous. isotropie and Iinearly elastic.

2) No slip between the implant and its supporting tissue was allowed.

3) The Young's moduli of the cortical bone 1 and the cortical bone 2 were the

parameters of the top of the alveolar ridge and the edge of the mandible respectively.

2.6 CALCULATION

SAP-VI structure analysis program was used in the calculation. The results were

given by the maxium principal stress (f max and minium principal stress (f min at the in­

terface between implant and supporting tissue.

3.RESULTS

Figure 4-6 are the culcalation results. Vertical axis represents principal stress in M Pa.

Positive value indicates tensile stress and negtive value is compressive stress. The

83

9 TONGUE CHEEK

9

......... II co UJ 0... .-J 0 2 -cn cn :::>

UJ cn I-UJ CI: UJ

9 I- > cn cn -l cn <{ UJ

0... CI: m u 0... 0 z 2

0 CI: U 0...

9

IV 0

9T--.---r--~-.---r--~-.---r--.-~

o 6 12 18 24 30 LOCATION(MM;

Fig. 4 Principal stress in bucilingual direction under vertical loading.

point 12.5mm is the bottom of the implant and the points 0 and 25mm are the top of

alveolar ridge.

3. 1 THE STRESS DISTRIBUTIONS OF FOUR KINDS OF IMPLANTS IN

IMPLANT-BONE INTERFACE UNDER VERTICAL LOADING

Stress distributions in the buccilingual direction in implant-bone interface were

shown in figure 4. The highest stress concentration in the crestal bone surrounding

!O LU a.. .....J 2 -C/)

C/) :::)

C/) LU

w ~ cr: w ~ > C/) C/) .....J C/)

« w a.. cr: U a.. Z 2 0: 0 a.. U

36

18

0

18

36

18

0

18

36

18

0

18

36

18

TONGUE

/

o 6

84

CHEEK

~

18 12

LOCATION(MM)

24

I

[[

m

Fig. 5 Principal stress in buccilingual direction under horizontalloading .

the neck part of the implants was observed in type m (conical) implant. The princi­

par stresses in the buccal crestal bone region of the bone-implant interface were

O"max =-6.83MPa and O"min =-8.46MPa. and the principal stresses in the lingual

crestal bone regions were O"max =-5.70 MPa and O"min =-7.51 MPa. The type I

(modified Branemark) implant had the lowest principal stresses in crestal bone re-

'" c.. 2

(/)

(/)

L:..l n:: I-(.lI

--l <: Cl..

U z Cl:: a..

o o

TONGUE

85

CHEEK

A g~f-~--'-~~~~~~~--~~~-----o

o o

0 0

6

Wo -" 0

(/) 20;: >-

0 UJC

>c (/) (/) I..I.,;c

crC: a... .,... -6-u

o o

B

d

..... . c ~ 6

L).----CTlIlß

~..x

7+-----------~r-~--.---r-~---r--. 0.00 6.00 12.00 18.00 2 4 .00 30.00

LOCA C (mm)

Fig. 6 Principal stress of HA-TI implant in buccilingual diretion under

verticalloading in different implantation period of time.

gion of the bone-implant interface. For buccal side they were umax = -2.36 M Pa

and Umin = -5.77 M Pa. and for lingual side umax = -1.84 M Pa and Umin = -5.34

MPa . The stresses in crestal bone region in sequence of magnitude are type m > type II > type N > type I . For all cases umax and umin decreased in magnitude as

the distance from the loading site increased. There were clear stress peaks in the in­

terface between bone and the body of implant for type I and type II . The stress

86

distribution of type IV implant was the most uniform. The stresses of type I. II and

IV in the bone-implant interface surrounding the bottoms of the implants were

compressive stress. and for type mimplant C1max was a little tensile stress and

C1min was compressive stress.

The stress distributions of all the four kinds of implants in the medial-distal di.

rection were similar to that in buccilingual direction.

3. 2 THE STRESS DISTRIBUTIONS OF FOUR KINDS OF IMPLANTS IN

IMPLANT-BONE INTERFACE UNDER HORIZONTAL LOADING

The stress distributions in the buccilingual derection in implant-bone interface were

presented in figure 5. The stresses of all the implants in the forcing side (buccal side)

of the crestal bone regions were tensile stress. and the stresses in the opposite side

(lingual side) were compressive stress. Among the four kinds of implants. type I had the maximun tensile and compressive stresses in the crestal bone region. The

stresses in sequence of magnitude were arranged as type I > type m > type II > type IV.

3.3 THE STRESS DISTRIBUTIONS OF CYLINDRICAL HA COATED TITANIUM

IMPLANT UNDER VERTICAL LOADING

The stress distributions in the buccilingual direction at the four possible stages of

implant-bone interface were shown in figure 6. High stress concentrations were ob·

served at the crestal bone region. At stage A (implant-cancellous bone interface)

among the four stages. the stress in the crestal bone region was the highest and

gradually decreased with the increase of thickness of the lamina dura around the im·

plant. On the other hand. the stress in the interface surrounding the body part of the

implant was the smallest at stage A and gradually increased with the increase of the

thickness of the lamina dura. Therefore. the stress distribution in the interface be·

came more uniform as the implant-bone interface developed.

4. DISCUSSION

The results showed that compressive stresses were concentrated in various degrees

in the crestal bone regions for all the four kinds of implants under verticalloading . In

vivo and in clinical application. the saucerization of alveolar bone was often ob­

served. The saucerization may result from the stress concentration. Different implants

had different stress distributions. The highest stress was observed in the crestal bone

region surrounding the neck part of type mimplant. There were obvious stress peaks

87

in the implant-bone interface of bodies of type I and type n implants. This result

was in agreement with skalak's and Rieger's theoretical and stress analysis(4) (6) .

The qmax of type m and type N implants in the interface surrounding the body of im­

plants were very small and qminof both m and N implants were compressive stress.

Generally, the stress curve of type N implant in the implant-bone interface was the

most smooth, and the result indicated that the stress distribution of type N implant

was the most uniform.

This research also showed that when the horizontal loading was equal to the

vertical loading in magnitude, the stress in the crestal bone region under horizontal

loading was 4-5 times larger than that under vertical loading. In order to decrease

the possible damage to implant-bone interface, special attention must be paid to de­

creasing the horizontal component of the loading in clinical application. That we as­

sumed the horizontalloading was equal to the verticalloading was only for the con­

venience of comparison. In fact. the horizontal component of occlusal force during

function was usually much less than vertical component(6) .

The results of stress analysis showed that the cylindrical implant with smooth

surface was superior to the other three kinds of implants in the uniformity of stress

distribution. So the cylindrical implants was chosen as the main implants to be de­

veloped in our laboratory.

After the cylindrical HA coated titanium implant was implanted, the thickness of

lamina dura surrounding the implant would increased as time went on. At the same

time the stress distribution in the interface became more uniform. So a healing time

(about 3 months) after implanting was needed before the crown was fitted on the

crown post of the implant. Three months after implanting, the shear strength of im­

plant-bone interface was 10.68 M Pa according to animal experiment (refer to the

part n of this paper) and the highest shear stress Tmax in the bone-implant interface

was about 1.2 MPa (figure 6). This means that the stress during function was not

high enough to break down the interface.

REFERENCES

1. Weiss, C. M. The physiologie, anatomie and physical basis of oral endosseous

implant design, J .Oral implantol. 1982, 10, 459-486.

2. Hench, L. L. and wilson J. Surface active biomaterials, Science, 1984, 226.

630-636.

3. Suetsugu, T .• Stress analysis of blade implant. mechanical properties of implant

88

materials and stress distribution. J. orallmplantol1979. 8. 380-392.

4. Skalak. R. B .. Biomechanical condideration in osseointegrated prostheses.

J. Prosthet. Dent. 1983.49.843-848.

5. Rieger. M. R. . Bone stress distribution for three endosseous implants. J. Prothet.

Dent1989.61.223-228.

6. Brunski. J. B .• Biomaterial and biomechanics in dental implant design. Int. J. oral

maxillfac. implants 1988. 3. 85-97.

89

A STUDY ON HA-COATED TITANIUM DENTAL IMPLANTS PART 11 COATING

PROPERTIES IN VIVO, IMPLANT DESIGN AND CLINICAL EVALUATION (1)

JIYONG CHEN, JIMING ZHOU, XINGDONG ZHANG, DER I WAN Institute of Material Science and Technology Sichuan

University, Chengdu 610064, P.R. China SHAOAN WANG and ANYU CHEN

College of Stomatology, West China University of Medical Science

Chengdu 610041, P.R. China

ABSTRACT

HA coating was obtained by means of plasma spray technique. The properties of HA coating were studied by XRD, IR, SEM and mechanical tests. The coating may be pure HA, pure oxyapatite or a mixture of HA and oxyapatite by different spray-techniques. The porosity, surface roughness, SET surface are analysis and strength are raltive to the size of sprayed HA greatly. Animal experiments have demostrated that the HA-coated dental implants was a promising one because of i ts superior properties in biomechanical compatibility as weIl as biocompatibility. Having considered the properties of materials and the biomechanical requirements of dental implants sufficiently, a set of two-piece cylindrical HA-coated titanium dental implant were designed. Over 500 implants have been used in clinical application. A two-year follow-up study has shown that the resul t is satisfactory and the success rate is 98%.

(1) This project was supported by the National Natural Science Foundation of China and the National Educational Committee of China.

90

INTRODUC'l'ION

For the success of endosseous dental implants it is important to choose both appropriate implant materials and a good shape design of the implant. The inherent ability of metals to resist mechanical failure in biomaterial application has made the materials of choice for critical structures subject to high stress in service. titanium dental implants have been used clinically for more than 20 years. In order to increase the probability of rigid fixation through mechanical interlocking of bone wi th an implant surface, porous and rough surface endosseous dental implants have been developped /1-4/. However, the higher surface can result in a higher rate of metal ion release to host tissue. On the other hand, hydroxyapatite (HA) is of excellent biocompatibility and biological fixation and has no harmful metal ion release. But the clinical use in load-bearing cases has been limi ted by the mechanical properties of the material. Coating techniques opened a new possibili ty for implant application of HA /5/: HA coating on metals results in significant reduction of metal ion release and can make the implant surface bioactive. So HA-coated ti tanium implants make use of the advantages of both HA and ti tanium. In this paper, the properties of HA coating were studied by XRD, IR, SEM and mechanical tests. Animal experiments were carried out. Having sufficiently considered the properties of meterials and the biomechanical requirement of dental implants, a two-piece cylindrical HA-coated titanium dental implant was designed. Aseries of surgery tools was also developped. The dental implants were used clinically and the results were satisfactory.

THE PROPERTIES OF HA COATING ON TITANIUM

SPECIMEN PREPARATION The plasma spray technique was used for HA coating on titanium. The HA powder used in the study was prepared in authors' laboratory with the following reaction:

The Ca/P ratio of synthetic HA is 1.67, which means that the starting material is pure HA. After drying the HA powder was screened with sizing screens. Before spraying the surfaces of ti tanium, the implants were processed wi th grit blasting. A rough

91

titanium surface offers an effective means of increasing the adhesive strength between HA coating and titanium thanks to the higher surface area and mechanical interlocking.

STABILITY OF HA COATINGS The stabili ty of a HA coating after plasma spray is a considerable problem. Many investigators have reported their resul ts. E. Munting et al /6/ believed that the coating was constituted of hydroxyapati te and P. Ducheyne ed al /3/ reported that significant changes of the chemical and physical characteristics were observed as a result of the plasma spray process and HA was transformed into a mixture containing in addition tetracalcium phosphate and <l and ß-tricalcium phosphate (TCP) • For investigating this problem, two specimens were prepared. Before spraying, specimen A was treated below 700°C and specimen B was sintered at above 800°C. After spraying, both specimen A and B were tested by XRD and IR. Fig. 1 and 2 show the XRD and IR spectra be fore and after spraying. The XRD spectra of specimen A and Bare similar.

10 20 30

2(J -

40 5U

Fig. 1 XRD spectra of HA coatings. a) Specimen A be fore Spraying; b) Specimen B before spraying; c) Specimen A after spraying and d) Specimen B after spraying.

92

The notable difference is that the crystallization of specimen B is better than that of specimen A before spraying and there is very li ttle of spectra moving at light 29 angle after spraying. There is no a orß TCP before and after Spraying. The IR spectra of specimens A and Bare qui te different after spraying. Specimen A keeps its OH- group and specimen B loses its OH- completely . The authors have reported /7/8/ that HA lost its OH- group and became oxyapati te. The two materials had the same space structure and had almost the same XRD pattern. The XRD method could not determine the loss of the OH- group. So from Figure 1. we cannot know confidently whether the material is HA or oxyapatite. The IR spectrum was very sens i ti ve in determining the OH- group. Figure 2 shows that a different preparation of the HA powder be fore spraying can obtain a different coating. The authors also reported /7/8/ that HA is stable below a 1200 °C temperature and HA cannot be transformed into a or ß-TCP below this temperature. The result of the fact that the HA coating contains a or ß-TCP after spraying implies that the CA / P ratio of the green HA powder is lower than 1.67, which is the value of pure HA. But the a or ß-TCP in HA also cannot be determined by XRD be fore HA is treated beyond 850 °C. In this experiment, we used pure HA, so a or ß-TCP could not be found after HA was sprayed.

o

c

B

A

4000 2600 1666 1000 400 WAVENUMBER cm- 1

Fig. 2 IR spectra of HA coatings. a) Specimen A be fore spraying; b) Specimen B be fore spraying; c) Specimen A after spraying and d) Specimen B after spraying.

93

From the above discussion we know that plasma-spray coating may be applied to HA, oxyapatite or the mixture of the two materials according to the different methods of preparation of the starting HA powder. If the starting material is treated below 700 °C the HO- group may be conserved, and if the treatment temperature is higher, the OH- group may be lost in the plasma-spray process. If the Ca / P ratio of the green material is lower than 1.67, the coating may contain a or ß-TCP.

THE PROPERTIES OF HA COATING In this section the authors investigated in detail the relationship between different-sized HA powders in terms of porosi ty, surface roughness, BET surface area and adhesion strength of the coating. Three types of HA powders sized 60-751J1l1, 40-451J1l1 and less than 301J1l1 were used to make HA coatings. Aseries of factors were determined by the following methods. Apparent densi ty: by measuring the volume of coatings and weighing them.

Fig. 3 SEM micrographs of HA coating. a) Surface structure; b) cross section.

94

TABLE 1

The relationship of different sized spray powders and coating

properties.

HA partic1e size <30 40-45 60-75

(IJIII )

Apparent density 2.63 1.96 1.80

(g/cm3 )

Porosity 17 38 43

(t)

Surface roughness 5-10 20-30 30-40

( IJIII )

BET surface area -- 0.78 --(M2/g)

Tensile strength 14 17 19

(MPa)

Shear strength -- 9.17 --(MPa)

Porosity: comparing the theory dens~ty 3.156g/cm3 w~th apparent

density.

P . 3.156- apparent densityx 10001. oroslty - 3.156 '"

Surface roughness: by using SEM cross section micrography and

measuring the difference between the highest and the lowest.

Tensile strength: mechanical test meter, 5 specimens per group,

speed: 2mm/min.

Shear strength: mechanical test meter, 5 specimens per group,

speed: 2mm/min.

Specific surface area: BET surface area meter.

From the results we conclude that the properties of coatings

relate closely to the size of the sprayed HA powder, and the

larger the size of the HA power, the higher the porosi ty, surface

roughness and adhesive strength. Sizes of HA particles and

95

preprocessing may significantly influence the implant properties

in terms of biology and mechanics. So the size of sprayed HA

powder should not be too fine to obtain a good coating.

ANIMAL EXPERIMENT

A total of ninety implants were used in this investigation. The implants were 7.0 mm in length. In order to avoid mechanical retentions during the push-out test, all the implants were made in a slightly conical shape with a top of 4.00 mm and a bot tom of 3.0 mm in diameter. Twelve endosseous dental implants, including six cylindrical HA-coated ones and six non-coated ones, were prepared. They were 10.0 mm long and 4.0 mm in diameter. The plasma spray coatings were about 100 ~ thick. The implants were sterilized prior to implantation.

Fig. 4 Photograph of interface of coating and bone tissue 12 weeks of post implantation.

Six adult dogs and two monkeys were used as animal models. Three months after the extractions, the dental and the cone implants were implanted into the mandibles and the femora of the animals respectively. The animals were sacrificed 12 and 24 weeks after implantation. The samples were prepared for examinations according to requ!rements. 12 weeks after implantation, the values of mean shear strength were 11.74 MPa (HA-coated), 10.68 MPa (sintered HA), and 2.68 MPa (uncoated titanium). 24 weeks after implantation, the mean shear strengths were 12.34 MPa (Ha-coated), 11.52 MPa (sintered HA) and 3.14 MPa (uncoated titanium). Based on the statistical

96

analysis, there was no significant difference for the HA-coated and the HA implants, while the uncoated implants showed a significant difference compared wi th the others. The newly formed bone close to the implants showed a progressive remodeling. Some HA grains were observed at the implant-bone interfaces both in the HA and the Ha-coated implants, but no obvious difference was detected. The uncoated titanium implant was surrounded by a thin layer of fibrous tissue. The Ca/P ratio of the bone around the uncoated titanium implant is lower than around the bone surrounding the HA or the Ha-coated implant. The results of this research have demonsrated that the Ha-coated dental implant is a promising one because of i ts superior properties of biomechanical compatibility as weIl as biocompatibility.

DESIGN OF HA-COATED TITANIUM DENTAL IMPLANTS

A kind of HA-coated titanium dental implant has been developed in our institute on the basis of three dimensional finite element stress analysis of endosseous dental implants, of the properties of HA coatings, and of the results of animal experiments. The implant consists of a root and a crown post. The root is a cylinder with a semispherical bottom. The crown post can be screwed in and fixed to the root. For convenience of implantation, a set of special tools, which included drills with different diameters, a depth detector rod and a screw driver has been designed. The implants root are 10.5 or 12 mm in length and 3.3, 3.5, 3.7, 4.0 and 4.2 mm in diameter. The crown posts are 6 or 8 mm in length and have the same size as the roots in diameter. The drills are separately 1.5, 3.2, 3.4, 3.6, 3.9 and 4.1 mm in diameter. The surface of the implant made from bioinert material is surrounded by fibrous tissue and the main stress concentrate on the neck and bottom of the implant. In the osteointegration implant, the stress concentrats mainly on the neck and only a little on the bottom. According to the results of part I, the cylindrical implant has minimum neck stress and i ts stress distribution is even. So the basic configuration of our implant is cylindrical. In the initial stage after implantation the neck stress is maximum. This stress may induce bone absorbtion and implant loosening. To guarantee a stable implant and bone interface, a two-piece structure is selected so that no stress

97

occurs at the interface after implantation. Two or three months later, dura grows to about 0.5 mm, the implant is fixed and the neck stress reduced. Then the crown post can be screwed and the dental prosthesis can be carried out. Plasma-spray HA coating should be prepared according to the following principles: a) It must be thick enough to cover the ti tanium core surface completely to prevent release of metal ions. b) The coating should not be too thick. The effects of coating are: covering the Ti surface, making the surface bioactive and conducting strees to the Ti core. If coating is too thick, it may be brittle, which is a common property of block ceramics. c) The adhensive strength at the interface between coating and titanium must be strong enough to resist tensile and shear stress. d) It should have suitable roughness and porosi ty so that new bone can ingrow into the coating and increase mechanical inter lock and the combined area between HA and bone tissue.

Fig. 5 HA-coated Ti dental implants and their surgery tools.

98

CLINICAL EVALUATION

c

B

A

Fig. 6 Clinical application. a) Implanting; b) 12 weeks after implantation and be fore screwing the crown post and c) after the prosthesis.

99

Over 500 such dental implants have been used clinically. By way of example, Figure 6(a) shows the implant being put into the maxilla of a young patient. Figure 6(b) is a photograph taken three months after implantation, be fore screwing the crown post. Figure 6(c) is after the prosthesis. The clinical results are promising after two years' follow-up, and the success rate is 98%.

SUMMARY

Stress distribution was carried out on four kinds of different

optimal configuration implant by three dimensional finite element

stress analysis. The results show that in the osteointegration

implant the stress distribution in the cylinder is evenest. The

HA coating may be pure HA, oxyapatite or a mixture of the two

by different plasma spray processes • Proper techniques can control

the resolution of HA. The adhesive strength, porosity, surface

roughness and BET surface area are closely connected with the

size of sprayed HA. Animal experiments show that HA-Ti implants

have promising biocompatibili ty and biomechanical property. After

sufficient evaluation of the properties of materials and the

biomechanical requirements of dental implant, a two-piece cylindrical HA-Ti dental implant and the relative surgery tools

were developed. Over 500 implants have been used in clinical and

the results are satisfactory. The success rate is 98% after two

years' fOllow-up.

REFERENCE

/1/ R.M. Pilliar, Oral Implantology and Biomaterials, edited by H. Kamahara, Elsevier Science Publishers B.V., Amsterdam, 1989, P. 151.

/2/ P. Ducheyne, In Vivo Metal Ion Release from Porous Titanium-Fibre Material, J. Biomed. Mater. Res. 18 (1984) 293.

/3/ P. Ducheyne, J.M. Cuckler, S. Radin, K.E. Healy, E. Nazar, Bioactive Calcium Phosphate Ceramic Linings on Porous Metal Coatings for Bone Ingrowth, The Third World Biomaterials Congress, April 21-25, 1988, Kyoto, Japan, P. 309.

100

/4/ R.M. Pi11iar, Powder Meta1-Made Orthopaedic Imp1ants with Porous Surface for Fixation by Tissue Ingrowth, C1in. Orthop. and Re1. Res. 176 (1983) 42.

/5/ K. De Groot, Hydroxyapatite Coatings for Imp1ants in Surgery, High Tech Ceramics, edi ted by P. Vincenzini, E1sevier Science Pub1ishers B.V., Amsterdam, 1987.

/6/ E. Munting, M. Verhe1pen, Li Feng, Morpho1ogical and Biomechanical Study of the Bone-Implant Interface of Hydroxyapati te Coated Implants. The Third Wor1d Biomateria1s Congress, 1988, Kyoto, Japan, P. 308.

/7/ Jiming Zhou, Jiyong Chen, Xingdong zhang, Pin Zhou, A Study of Tricalcium Phosphate in Hydroxyapatite. Proceedings of the C-MRS International '90 Ed. by H. Li, Elsevier Science Publishers Ltd., 1990, The Netherlands.

/8/ Jiming Zhou, xingdong Zhang, Jiyong Chen, High Temperature Characteristics of Synthetic Hydroxyapatite, Biomaterials, to be published.

101

REHABILITATION OF RADICAL MASTOIDECTOMY CAVITIES

WITH CALCIUM PHOSPHATE CERAMICS

Massimo DeI Bo - Arturo Zaghis

Istituto di Audiologia dell'Universita di Milano.

Ospedale Policlinico, Via F. Sforza, 35 MILANO

ABSTRACT

The presence of keratinizing stratified squamous epithelium within the middle ear or other pneumatized portions of the temporal bone is called a cholesteatoma. The surgery of cho­lesteatoma frequently results in the ereation of a radiealma­stoidectomy, an operation in which the mastoid cavity, epitym­panum, middle ear and external auditiry eanal are eonverted into a eommon eavity exteriorized throughthe external meatus. The problems of the open mastoid eavity are familiar to all otologists. Reeurrent infection, water avoidance and recurrent visits eleaning, are the indications to the rehabilitation of a radieal cavity. This artiele deseribes our method for the anatomieal revalidation of radical eavity: partial obliteration of the eavity and posterior canal wall reconstruction with tri­calcium phosphate ceramies and fibrin glue.

INTRODUCTION

Aural cholesteatoma is a cystie strueture produced by kerati­nizing squamous epithelium within the middle ear or other

102

pneumatized portions of the temporal bone; in other words, as Gray says, "skin in the wrong place". Surgical treatment in middle ear cholesteatoma, is the solution of choice. Radical mastoidectomy, when properly performed, has proved to be an efficient surgical procedure for treatment of cholesteatoma, although to-day its number is getting smaller. The technique consists of joining the tympanie cavity and the mastoid into one single cavity by removing the posterior canal wall. However, sometimes there are complications from the old mastoid cavities:

- recurrent otorrhea and infection; - water avoidance; - recurrent visits for cleaning and topical drug therapy; - hearing aid fitting.

To prevent such complications, reconstruction and obliteration of the mastoid cavity, where indicated, seems to be a good so­lution. Hearing improvement is desirable, but it's not the first objective.

This article describes our method for anatomical revalidation (reconstruction and obliteration) of radical cavity with calcium phosphate and fibrin glue. Until now, the most common material seems to be autograft muscle flaps, often in combination with bone chips or bone pate. Homograft museie and bone have also been used. But free museie or fat generally dissolves by suppu­ration. Bone pate mixed with fibrin glue does not warrant squa­mous epithelium or cholesteatoma being left behind. Calcium phosphates are the inorganic component of bone and teeth, and their biocompatibility in vive and in vitro is well esta­blished. They are conductors of osteogenesis and become perfectly integrated with the surrounding bone. (1-2-3).

MATERIAL

A two-stage procedure is usually needed for anatomical and functional restoration, with the second procedure coming about one year after the primary operation. We use a mixture (1:1) of two types of calcium phosphates:

- Hydroxylapatites granules. Ca10 (P04 )6 OH2 Its composition is like bone; lt is not resorbed and when porous acts as a scaffold for bone healing.

- Tricalcium phosphates. Ca3 (P04 )2

103

It is bioresorbed and its degradation products can be re­constituted by the body to form new bone material.

SURGICAL TECHNIQUE

The patient is supine under general anesthesia. Infiltration of the soft tissues is made behind the auricle and in the external auditory canal. The surgeon operates seated, and light is provided by the microscope raised up. Incision is sited in the supra - and postaural sulcus from its anterior extremity (above the root of the helix) to a level corresponding to the floor of the external auditory meatus. Hemostasis is then se­eured by diathermy. The surgeon then uses an elevator to de­tach the soft tissues from the bone. The skin is then carefully separed from the walls of the cavi­ty, in such a way as to obtein a posterior meatal flap which is folded upward. The remaining part of the skin of the eanal, lining the anterior wall, is elevated superiorly. A anterior peduneulated tragal flap is thus obtained. (Fig. 1 A-B). Cleaning of the tympanic cavity is performed and after removal of all ectopic epithelium the mastoid cavity is revised with the cutting burr. At this point the patency of the Eustachian tube is verified using catheterism and washing. After this eu­retive step, when a seeond stage proeedure is suggested, we introduee a thiek Silastic sheeting whieh cover the medial wall of the eavity, extending both into the Eustachian tube and through the aditus area into the mastoid eavity. Another Lio­dura sheet is preshaped in the curve of the posterior eanal wall, perpendieularly to the Silastic; it serves as a retainer for the calcium phosphate granules. (Fig. 1 cl. Next, the mastoid cavity is filled with the calcium phosphate granules ineorporated with the fibrin adhesive (Tissueol). The obtained compound is malleable and ean be used to fill the majority of the cavity. Tissucol is an ideal binding material, attaching the granules to one another and to the cavity walls. (Fig. 1 D). Fresh temporalis faseia is cut out and accurately positioned behind the tympanomeatal flaps. (Fig. 1 E). Faseia is, in our hands, the best tissue for reconstructing the tympanie membrane, being both supple and elastic. The pre­viously elevated anterior and posterior flaps are then repla­ced, eovering apart of the new drum. (Fig. 1 F). The operation is now eompleted having closed the retroauricular approach and having paeked the new auditory eanal with Gelfoam. With this technique, a single material ean be used both to fill the ea­vity and to reconstruct the posterior wall.

Fig. 1

104

A B

c o

E F

Surgical technique for the rehabilitation of radi­cal mastoidectomy cavities with calcium phosphate ceramies.

105

RESULTS

This procedure has now performed in 8 cases, the median lenght of follow-up was 24 months. There were 3 female and 5 male patients. The youngest patient was 32 years old and the eldest 56 years old. In two cases the trans plan ted temporalis fascia vanished due to infection, despite postoperative antibiotic treatment, and in one case the granula had to be removed individually from the discharging cavity. Six patients had no problems in the postoperative control period. There were no extrusion problems with the implants. The clinical tolerance has been very satisfying and the ma­terial remains stable, does not retract and does not show any inflammatory reaction. (Fig. 2). This paper is not intended to study the hearing results of our surgical treatment; hearing improvement is the objective of the second look.

Fig. 2 Computed Tomography of the left temporal bone, 6 monthsafter operation. Excellent stability and biocompatibility of the granules, without any foreign body reaction.

1~

CONCLUSIONS

Certain salient features stand out from our work: - Calcium phosphate granules proved their biocompatibility

in the ear; - They do not shrink; - They prevent recurrent cholesteatoma in the mastoid

cavity; - They are very suitable for reconstruction of the posterior

ear canal wall; - The aim of preventing recurrent infections was achieved by

obliteration with calcium phosphate ceramics.

We insist on the necessity that a long fOllow-up is essential to assess any results, but the initial results have been very satisfactory enough to continue practicing the ceramic implants.

ACKNOWLEDGEMENTS It is a pleasant duty to thank the "Poliambulatorio Odontoia­trico Venini" and "Ditta Dispo" who together provided both the calcium phosphate ceramics and the assistence in our first steps.

REFERENCES

1. Daculsi G. : Biological properties of calcium phosphate biomaterials. In Babighian-Veldman (Eds) Transplants and Implants in Otology. Kugler-Ghedini Pub. Milano, 1988.

2. Denissen H. - Mangano C. - Venini G. : L'idrossilapatite in implantologia. Piccin Nuova Libreria, Padova, 1985.

3. Hormann K. : Mastoid obliteration. In Tos-Thomsen-Peitersen (Eds) Cholesteatoma and mastpoid surgery. Kugler-Ghedini Pub. Milano, 1989.

4. Gersdorff M. - Robillard T.A.J. : Revalidation des cavites d'evidement par reconstruction osseuse and bio- materiaux. Acta Oto-Rhino-Laryngol. Belg. 1988, 42, 35-39.

5. Olaizola F. - Arroyo R. : Results obtained in the recon­struction of the temporal bone with biocompatible material after 18 months. In Tos-Thomsen-Peitersen (Eds) Cholesteatoma and mastoid surgery. Milano, 1989.

Kugler-Ghedini Pub.

107

CLINICAL RESULTS OF IMZ DENTAL IMPLANTS

P. PASSI University of Padua, Dental Clinic

Head: Prof. A. BeItrame

Abstract

In the last four years, the author inserted over one hundred IMZ cylinder dental implants. They were either hydroxylapatite or pure ~itanium-coated. These implants are of the "osseointegrated" type, needing a two-phase technique. The c1inical aim was to replace single teeth, to build olTerdentures or partial and full-arch fixed reconstructions. Prosthetic therapy was carried out by the same person who surgically placed the implants. The original prosthetic method, wh ich utilizes a shock-absorbing cylinder, was replaced by a cast pin abutment in most patients. Clinical results were very good, because more than 90% of the implants are still functioning well, with minimal bone loss.

Introduction

IMZ (Intra Mobile Zylinder) dental implants belong to the so-ealled "osseointegrated" category, that today are believed to be the most reliable from a clinical point of view [1,2]. They are eylindrical, and eoated by plasma-spray technique [3] either with pure titanium or hydroxylapatite (HA). The shape and surgical teehnique of both are identical. They are available in four lengths (8, 10, 13, and 15 mm), and in two diameters (4 and 3.3 mm). They are utilized with a two-phase technique [4,5]. Insertion is carried out raising a surgieal flap, the implants are placed after drilling the bone with a set of specific burs, then the tissues are sutured over the implant heads. After a minimum of three months, the tissues are surgically re-opened, and the prosthetic reconstruetion is performed by means of several abutments, depending on the ease. The original and typical design of IMZ implants has an internal shoek­absorber eylinder. in order to give the im plant a resilienee like that of a natural tooth. This eylinder is made with polyoxymethylene (POMI, and needs to be replaeed at least onee a year.

108

However, it is also possible to carry out the prosthetic reconstruction by means of stiff abutments. In this ",ay, the need to replace the shock-absorber cylinder is avoided. This paper deals with 1061HZ implants placed in foul' years. This is not a large number, but all these im1Jlants were placed; prustheticaily reconstructed , and carefully controlled over aperiod of time b~· one person only (the author himselfl.

Clinical experience

In the period April 1987-!>Ial'ch 1991, 106 implants were inserted in 41 patients. They were placed in order to build overdentures, or fixed bridge abutments (Fig.l), or for single tooth re placement. Three fulI-arch implant fixed reconstructions were carried out, two mandibular and one maxillal'Y (Fig.2). The types of prosthetic reconstruction are sho,,-n in table 1.

TABLE I

Prosthetic reconstructions built on IMZ implants

Overdentures 13 Fixed bridges 28 Full-arch fixed reconstructions 3 Single tooth replacements 3

Total 47

The total of reconstructions is not equal to the number of patients treated (41), because some patients had more than one prosthesis (for instance, a fixed bridge and an overdenture), and five patients with 14 implants are still in the heaIing phase. The number of implants placed per year is shown in table H, where failures are also reported. "Early" failures occurred in the heaIing phase (first three months), while "late" failures occurred after the implants were prosthetically loaded.

TABLE II

Implants placed per year and failures

Year Implants inserted Failures Early Late

1987 10 2 1988 21 1 2 1989 20 1990 41 4 1991 14

Total 106 9 (8.49%)

109

TL lS note,,-Tol'Lhy tnat. onlr t\'\-o failures occurred after the healing l:>hase; Z1!1d ,-.. el'e dul.:! to inlplant fracture.. One probably occurred for exces~-3i\-E' lC':J.d, 1)f~cause a single irnplant supported a maxillary ü\'~~l·d(_'nturf.? in a patient \,Ohere the extreme bone thinness made it possibh-: Lo place only one iInplant in the front alyeolar ridge. Breakage of the other implant. occurl'ed because of manoeuvres for !'emoval of a bl'üken PO!'{ snack-absorber. In the seven railures that occurred in the early healing phase, the implants became loose and came trough the mucosa; with minimal pain ror tht~ patient. In these cases; lhe loosened implant is easily removed, and it is generallJ" possible to insert another one, near the rormer 01'

even in the same place, arter several weeks. Most implants were of the titanium-coated type, and only ten were hydl'oxylapatite-coated. No railure occurred in this latter group. The clinical conditions of the implants that passed the healing phase are very good. There is minimal 01' no bone loss, and the gingival condition is excellent (Fig.31; this is also true for the implants placed four years ago (Fig.4 I. At this time, none of the implants that were prosthetically reconstructed show clinical 01' radiological signs of possible failure.

Discussion and conclusions.

The 1M2 implant system is easy to handle and clinically very reliable [6J. Its only weak point may be the POM shock-absorber. In fact, this may fracture, and tends to accumulate bacterial plaque and calculus and becomes malodorous with time. For these reasons, its periodic replacement is mandatory. Considering that equally good results are obtained with similaI' implant systems that utilize stiff abutments, in the last two years the POM cylinders have been replaced with cast metal pins, 01' with titanium screw cylinders. This latte l' solution is particularly fit in cases where the difference of inclination of the implants exceeds 15 degrees, 01' in order to build overdentures in patients where the lack of vertical space prevents performing telescopie crowns. As far as the implant coating is concerned, several investigations underlined the so-called "biointegration" of hydroxylapatite [7], whieh appears to bond with bone more tightly than titanium [8J. For this reason, hydroxylapatite might be the best material for endosseous implant coating. However, clinical results obtained with titanium-coated implants are very similaI' [9J. The bettel' results obtained in this clinical study with HA-coated implants are to be considered as not significant, owing to the small number of these implants inserted. In reconstructing totally edentulous jaws with fixed bridges, bone resorption may make it necessary to build very long prosthetic crowns with pOOl' aesthetics, 01' to provide an acrylic artificial gingiva at the base of the bridge. In such a situation, a removable overdenture may be the best solution in terms of good hygiene, aesthetics and low cost. On the basis of this clinical experience, the IMZ implant system seems to be very reliable. These clinical results confirm those of wider clinical studies on the same implants [9,10J and others of the oseointegrated type [11,12,13,14J , where the overall success rate is over 90%, even after five years and more.

110

Fig.l. Two distal maxillary edentulous spans reeonstrueted with lour implants as abutments. One-year radiographie eontrol.

Fig.2 A lull-areh lixed maxillary reeonstruetion alter 18 months. It was earried out by means 01 live implants.

111

Fig.3. Good clinical condition around an implant inserted two years before, reconstructed by means of a cast metal pin. It sets as distal abutment for a fbced bridge.

Fig.4. Implants inserted four years ago, showing no bone resorption. They are connected to a Dolder's bar with titanium screw abutments, and support an overdenture.

112

References

1. Albrektsson T., Lekholm U.: Osseointegration:current state of art. Dent. Cl. N.Am., 33(4):537, 1989

2. Passi P., Terribile Wiel Marin V., Miotti A. Histologie investigation on two titanium screw dental implants in humans. Quintessence Int., 20(6):429, 1989

3.Babbush C.A.:Titanium plasma spray screw implant system for reconstruction of the edentulous mandible. Dent. Cl. N. Am.,30(1):117,1986

4. Kirsch A., Mentag P.J.:The IMZ endosseous two phase implant system.A complete oral rehabilitation treatment concept. Oral Implantology, vol.XII n.4, 1986

5.Malevez C.,Daelemans P.,Kovacs B.,Kurilalo A.: Les implants IMZ de Kirsch. Actualites Odonto-Stomatologiques, 159:539, 1987

6.Passi P., De Polo G., Ferronato G., Miotti A.:Gli impianti IMZ neUa pratiea cliniea. Il Dentista Moderno, n.1:163, 1991

7.Passi P., Terribile Wiel Marin V., Parenti A., Miotti A. Ultrastructural findings on the interface between hydroxylapatite and oral tissues. Quintessence Int.,22 n. 3:193, 1991 8.Meffert R.M., Block M.S., Kent J.N.,: What is osseointegration?

Int. J. Period. Rest. Dent., 4:9, 1987 9.Kirsch A., Ackermann K.L.: The IMZ osseointegrated implant system.

Dent. Cl. N. Am., 33(4):733, 1989 10. Koch W.: Die Zweiphasige enossale Implantation von intramobilen

Zylinder-implantaten IMZ. Quintessenz, 1:24, 1986 11. Adell R.:Clinieal results of osseointegrated implants supporting

fixed prostheses in edentulous jaws. J. Prosthet. Dent.,50:251,1983 12. Cordioli G.P., Beltrame A., Favero G.A.: Studio clinico longitudinale

su protesi osteointegrate secondo Bränemark. Il Dentista Moderno, n.4:909, 1989

13.Golec T.S.:Three year clinical review of HA coated titanium cylinder implants. Oral Implant., XIV (4), 1988

14.Schroeder A., Suffer F., Krekeler G.: Orale Implantologie. G. Thieme Verlag, Stuttgart, New York 1988

.4uthor's address: Dr. Piero Passi, Via C. Battisti 3, 35100 Padova, ltaly

113

DIRECT COMPOSITE-CERAMIC RESTORATIONS: A CLINICAL STUDY

CARLO PRATI, GIANNI MONTANARI, EUGENIO TOSCID, ANTONIO SA VINO. School of Dentistry, Department of Operative Dentistry

University of Bologna, ltaly Via San Vitale 59, 40125 Bologna, ltaly

ABSTRACT

A clinical study was carried out to evaluate the conditions of Class land Class 11 composite resin-ceramic restorations. A follow up of one, two and three years was made to evaluate the anatomical form, marginal discoloration, interproximal contact and secondary caries of 224 restorations placed in 53 healthy subjects. The USPHS system was used to score the clinical evaluations. Four different composite resins were used: P 50 (a small particle zirconia filled BIS GMA composite resin bonded ceramic), P 30 (a small particle BIS-GMA composite resin bonded ceramic), Clearfil PhotoPosterior (a fumed silica particle BIS-GMA, TEG-GMA-UTMA composite resin) and HELIOMOLAR (a silica microfilled BIS-GMA/poly-urethane composite). After three years, all the restorations showed several types of alterations, but only 6% of sampies were replaced for fractures and other aIterations.

INTRODUCTION

In recent years, several types of composite resin-ceramic restorations have been developed due to the great attention to aesthetics in anterior and posterior teeth. Several clinical studies have suggested that occlusal wear and marginal deteriorations were the most important problems connected with composite resins (1). These problems have been weIl documented in posterior teeth (1, 2, 3). Two types of light­cured Resin Bonded Ceramic with Advanced Particle Coupling (APC) (P 30 and P 50, 3M Co., MN, USA) have been recently marketed The aim of this study was to evaluate the clinical performances of P 50 and P 30 with two other composite resins: Clearfil PhotoPosterior (Kuraray, Japan) and Heliomolar (VIVODENT/lVOCLAR, Liechtenstein).

MATERIALS AND METHODS

Fifty-three subjects with similar oral conditions were selected for the study. Each patient had at least two or three old restorations or caries replaced with composite resin materials.

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The materials selected were: P 30 a light-eured resin bonded eeramie with advaneed partiele eoupling (APC) N=50.( 3M Co. MN, USA), P 50, a small partiele zireonia filled Bis-Gma eomposite resin bonded ceramie N=73 (3M, Co. MN, USA), Clearfil PhotoPosterior a fumed siliea partiele BIS-GMA, TEGMA-UTMA eomposite resin N=60 (Kuraray, Japan) and Heliomolar, a siliea mierofilled BIS-GMA/Poly­urethane eomposite N=40 (Vivadent/Ivoelar, Leiehtenstein) (see Table 1). The eavities were prepared aeeording to eonservative proeedures. After 6 months, 1,2 and 3 years eaeh restorations was evaluated aeeording to the USPHS system (Table 2).

RESULTS

After 6 months no restorations were replaeed or were rated as Charlie. After 1 year only, P 30 presented one restoration lost due to a rapid oeelusal wear. After two and three years eaeh group presented same sampies with elinieally detectable alterations. Oeelusal wear or marginal Craetures was the most eommon damage observed in the restorations. No statistieal differenees were observed among the Cour groups.

Material

P30 P50 Clearfil PhotoPosterior Heliomolar

TABLEI Materials used in the study

ManuCaeturer

3M Co, MN, USA 3M Co, MN, USA Kuraray, Japan VIV ADENTßVOCLAR, Leiehteinstein

DISCUSSION

The present data eonfirm that oeelusal wear is the most important problem. It is possible that atrition and abrasion oC tooth surfaee ean eause the modifieation of tooth surfaee with small fraetures and eraeks along the interface between inorganie and polymerie phase (4). The Craetures along this surfaee may be modified by hydrolitie degradation with the release oC inorganie (eeramie) eomponents in the oral fluids (5). In Caet, the surface roughtes oC our sampies inereased during that time. An in vivo study has shown that wear rate oC Heliomolar and P 50 was less than 12 mieron during a one-year CoIlow-up (6). This study eorrelates weIl with our data on the same materials. Our study on P 10 and Sitar reported simitar elinieally deteetable alterations. This means that the present materials may be earefully used when a large oeelusal surfaee is exposed to oral environment. As regards marginal adaptation, the laek oC problems suggested that the dentinal bonding agents used in this study are able to prevent mieroleakage and marginal de­bonding. Caries resistanee and marginal diseoloration may be related to the good marginal adaptation oC restorations. Several in vitro studies evaluated mieroleakage (8, 9) and the Craeture resistance of teeth restored with eomposite resins (10, 11), suggesting that light-cured resin bonded ceramics have properties similar or better than the other composite resins. However, the lack oC a elear relationship between in vitro and in vivo studies suggests that other clinieal studies are needed for information on the long-term life-survival of these materials.

Anatomie form

ALFA

BRAVO

CHARLIE

115

TABLEl USPHS rating system as used in this study

The restoration is continuous with existing anatomie form

The restoration is discontinuous with the existing anatomie form, but the missing material is not sutlicient to expose dentin or base

Sumcient material lost to expose dentine or base

Marginal adaptation

ALFA

BRAVO

CHARLIE

No visible evidence of a crevice a10ng the margin into which the explorer will penetrate.

Visible evidence of a crevice along the margin into which the explorer will penetrate or catch

Explorer penetrates into a crevice and dentin or base is exposed.

Marginal discoloration

ALFA

BRAVO

CHARLIE

Color match

ALFA

BRAVO

CHARLIE

No discoloration anywhere on the margin between the restoration and the tooth structure.

The discoloration has not penetrated along the margin in a pulpal direction.

The discoloration has penetrated a10ng the margin in a pulpal direction.

The restoration matches adjacent tooth structure in color, shade and translucency.

The mismatch in color, shade or translucency is with­in the acceptable range of adjacent tooth structure.

The mismatch in color, shade or translucency is out­side the acceptable range of adjacent tooth structure.

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Interproximal contact

ALFA

BRAVO

CHARLIE

Contact is tight and it is difficult to pass dental floss between the restoration and the adjacent 1ooth.

Contact is light and it is relatively easy 10 pass dental floss between the restoration and the adjacent tooth.

There is no contact between the restoration and the adjacent tooth.

Secondary caries

ALFA

BRAVO

Materials

P 30 (N=50)

P 50 (N=73)

ClearfIlI Photo Post. (N=60)

Heliomolar (N=40)

No caries present.

Caries present.

TABLE3 Clinical results.

6m 12m

ALFA 49 45 BRAVO 1 4 CHARLIE 0 1

ALFA 70 70 BRAVO 3 3 CHARLIE 0 0

ALFA 60 60 BRAVO 0 0 CHARLIE 0 0

ALFA 40 36 BRAVO 0 4 CHARLIE 0 0

REFERENCES

24m 36m

40 34 9 10 1 6

68 60 3 1 2 9

57 51 2 8 1 1

34 33 4 4 2 3

1. Warren, J.A. and Clarkg, N.P., Posterior composite re4sin: current trends in restorative techniques. General Dentism, 1987,34,368.

117

2. Leinfelder, K.F.,Wilder, A.D. and Teixeira, L.D., Wear rate of posterior composite resin • .I.A.D.A, 1986, 112, 829-833.

3. Leinfelder, K.F., Composite resins in posterior teeth. Deut CUn North Am., 1981, 25,357-364.

4. Söderholm, J.M., Influence of silane treatmetn and liller fraction on thermal expansion of composite resins • .I Deut Res, 1984, 63, 1321-1326.

5. Kalacaandra, S., Influence of lillers on the water sorption of cOinposites. Deni MakI, 1989 5, 283-288.

6. Leinfelder, K.F., Posterior composite resins • .IA.DA" 1988, 114, 21-26.

7. Prati, C. and Montanari, G., Three-year clinical study oftwo composite resins and one non-gamma 2 conventional amalgam in posterior teeth. Schweizer Monastsschrif tür Zahanmedizil4 1988, 98, 120-125.

8. Darbyshire, P.A., Messer, L.B. and Douglas, W A., Microleakage in dass 11 composite restorations bonded to dentine using thermal and load cycling. J...Dent Res.., 1988, 67, 585-587.

9. Prati, C., Nucci, C., Davidson, L.C. and Montanari, G., Early marginalleakage and shear bond strength of adhesive restorative systems. Deut Mater., 1990, 6, 195-200.

10. Causton, B.F., Miller, B. and Sefton, S., The deformations of cusps by bonded posterior composite restorations: an in vitro study. Br Dent .I., 1985, 159,397.

11. Reel, D.C. and Mitcaell, R,J., Fracture resistance ofteeth restored with dass 11 composite restorations . .1 Prosthet Dent., 1989, 61, 177.

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ARTICULATION OF CERAMIC SURFACES AGAINST POLYETHYLENE

R. M. STREICHER, M. SEMLITSCH, R. SCHÖN SULZER Medical Technology Ud. CH-8401 Winterthur, Switzerland

ABSTRACT

Investigation was carried out on five grades of the oxide-ceramics alumina and zirconia in a pin-on-disc test against polyethylene for suitability as articulating components for total joint prostheses. The tests showed a difference in surface quality and energy between the various grades of ceramic. The polyethylene wear rate caused by the ceramic counterfaces was with the alumina qualities lowest and less than in combination with CoCrMo-alloy. The alternative Zrl>2 ceramics yielded unfavourable wear and friction results. Although mechanically preferable the tribological investigation seems to caution the use of high performance non­alumina ceramics for articulating implants at the present state.

INTRODUCTION

Since 1975 alumina bioceramic is used for prostheses in articulation against cups made of polyethylene. Long-term results have shown that this material combination reduces the clinical wear rate of the polymer to one half of the value for the pairing with CoCrMo (I).

The reduction in polyethylene wear is attributed to the superior surface finish and the

possibility of lubrlcation by chemisorption (2). Although the mechanical properties of

alumina are sufficient under certain circumstances fracture of alumina balls can occur. The technical advantages of other high performance oxide-ceramics with increased strength

and toughness has made them possible candidates for critical components of highly stressed implants. These ceramics are mainly transformation toughened zirconias, stabilized with the

oxides MgO or Y 203. They are characterized by high strength and toughness with values above 10 MN/m312 (3). Magnesium-oxide partially stabilized zirconia (Mg-PSZ) has shown

good biocompatibility (4) and excellent mechanical properties together with outstanding wear

resistance for technical applications. This ceramic consists of coarse grains with optimally

119

sized tetragonal zirconia precipitates in a cubic zirconia matrix (5).

MATERIAL AND METHOD

Discs of three different qualities of alpha-alumina bioceramic were obtained from two different

manufacturers, designated M and F. All qualities were conform to the requirements of the

ISO 6474 standard. Their hardness is typically in the range of 2300 HV. The other properties

are given in table 1.

TABLE 1 Properties of the alumina qualities used for the tribological investigation

Property Unit ISO 6474 ProducerM ProducerF ProducerF TypeB TypeBll

Purity % > 99.5 99.9 > 99.7 > 99.7 SiÜ2+Na20 % < 0.1 0.01 <0.02 <0.02 Density 'lJcrrtJ ~ 3.9 3.95 3.95 3.97 GrainSize ~ <7 2 4.3 3.6 Bending Strength MPa ~400 420 552 595 Compressive Strength MPa =4000 4400 5000 Young's Modulus OPa =380 420 380 Impact Strength Ncm/cm2 >40 >40 ~ Corrosion Resistance m'lJm2 < 0.1 0.008 < 0.1 klC MN/m312 4

Zirconia discs were obtained from manufacturer M and N. The Zr02 from producer M was

tetragonal zirconia polycrystalline (fZP) bioceramic stabilized with yttria. Manufacturer N

delivered magnesium-oxide partially stabilized zirconia (Mg-PSZ) discs. The properties of the

two zirconias are reported in comparison with another PSZ grade in table 2. Their hardness is around 1200 HV.

The surface of all ceramic discs was characterized at least with profilometrie, light and

scanning electron microscopy and the wetting angle was determined. Pins of 3 mm diameter

of medical grade ultra high molecular weight polyethylene (UHMWPE), conform to

ISO 5834, were used. They were all machined from the same section of a sheet and cut with a

microtome perpendicular to their diameter prior to testing (1).

Tribological tests were conducted in a clean room using a pin-on-disc apparatus with a

continuous sliding velocity of 0.025 m/s and a mean surface pressure of 3.45 MPa in a

mixture of Ringer's solution with 30 % calf serum (1). From the inductively determined

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eontinuous displacement of the pins the linear regression was determined and the volumetrie

wear factor k3 eomputed . The eoefficient of frietion was recorded eontinuously from a

eompression load eell.

TABlE2 Properties of the zirconia qualities used for the tribologieal investigation

Property Unit 1ZP Mg-PSZ YPSZ

Purity % '71 96.5 95 Yß?JMgO % 3 mol 3.4 wt 5wt Density g/crß3 6.05 5.72 6.0 Grain Size (average) Jlßl 0.2 - 0.4 0.42 0.5 Bending Strength MPa 1000 800 ~920 Compressive Strength MPa 2000 1850 7500 Young's Modulus GPa 150 208 220 kle MN/m3f2 7 8 10

RESULTS

Most of the dises investigated exhibited pores from the manufaeturing process andIor scratches

from the polishing step with diamond paste. The quality of the surfaee finish varies from very

good (TZP and alumina type B TI) to unsatisfactory (Mg-PSZ, alumina type M and B). The

wetting angle of all eeramie dises is at least 30 % lower eompared to the metal CoCrMo.

All results of the tribologieal investigation with pins made of UHMWPE showed stable

artieulation after a running in period of 24 to 40 hours. The wear rate of the polyethylene

displayed then linear increase and the value for the eoeffieient of frietion remained eonstant

over the testing period. The results with a minimum of 6 tests per material eombination are

summarized in table 3. Within all the single results of a material eombination no eorrelation

between the UHMWPE wear rate, the eoefficient of frietion and the surface roughness of the

discs was detected.

The average wear factor of polyethylene against all Ab03 eeramies was at least 20 %

lower than for the eombination against CoCrMo cast alloy. It is not dependent on the surfaee

roughness of the alumina dises. The enhaneed quality of alumina showed in this test a non

signifieant decrease in polyethylene wear rate together with inereased frietional values, while

the other qualities have a redueed eoeffieient of frietion of 10-15 % eompared to the metal

dise, displayed in table 3. The finer grained, smooth alumina quality M exhibited a non

significant decreased wear rate of the UHMWPE partner eompared to those with eoarser grains.

121

The wetting angle of the enhanced alumina quality is higher, indicating a reduced surface

energy for this quality.

TABLE3 Results of all tests with the pin-on-disc machine against UHMWPE pins

Property

Co-28Cr-6Mo Ah03 M Ah03 F/B Ah03 F/BII Zr<h 1ZP Zr<h Mg-PSZ

Roughness Rzijun)

0.304 ± 0.075 0.064 ± 0.029 0.263 ± 0.044 0.305 ± 0.105 0.155 ± 0.018 0.848 ± 0.247

Wetting angle(~

93 51 39 '!f:)

49 54

Wear factor of PE (nnn3jNm*lo-7)

2.72 ± 0 98 2.13 ± 0.40 2.26 ± 0.51 2.17 ± 0.74 3.26 ± 0.88 2.66 ± 1.81

Coeff.of friction

0.094 ± 0.016 0.079 ± 0.014 0.083 ± 0.002 0.119 ± 0.012 0.114 ± 0.020 0.093 ± 0.023

The roughness of the tetragonal zirconia polycrystals (TZP) and the contact angle for the

lubricating medium was in the same range as for the AhÜ3 ceramics. Nevertheless this

proposed alternative ceramic caused an increased UHMWPE wear rate. The values observed

were even higher than for the carbide containing CoCrMo alloy Protasul-2 which has an

increased roughness. Also the coefficient of friction was increased by about 20 % and comparable to the Co-basis alloy. Mg-PSZ with its extreme roughness values gave the best tribological result of the non-alumina ceramics.

DISCUSSION AND CONCLUSIONS

AhÜ3 bioceramic has dernonstrated c1inically a superior wear and frictional behaviour against

UHMWPE, when compared to the standard combination with mirror polished CoCrMoC cast alloy. There are at least two reasons for this: (i) Because of the single phase fine grained structure of alumina the surface roughness is lower than for the carbide containing bi-phasic

CoCrMo-alloy. This roughness is mainly inverse instead of the protruding asperities

encountered with the hard M7C3 carbides. (ü) The surface energy is high and has therefore the

ability to bind a lubricating molecular layer on the surface, which reduces wear and friction of

this material against the non polar polyethylene (2).

This investigation showed that all ceramic discs tested bad an unfavourable surface quality

when compared to standard ball heads. The roughness of the discs varies from Ra 0.01 to

0.27 ~m and the average roughness Rz from 0.06 to 0.85 ~. Nevertheless the wear rate and

122

frietion against polyethylene did not exhibit any eorrelation with the roughness of the different

ceramies. This is diverse to metal discs and ean be explained by the origin of the roughness

and the differences in surface energy of the two material groups. Protruding roughness, under

eonditions of mixed fIlm lubrieation, increases the abrasive wear while inverse roughness like

voids together with high polarity allows for better lubrieation and reduced adhesive wear.

The reduetion in polyethylene wear rate with alumina eompared to CoCrMo was only in

the range of 20 to 25 %, eompared to 35 - 75 % elinieally. This is the eonsequence of the

problems with manufacturing and polishing the flat ceramie discs to the same perfect quality

as the ball heads for the elinieal use. The roughness of the ceramie discs is a faetor of two

higher than that of the balls, while there is no difference between those two geometries for

Co-basis alloys. No signifIeant differenee in wear rate was detected between the three qualities

of bio-eeramies tested. The eoeffieient of frietion is 10-15 % redueed for two grades of

alumina eompared to the metal discs. This is in agreement with results of pendulum tests,

reported elsewhere (6).

The mechanieal properties, especially toughness and bending strength together with a

decreased modulus of elastieity (3), of the alternative ceramies TZP and Mg-PSZ make them

attraetive substitutes for alumina. This study exhibited unfavourable wear and frietion

behaviour of these high performance eeramies in artieulation against polyethylene, although

their roughness was eomparable to the AbÜ3 discs. The reason of this inferior tribologieal

behaviour against the non-polar polymer UHMWPE is not understood yet because no distinet

difference in the properties investigated has been traced. The result of increased friction of

zirconia in eombination with polyethylene presented in this investigation is in agreement

with another publieation showing an increase of more than 20 % eompared to alumina (7).

Non published results in a different laboratory with zirconia and alumina balls have shown

that the frietional efforts are remarkable higher for zirconia, and the wear rate of UHMWPE is

increased (8).

The deleterious transformation of tetragonal to monoclinie phase in yttria doped zireonias

due to ageing in water or humid environment at elevated temperarures, and the thus

dramatieally redueed toughness, is weIl established (9). Although short term laboratory and

animal tests (3, 10) have demonstrated the stability of zireonia in simulated or aetual biologie

environment there are also reports about increase in monoclinie phase oceuring at 40°C in

long-term testing (11). There are also reports about radioactivity associated with zirconia (12).

Although the aetivity is very small there still is some eoneern about the long-term effects.

It is therefore eonc1uded that zirconia is not recommended for elinieal use as a substitute for

alumina at this stage. After more than 15 years of sueeessful elinieal applieation in more than

250'000 eases Ab03 still is the ideal partner for ball heads of hip joint endoprostheses in

combination with UHMWPE. Improved alumina grades with fIner grain size have enhaneed

123

mechanical properties and allow for fracture proof components without effecting the tribological properties.

REFERENCES

1. Streicher, R.M. and Schön, R., Tribology in medicine. In ~ Biomaterialien fiU: ~ Endo!>rothetik. ed. G. Hofmann, Praxis-Forum. Berlin, 1989,19, pp. 20-39.

2. Dawihl, W. and Dörre, E., Adsorption behaviour of high-density alumina ceramics exposed to fluids. In Evaluation m Biomaterials. J. Wiley & Sons, Chichester, 1980, pp. 239-245.

3. Christei, P., Meunier, A., Heller, M., Torre, 1. P. and Peille, C., N., Mechanical properties and short-term in vivo evaluation of yttrium-oxide-partially-stabilized zirconia. L. Biomed. Mat. Rä, 1989, 23, 45-61.

4. Garvie, R. C., Urbani, C., Kennedy, D. R. and McNeuer, J. C., Biocompatibility of magnesia-partially stabilized zirconia (Mg-PSZ) ceramics, L.Ma1.& 1984,19,3224-3228.

5. Swain, M. V., Zelizko, V., Lam, S. and Marmach, M., Comparison of the fatigue behaviour of Mg-PSZ and alumina in Ringer's solution, Proceedings of the MRS meeting on advanced materials, Tokyo, 1988.

6. Streicher, R. M., Schön, R. and Semlitsch, M., Untersuchung des tribologischen Verhaltens von Metall/Metall-Kombinationen für künstliche Hüftgelenke. Biomedizinische Technik, 1990,35 (5), 107-111.

7. Davidson, 1. A., The effect offemoral head size and hardness on the frictional moment during articulation. In Advances in Bioeneineerin~, Proc. ASME, 1989, pp. 34-35.

8. Dörre, E., Private communication.

9. Sato, T. and Shimada, M., Transformation ofyttria-doped zirconia polycrystals by annealing in water, 1.. Am . .Qmun.~, 1985,68,356-359.

10. Mandrino, A., Moyen, B., Ben Abdallah, A, Treheurex, D. and Orange, D., Aluminas with dispersoids. Tribologic properties and in vivo ageing, Biomaterials, 1990, 11,88-91.

11. Cales, B. and Peille, C. N., Radioactive properties of ceramic hip joint heads. In Bioceramics Vol. 2. ed. U. Soltesz, German Ceramic Society, Cologne, 1990, pp. 152-159.

12. Thompson I. and Rawlings, R. D., Mechanical behaviour of zirconia and zirconia­toughened alumina in a simulated body environment, Biomaterials, 1990,11,505-508.

124

HIP ANATOMICAL UNCEMENTED CERAMIC ARTHROPLASTY (AN.C.A.): RESULTS AT A 3 YEARS FOLLOW-UP

L. Specchia, A. Moroni, L. Ponziani, G.Rollo, S. Pavone, V. Vendemia 3rd Department of Orthopaedic Surgery, Rizzoli Institute,

Bologna, Italy.

ABSTRACT

70 uncemented total hip re placements with bioceramic components have been perjormedlrom 1986 to 1988 at the third Department olOrthopedic Surgery 01 the Rizzoli Institute. 47 hip were evaluated at 3.8 years averagelollow-up (range 3-5 years). 38 01 these patients were lemale, 9 male. Age rangedlrom 42 to 65 years, average age was 54 years. In 20 cases (group A) a total AN.CA. uncemented prosthesis was implanted. In 27 cases (group B) mixed implants were perjormed: titanium uncemented acetabular component and AN.CA. stem. In the total AN.CA. devices an alumina prosthetic head was utilized, in the mixed implants a Ti NO 4 coated titanium prosthetic head. 40 surgical procedures were primary total hip replacements, 7 were revisions ollailed cemented prostheses. The patients have been evaluated according to the Harris Hip Scoring System. X-rays were analyzed according to the Engh criteria. Fifteen cases (31.9%) were excellent, 12 (25.5%) good, 12 (25.5%)lair and 8 (17%) poor. No statistical signijicative dijJerence was observed comparing the two groups 01 prostheses. Three acetabular components were evaluated as lailed: 2 01 these prostheses were revised. A serious damage 01 the prosthetic alumina head was noticed in these 3 cases.

INTRODUCTION

Ceramic materials can be divided in two groups: bioinert and bioactive ceramics. Bioceramies may be used to coat metallic prosthetic components and for the manufacturing of whole prostheses. The stern and the acetabular component of the AN.C.A. system are coated by alumina, the prosthetic head is alumina made and the acetabular component is manufactured by a compound of alumina and bioglass. (1)

Aim of the present study is to evaluate the clinical and radiographie results of aseries of AN.C.A. prostheses at a 3 years follow-up.

MATERIAL AND METHOD

The AN.C.A. prosthesis stern made by Chrome-Cobalt-Molibdenum alloy with "Wax" casting technology, has an anatomical shape, thus there are two specular prostheses for each different size required. In the proximal part, the surface is covered with a single layer of beads, with the exclusion of the lateral part, which is kept smooth. The entire surface of the stern is covered with alumina-oxide, applied using the air plasma spray technique.

125

The acetabular component, spherical in shape, is built with a composite structure made up of ceramic (alumina), titanium alloy and ceramic coating with tridimensional porosity (Al spheres with bioglass alloying element). The prosthetic 32 mm diameter head is Alumina made. (1)

Between 1986 and 1988,70 uncemented AN.C.A. hip prostheses were performed. 47 hips (38 females and 9 males) were evaluated at a 3.8 years average follow-up (range 3-5 years).

Lateral direct approach with the partial medius gluteus detachment and the total minimus gluteus was always used. All the patients were operated on by the same surgeon. In 20 (42.5%) cases (group A) a total AN.C.A. prosthesis was implanted, in the other 27 (57.4%) cases (group B) a titanium uncemented screwed socket associated to the AN.C.A. stern and to a titanium-N04 head was used. The primary implants were 40 (85.1 %): 32 patients were affected by osteoarthritis, 6 by femoral neck fracture and 2 by femoral head necrosis; the revisions were 7 (14.8%).

Clinical results were analyzed according to the Harris scoring system (2), evaluating the results as: excellent, when the range was from 100 to 90 points, good 89-80, fair 79-70 and poor< 70.

Radiographic results were analyzed considering the bone-prosthesis interface and the type of stabilization according to the Engh et al criteria (4). The functional and radiographic results of group A and group B were compared using the T-Student Test.

RESULTS

Fifteen cases (31.9%) were evaluated as excellent, 12 (25.5) good, 12 (25.5) fair and 8 (17%) poor.

Poor results were due to: mid-thigh pain in 3 cases, severe damage of the Alumina head associated to a failure of the acetabular component in 3 patients (Fig. 1-2); excessive lengthening of the operated limb causing a reduced range of motion and a severe limp in 2 cases. A slight lengthening of the operated limb (less than 1.5 cm) was present in 13 patients (27.6%).

As early postoperative complications we had: 1 calcar breakage; 2 pulmonary embolia (the ftrst patient died 3 days after the operation, in the second case a complete healing with suitable therapy was obtained) 1 phlebitis and 3 urinary tract infections.

Radiographie evaluation showed 3 failures of the acetabular component, 2 of these cases have been already revised (Fig. 1). During the revision procedure a serious wear damage of the alumina head was noticed. Damage of the alumina head was present in the third failed prosthesis too (Fig. 2). In 7 patients a radiolucency line in 3-4-5 zones of the stern was observed (3). In 1 case a slight subsidence of the stern was present. In all the other cases the bone prosthesis interface demonstrated a good bony stabilization but bone remodeling signs were present also (Fig. 3).

No signiftcative differences of the results were observed comparing the two groups of prostheses.

DISCUSSION

These data lead us to emphasize some topics that, in our opinion, affected dramatically the long term results.

First of all, the socket coils are too few to guarantee the necessary initial ftrm mechanical stability of the implant. Moreover, the surgical implantation technique, of the acetabular component is not easy, because the device for the prosthesis insertion often does not allow a correct screwing in the chosen position. The stern implantation technique is difftcult also,

126

Figure 1. a) failed cemented MuHer prosthesis in a 54 year old male patient. b) AN.C.A. prosthesis at I year. c) loosening ofthe acetabular component at 2 years, notice also the

remodeling around the stern. d) X-Ray image 1 year after the revision.

127

Figure 2. a) failed cemented Muller prosthesis in a 62 year old female patient. b) AN.C.A. pros thesis at a 1 year follow-up. c) loosening of the acetabular component at 18 months. d) dislocation of the prosthetic head along with rupture and severe wear damage at 24 months.

The patient refused arevision procedure.

128

Figure 3. good result at a 3 years follow-up observed in a 52 years old female patient Notice the hone remodeling around the stern.

expecially in the flrst cases, because the anatomical stem shape combined to its lengthen.The available pros thesis sizes are, in our opinion, too restricted, obliging in several cases to use an unflt stem. Furthermore the neck is probably too long and it should be responsable of the cases of limb lengthening.

Moreover we emphasize that problems related to the ceramic materials were detected in all the 3 failures. In the 2 revised cases a very severe damage of the prosthetic head with a severe alumina wear debris was observed during the revision. In the other failed prosthesis a severe damage and ropture of the Alumina head was radiographically observed. These data conflrm the results published by Christel et al.(5) suggesting that wear is very low in the alumina prosthetic joint components but also that this material is very sensitive to any possible malalignement which could be capable to cause a severe damage of the ceramic components and the final failure of the implant. Thus we emphasize that the ceramic materials of the AN.C.A. system, as weH as the design and the availability of other shapes and sizes, should be improved. BasicaHy we appreciate some biomechanical principles of the AN.C.A. system, like the ceramic coating of the metallic substrate and the ceramic-ceramic contact. Anyway the surgical experience and the clinico-radiographic results emphasize also how the use of these materials in the hip total joint replacement can be unforgiving.

129

REFERENCES

1. A.Toni, A.Sudanese, D.Cioni, D.Da11ari, T.Greggi, A. Giunti, L'artroprotesi anatomiea di eeramiea d'anea (AN.C.A.): esperienze preliminari eon una nuova protesi non eementata. Chir. Org. Mov.,1990, 75, 81-97.

2. Harris W.H., Traumatie arthritis of the hip after dislocation and acetabular fraetures: treatment by mold arthroplasty. J.B.J.S., 1969, SI-A, 4.

3. Gruen T.A., Me Neiee G.M., Amstutz H.C., Modes of failure of cemented stem-type femoral components: a radiographie analysis of loosening. CHn. Orthop., 1979,141, 17-27.

4. Engh C.A., Bobyn J.D., Glassman A.H., Porous-eoated hip replacement. The faetors governing bone ingrowth, stress shielding and clinieal results. L Bone Joint Surg., 1987, 69/B, 45-55.

5. Christel P., Meunier A., Dodot J.M., Crolet J.M., Witvoet J., Sedel L., Boutin P., Biomechanical compatibility and design of eeramie implants for orthopaedie surgery. Bioceramics Material Charaeteristics. Acad. Sei. New York, 1987.

1~

FLUORAPATITE AND HYDROXYAPATITE HEAT-TREATED COATINGS FOR DENTAL IMPLANTS

H.W. DENISSEN, H.M. DE NIEUPORT, W. KALK, H.G. SCHAEKEN and A. VAN DEN HOOFF

TRIKON: Institute for Dental Clinical Research University of Nijmegen, PO Box 9101,

6500 HB Nijmegen, The Netherlands

ABSTRACT

Sintered fluorapatite and hydroxyapatite ceramic implants have been shown to be non-degradable in animal and human research. It was found however that plasma-sprayed coatings of hydroxyapatite ceramic degraded in vivo, possibly because the ceramic lost almost 50% of its crystallinity during the plasma-spraying procedure. A heat treatment at 6000 C proved to be the method op choice to reconstitute the crystallinity of the hydroxyapatite coating and to promote the transforma­tion of a -tricalcium phosphate into the more stable ß -tricalcium phosphate. Fluorapatite coatings did not loose their crystallinity during plasma-spraying and had the added advantage of stability due to the fluoride ion. The solubility in vitro of the fluorapatite and hydroxyapatite heat-treated coatings were compared to the sintered hydroxyapatite ceramic and the untreated hydroxyapatite coatings. Heat treatment reduced the dissolution rate of the hydroxyapatite coating which originally was 92%, to only 9%, while 13% of the fluorapatite coating was found to dissolve.

INTRODUCTION

Fluorapatite (FA) and hydroxyapatite (HA) powders can be

compressed and sintered to yield bulk ceramic materials. Bulk

FA and HA have been found to be non-degradable in animal

research. 1,2 In long-term research bulk HA has also been

shown to be non-degradable when inserted in human alveolar

bone or soft tissues (Figs. 1 and 2). 2

131

t Figure 1. Radiograph taken of an endosteal bulk HA ceramic implant placed in the incisor region of the mandible of a patient in order to maintain the height of the alveolar ridge under the fixed bridge. The implant has the shape of a

cylinder (arrows).

Figure 2. Radiograph of the same patient as presented in figure 1, made 15 years after implantation. No radiographie change in the outline of the implant surface can be detected,

neither in the bony nor in the subperiosteal environment.

However bulk FA and HA ceramies as such cannot be used in

load-bearing situations because of their poor mechanical

properties. For this reason FA and HA ceramies are applied on

a titanium substrate by a plasma-spraying technique. 3 We

have reported on mandibular bone response to plasma-sprayed

coatings of HA. 4 From radiologieal, macroscopical and

132

microscopical observations it was concluded that the biologi­

cal properties of plasma-sprayed coatings of HA are the same

as those of sintered HA ceramic. Free particles of HA were

observed in the surrounding bone tissue, presumably as a

result of failure of the coherence of the HA particles at the

outer layer of the coating (Fig. 3). 4 Other research groups

found particles in the regional lymph glands as welle 5

Figure 3. Undecalcified section of bone (A) adjacent to HA­coated dental implant (B) that has been in situ for 9 months in the mandible of a doge A very tight connection between the lamellar bone and the HA coating is found. However clusters of HA coating particles are present in the bone marrow (arrows).

As long as we don't know the ultimate fate of the lost HA

material in the human body and although no adverse effects have

yet been reported, it seems obvious that research should be

directed towards the development of stable bioactive coatings

of implants to be inserted into alveolar bone. Furthermore it has been shown by mechanical push-out tests that highly

degradable calcium phosphate coatings lead to a low shear

strength between coating and adjacent bone,which can adversely

affect the functioning of a dental implant under loading

conditions.6 If the coating would disappear completely the

advantage of the bioactive interface would be lost and a

133

direct contact between the metal substrate and bone would

result, undoing all advantages.

Therefore we focussed our research on improving the

stability of the coating in two ways. First, by replacement of

the OH-ion by the F-ion in the coating powder, thus obtaining

a FA coating. Second, by heat treatment of the HA coating.

In this study we evaluated the material characteristics

of FA coatings and HA heat-treated coatings and compared the

solubility behaviour of these coatings with bulk HA ceramic

and untreated HA coatings.

MATERIALS

The cylinder-shaped bulk HA ceramic was made of a

commercially available HA powder, which was precompressed and

sintered at 13000 C during seven hours. The particles used for

the plasma-spraying procedure of the HA coating were prepared

from granulated, ground and sieved bulk HA ceramic. 7

The particles for the FA coating were prepared in the same way

from compressed and sintered amounts of calcium fluoride,

calcium hydrogen monophosphate and calcium carbonate at 12500 C.

The titanium dental implants used for coating were cylindrical

and screw-shaped. Our purpose was to examine a design which

was in clinical use i.e. not specially prepared for X-ray

diffraction or other laboratory experiments. We thus avoided

maximizing results which might be at variance with the results

obtained with clinically used implants.

METHODS

Plasma-spraying of the FA and HA coating was done with a

Metro equipment, using a standard powder port and hopper. As

arc gases nitrogen, or argon were used. Part of the sampIes of

the HA-coated implants were used for experiments with heat

treatment. Two temperatures were used for heat treatment:

6000 C and 9250 C.

The FA coatings remained untreated.

134

For X-ray diffraction studies we used a Philips PW 1050

diffractometer. The scanning rate was 0,250 28 per minute.

Registration started at 230 2 8; 40 KV and 30 mA was used.

Diffraction registrations were made of bulk HA and of plasma­

sprayed HA, FA and HA heat-treated coatings.

Dissolution rates of the same materials were measured in

a 0,5% acetic acid solution at 220 C during 30 minutes at pH 3.

The dissolution fluids were stirred continuously. After the

test the materials were rinsed in distilled water and dried at

1200 C. Before and after dissolution tests the weights of the

samples were determined on a balance (Mettler AT 250). The

measuring error was about 0,2 mg. Thereafter the residual

coating was brushed with a soft brush in order to remove loose

material from the coating, and weighted again. The weight loss

was related to the total weight of the bulk HA, or plasma­

sprayed coating and to the total surface area of the sample.

This surface area was the macroscopic surface, in other words

it was assumed that the surface did not have pores.

Animal Experiments

FA and HA heat-treated coated dental implants were placed

jn the premolar region of adult Beagle dogs as has been

described earlier for untreated HA-coated implants (Fig. 4). 4

The implants were clinically and radiographically followed-up

during aperiod of 18 months.

Figure 4. Radiograph of part of the mandible of a dog with 2 FA-coated implants directly after implantation. The same implants were used for our X-ray diffraction and dissolution tests. Radiolucencies are present around the implants because

the implants were placed in cylindrical implant beds.

135

RESULTS

The X-ray diffraction studies of the sintered bulk HA

implants showed patterns which were characteristic for HA.

ß -tricalcium phosphate (ß -TCP) was also present. However, in

the untreated plasma-sprayed HA coatings we found a -tri­

calcium phosphate ~-TCP)instead of ß -TCP.

When comparing the pattern of a sintered bulk HA implant with

that of a plasma-sprayed HA implant it is evident that the

ratios between height and width of the peaks- and thus the

crystallinity- had markedly decreased during plasma-spraying

(fig.5).

H BULI HA

H

35 33 30 !1

Figure 5. X-ray diffraction pattern of a bulk HA implant and of an untreated HA-coated implant. Because of the cylindrical shape of both implants the patterns are less pronounced than patterns from specially prepared flat sampies. However it is clear that the patterns show characteristic HA features but that the degree of crystallinity of the plasma-sprayed HA coating has decreased considerably compared with that of the bulk HA implant material, which was a starting material for

the plasma-sprayed coating.

Heat treatment at 6000 C or at 9250 C improved the

crystallini ty of the HA coating. Figure 6 shows the differences

in effect of the two types of heat treatment and compares

the patterns to the pattern of a FA coating.

136

HA HEAT-TIIEUE

FA -I ~-

35 3S 30 27

Figure 6. X-ray diffraction patterns of a HA coating heat­treated at 6000 C (top), at 9250 C (middle) and an untreated FA coating (bottom). The tracings have to be compared with the bottom tracing in figure 5. Crystallinity had improved most with the 6000 C treatment. Crystallinity of the untreated FA coating, as far as can be seen from these patterns, is only slightly less than of the heat-treated HA samples. Therefore

we saw no reason to give the FA coating a heat treatment.

Tables 1 and 2 show the results of the dissolution tests

in 0,5% acetic acid.

TABLE 1 Average weight loss of bulk HA and plasma-sprayed HA, FA and HA heat-treated coatings in 0,5% acetic acid solution at 220 C. The number of samples of each

coating was 5.

Coating Average weight Percentage of weighi loss loss (in mg) weight loss per mm (in mg)

(Bulk HA) 0,4 4 0,003 HA 8,3 92 0,059 FA 1,2 13 0,009 HA HT 0,8 9 0,006

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TABLE 2 Average weight loss after brushing. The number of

samples of each coating was 5.

Coating

(Bulk HA) HA FA HA HT

Average weight loss (in mg)

2,5 0,9

* no remains of coating left

Percentage of weight loss

8 * 2

The radiographie results in experimental research are

illustrated by a typical case in figure 7. From a radio­

graphical and clinical point of view the implants behaved like bulk HA implants and HA untreated coatings. 4

Figure 7. Radiograph of the same implants as in figure 4, one year after implantation. The radiolucencies have disappeared

and the FA implants are well tolerated.

DISCUSSION

One of the main objectives in dental implantology is to

obtain a non-degradable bioactive coating forming an inter-

138

face with alveolar bone. It is imperative that no clusters of

particles of the ultra-thin coating should be lost in the

surrounding tissue, because in case the coating desintegrates

completely the implant is only retained by mechanical

retention and might loosen.

It seems that crystallinity is an important factor in the

solubility behaviour of a ceramic material~Bulk HA ceramic

was a stable material because the sintering temperature did

not exceed 13000 C. Consequently the size of the HA

crystallites, and thus the crystallinity, was maintained

during the sintering procedure. 9 This resulted in a stable

implant surface in the dissolution test and in our clinical

research. 2 However the temperatures reached during plasma­

spraying, reduced the crystallinity of the HA starting

material to almost 50% of its original value. 9 This resulted

in a HA coating which was soluble in vitro and degradable in

animal experiments. 4

According to the X-ray diffraction results heat treatment

increased the crystallinity of the HA coating. Consequently

the coating was much more resistent to dissolution. This is in

agreement with the findings of other research groups who

concluded that the dissolution rate of calcium phosphate

ceramics depends on its crystallography and stoichiometry of

the material. 8 The ß -TCP component of bulk HA ceramic

turned completely into a -TCP at 11000 C during the plasma­

spraying procedure. Heat treatment reversed this process

because no a -TCP was present in the heat-treated coating.

Because ß -TCP is less soluble it is not surprising that heat

treatment has a beneficial influence on the stability of a HA

coating.

The stability of the FA coating can be explained in terms

of different behaviour as regards transformations during

plasma-spraying. As distinct from HA, the FA has a molten

phase and does not change over the critical temperature of

13000 C which occurs during the plasma-spraying procedure.

Its high crystallinity, although not 100%, made it much more

stable in our solubility tests than untreated HA. A 0,5%

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acetic acid solution is not a physiologie environment for the

coating surface. However it was evident from the results that

there was a similarity between the in vitro behaviours of the

bulk HA and the untreated HA coating on the one hand and the

in vive behaviours in human and animals on the other hand.

Both in vitro and in vivo, sintered bulk HA appeared to be

stable and HA plasma-sprayed coatings unstable. This might be

an indication that untreated FA and HA heat-treated coatings

which in vitro behave similarly to bulk HA, also exhibit a

much more stable pattern on the tissue level.

CONCLUSIONS

The high crystallinity of the FA ceramic coatings was

maintained during the plasma-spraying procedure. These

coatings had the added advantage of stability due to the

presence of the F-ion.

A heat treatment of a HA coating at a temperature of 6000 C

enhanced the crystallinity most and promoted the transforma­

tion of a-TCP into the more stable ß-TCP.

These phenomena resulted in a reduction of the dissolution

rate of the HA coating which originally was 92% to only 9%.

The FA coating was almost as stable and dissolved 13%.

To definitely substantiate the benefits of the untreated

FA coatings and the heat-treated HA coatings in alveolar bone

an additional histological investigation is necessary.

REFERENCES

1. Heling, L., Heindel, R., Merin, B., Calcium fluorapatite,a new material for bone implants. J. Oral Implant, 1981, 9, 548-55.

2. Denissen, H.W., Kalk, W., Veldhuis, A.A.H.,Van den Hooff,A. Eleven-year study of hydroxyapatite implants. ~. Prosthet Dent, 1989, 61, 706-12.

140

3. De Groot, K., Geesink, R., Klein, C.P.A.T. and Serekian, P., Plasma-sprayed coatings of hydroxyapatite. !l.. Biomed Mater Res, 1987, 21, 1375-81.

4. Denissen, H.W., Kalk, W., De Nieuport, H.M., Maltha, J.C., Van de Hooff, A., Mandibular bone response to plasma­sprayed coatings of hydroxyapatite. Int. J. Prosthodont, 1990, 3, 53-8. --- -

5. Klein, C.P.A.T., Patka, P., Den Hollander, W., Macroporous calcium phosphate bioceramics in dog femora: a histological study of interface and biodegration. Biomaterials, 1989, 10, 59-62.

6. Klein, C.P.A.T., Patka, P., Van der Lubbe, H.B.M., Wolke, J.G.C. and De Groot, K., Plasma-sprayed coatings of tetra­calcium-phosphate, hydroxyl-apatite, and a TCP on titanium: An interface study. !l.. Biomed Mat Res, 1991, 25, 53-65.

7. Denissen, H.W., Veldhuis, H.A., Rejda, B.V., Dense apatite ceramic implant systems: a preliminary report. !l.. Prosthet Dent 1983, 49, 229-33.

8. Klein, C.P.A.T., Driessen, A.A., De Groot, K., Relation­ship between the degradation behaviour of calcium phosphate ceramics and their physical-chemical characteristics and ultra-structural geometry. Biomaterials, 1984, 5 157-60.

9. Koch, B., Wolke, J.G.C. and De Groot., X-Ray diffraction studies on plasma-sprayed calcium phosphate-coated implants. !l.. Biomed Mat Res 1990, 24, 655-67.

141

POROUS TITANIUM IMPLANTS WITH AND WITHOUT HYDROXYAPATITE COATING

A. Moroni, V. Caja, E. Egger, F. Gottsauner Wolf, L. Trinchese, G. Rollo, E.Y. Chao

Rizwli Orthopaedic Institute,ViaPupilli 1,40136 Bologna, ltaly. Orthopaedic Biomechanics Laboratory, Mayo Clinic, Rochester, MN, USA.

ABSTRACT

The response to unloaded titanium porous coated implants in the canine femoral cancellous bone was studied. The implants beads size was 500-700 micron. Seven samples were coated with hydroxyapatite (HA) by air plasma spray and equal number ofuncoated implants were used as the controls. Dogs were euthanized after 12 weeks. No statistical significative differences were observed between the HA coated and uncoated implants in the push-out test results. Statistical significative greater direct bone apposition to the implant surjace was observed in the coated samples compared to the uncoated ones.

INTRODUCTION

Bone fixation of joint prostheses can be achieved with or without bone cement. Cementless

fixation is obtained through metallic devices with porous coating to achieve bone ingrowth.

Short term results are usually positive but long term results are less satisfactory. Failure is

related to biomechanical and biological reasons. To enhance implant fixation, hydroxyapatite

(HA) coating of metallic prosthetic surfaces (1-2) could be usefull preventing the release of

metallic ions and improving the dirett bone apposition to the prosthesis surface. The aim of

our study was to evaluate the response of the canine femoral cancellous bone to HA coated

and uncoated porous titanium implants.

MATERIALS AND METHODS

Fourteen titanium porous coated sampIes supplied by De Puy, USA, were utilized. The

sampIes had a cylindrical shape, the length was 22 mm and the diameter 5 +/- 0.3 mm.

142

Beads size was 500-700 micron. Seven sampIes were HA coated by air plasma spray

technique (Biocoatings, Flametal, Fornovo Taro, ltaly).The thiekness of the HA coating was

40-80 micron.

SampIes were randomized bilaterally implanted in the distal femoral cancellous hone of

7 mature mixed breed dogs. Through a lateral approach a 3.5 mm drill was advanced to a

depth of 22 mm perpendicularly to the long axis of the femur. The hole was reamed to a

diameter of 0.2 mm smaller than the sampIe to be implanted. Then the cylinder was press

fitted into the hole to lie in the cancellous hone medullary canal.

Dogs were kept in cages, allowed for free movements. Monthly, radiographie analysis

was performed.

Dogs were euthanized 12 weeks after surgery and femurs were removed and dissected

free of soft tissue. The implant hone block was cut four times, perpendicularly to the

longitudinal axis of the cylinders, in order to get three sampIes discarding the block ends.

Two sampIes, approximately 3 mm thick, were obtained from each implant and centered

over a 7 mm hole in a elevated aluminum plate and push-out of the surrounding hone at a rate

of 2 mm/min in the material testing machine (Instron Mod. 1125, Instron Corporation,

Canton, Mass., USA). In each push-out sampIe the thickness, diameter and surface area

was calculated and used to compute norma1ized mechanical results.

The remaining sampIes were processed for undecalcified sectioning and embedded in

polymethylmetacrilate. Cross sections were cut with a diamond saw (Leco Vari Cut VC50,

USA; No. 114245 Diamond Wafering Blade). After ground 100 micron thickness sections

were obtained and polished with alumina powder. Microradiographies of each sections were

performed in a standard technique. The microradiographic cross sections were quantitatively

and qualitatively evaluated. Histomorphometric analysis was performed using a Merz grid. (3)

This analysis was performed on a 90° randomized quadrant in each one of the 28 available

sampIes. A 146 points counting analysis for each quadrant was used to determine the

amount of hone in the ingrowth area (hone % in the section extending 500 micron from the

substrate), hone ingrowth (hone % in the available space of the ingrowth area) and porosity

( available area for hone ingrowth in the ingrowth area). A 59 counting points analysis was

used to evaluate the hone ongrowth (hone % in the section extending 220 micron from the

ingrowth area).

The perimeter of the implant in each microradiographic sector was digitazed with asonie

device (Graph Pen Sonie Digitizer, SAC, USA) in order to measure the percentage of the

surface of the implant with direct hone contact at 35 x magnifications (no visibile space

between hone and implant). Paired t-test were used to compare HA coated vs uncoated

results.

143

Figure 1. Photography showing one HA coated cylinder.

Figure 2. Macrographic test showing the beads in a HA coated sampie.

RESULTS

Push-out test results are shown in Table I. No statistically significant differences were found

in raw force for push-out test results comparing the HA coated and uncoated sampies.

Likewise, no differences were found when force was normalized to take into account either

plug diameter, plug length or calculated surface area

A different morphology of the ttabeculae close to the implants was observed in the HA

coated sampies compared to the uncoated ones. In the uncoated cylinders ttabeculae close to

the implants show an arc shape with a partial contact with the titanium surface (Fig. 4). In the

144

HA eoated eylinders, trabeculae close to the implants show a mushroom shape with a wider

eontact with the implant (Fig.3-5)

Histomorphometrie results (Table I) show no statistical signifieative differenees in

amount of bone in the ingrowth area, bone ingrowth, porosity and bone ongrowth comparing

eoated and uneoated HA sampIes. Statistieal signifieative inerease of the direet bone

apposition to the implant surfaee between the HA eoated and the uneoated sampIes was

observed.

Figure 3. Microradiography showing direct bone apposition through mushroom shape

trabeculae in a HA eoated sampIe.

Figure 4. Histologieal specimen showing an HA uneoateQ sampIe. Notice the spaee

between arc shape trabeculae and titanium.

145

Figure 5. Histological specimen showing a mushroom shape trabecula in aHA coated

sampie. Notice the direct bone apposition to the HA coating.

DISCUSSION AND CONCLUSIONS

HA coating of smooth metallic surface still presents problems as fatigue fractures of the HA

coating. (4) Because of that, this study is dealing with porous coated sampies. Titanium was

used because titanium is more reactive chemically in comparison to Cr-Co and other metallic

alloys allowing a stronger link with the sprayed HA. (2)

HA coating of titanium porous prosthesis has been proposed recently by several authors

to improve the bone prosthetic ftxation. (24·5·6) Their results demonstrate that HA coating of

a porous metal surface is effective for early stronger ftxation. (5·6) Improvement of the shear

strength of the bone implant ftxation was reported at 2 and 6 weeks after the implantation. (4)

On the contrary, at 12 weeks since implantation the shear strength of the bone implant

interface of HA coated versus HA uncoated porous titanium implants did not show

signiftcative differences. (4) Oonishi et al. (4) suggested that at this time the push-out test

could measure the shear sthength of the bone around the implant more than the shear strength

of the bone implant interface. All these data demonstrate the importance of the implantation

time on the bone implant strength ftxation.

146

In Out study also no differences of the shear strength were observed between the HA coated and the HA uncoated porous implants despite the greater bone implant direct contact observed in the HA coated sampIes. Comparison of Out push-out results with the results

reported by the other authors is not possible because the different size and shape of the implants, technique of implantation and push-out model.

In all sampIes a good osteointegration has been observed, however, the morphological

results were different in the two groups of sampIes. Mushroom shape trabeculae with direct

bone apposition onto the implant surface were seen in the HA coated implants. This fact is related to the significant differences on bone contact perimeter observed in the HA coated versus the uncoated samples and confmns the HA osteoconductive properties which, in Out

opinion, are definitely an advantage of this coating.

TADLE 1 PUSH OUT AND mSTOMORPHOMETRIC RESULTS

HA Coated Sampies Uncoated Sampies

Shear strength 6.84 +/- 3.3 MPa 6.63 +/- 3.9 MPa

Boneamount ingrowth area 12.6 +/- 1.9% 17.8 +/- 5.5%

Bone ingrowth 52.5 +/- 12.6% 49.5 +/- 15.6%

Porosity 33.7 +/- 5.8% 26.9 +/- 6.2%

Bone ongrowth 44.3 +/- 13.9% 44.8 +/- 18.7%

Direct contact bone-implant 28.5 +/- 12.0% * 49.6 +/- 14.2% *

* P<0.012

147

REFERENCES

1. Denissen, H.W., Oe Groot, K., Makkes, P.C.H., Van Oe Hoff, Klopper, P.l., Tissue response to dense apatite implants in rats.l Biomed Mat Res. 1980,5,113-721.

2. Ducheynne, P., Hench, L.L., Kagan, A., Martens, M., Bursens, A., Mulier, I.C., Effects of hydroxyapatite impregnation on scheletal bonding of porous coated implants.l Biomed Mat Res. 1980, 14, 225-237.

3. Kimmei, D.B., Webster. S.S.I., Measurements of area perimeter and distance: Details of data collection in bone histomorphometry. In : Recker RR, Ed. l!.2.n..s; hystomorohometry techniQues and inteTJ!retation : CRC Press, 1 st Ed. Boca Raton, 1983, 89-108.

4. Oonishi, H., Yamamoto, M., Ishimaru, M., Tsuji, E., Kushitani, S., Aono, M., Ukon, Y., Tbe effect of hydroxyapatite coating of bone growth in to porous titanium alloy implants. 1 Bone Joint Surg .1989, 11 B, 213-216.

5. Hench, L.L., Splinter, R.l., Allen, W.C., Bonding mechanisms at the interface of ceramic prosthetic materials: 1 Biomed Mat Res Symposium 1971,2 (part 1),117-141.

6. ,Ducheynne, P., Structural analysis of hydroxyapatite coatings on titanium. Biomaterials §1986, 7, 97-103.

148

SURFACE REACTIVITY AND BIOCOMPATIBILITY OF BULK GLASS AND GLASS COATINGS

BRUNO LOCARDI Stazione Sperimentale deI Vetro

Via Briati, 10 30141 Murano (Venezia), ITALIA

ABSTRACT

Following a comparison of conventional and biofunctional glasses, the bioactivity of glass was examined associated with a hydrolytic degradation process and to the development of a hydroxyapatite calcium phosphate film at the glass/bone tissue interface. The main compositions of bioactive glasses are reported. The author's experiments in the field of reactivity and biocompatibility of these glasses are described as a part of a cooperative work carried out by the Stazione Sperimentale deI Vetro and GSB (Gruppo di Studio Biovetri). A "BIOVETRO" has been developed exhibiting good osteoconductive properties, and i ts possible applications as coatings with the same features deposited by plasma spray onto metal substrates have been evaluated.

I NTRODUCTI ON

Traditionally, glass has been considered as the passive element in the manufacture of containers and in the conveyance of light; this is contras ted by the active role played today by a whole new generation of glasses which are classified according to their roles in the fields of optics, electronics, chemistry, biology and biomedicine.

Bioglasses and bioglassceramics belong to the biologica11y active group of glasses , that is to say those glasses which, when placed in contact with cellular tissues, demonstrate a biocompatibility both in vivo and in vi trio and an absence of inflammatory and toxic processes. Furthermore, in the presence of precursory osteogenetic conditions, they also demonstrate an osteoconductive predisposition which tends to favour a particularly good biological bond at the interface between glass and bony tissues.

The bioactive category includes the bioinerts, characterized by a limited reaction at the interface, and the biodegradables, for which two processes : biodegradation of the implant - growth of the bony tissue occur simultaneously, ensuring good affinity between the tissue and the material of the implant.

What are the physical and chemical differences and similarities which distinguish or tie bioglass to traditional glass?

As for a11 other glas ses , bioglass is derived from a mixture of inorganic raw materials of analytical purity which, when heated to between 1,300° ands 1,600°C depending on their oxide composition,

149

trans form into homogenous molten 1 iquids completely devoid of crystall ine or gaseous inclusions. As this molten mass is carefully cooled, it exhibits a progressive increase in viscosity, evolving from its liquid high temperature form to asolid, amorphous substance at room temperature, without crystallization, the latter having been inhibited of slowed down.

The term solid describes he re a rigid material, that is to say one which is practically underformable, with a viscosity equal to or above some 1015 poise.

The term amorphous indicates that the three-dimensional organization of the structural units lacks the geometric order characteristic of the crystalline state.

To produce a vitreous state starting from a liquid state, it is necessary for the cooling rate of the substance under consideration once it has passed below the melting point to be greater that the crystallization rate. Since the vitreous state is not one of equilibrium, it turns out that glass is a thermodynamically unstable substance which tends to evolve towards more stable conditions, that is to say it tends to crystallize.

Glass ceramics are glasses whose crystallization tendencies have been exalted, that is to say materials which, having been produced in a vitreous state by traditional fusion methods, are given a subsequent thermal nucleation and controlled crystallization treatment, with the consequent development of a ceramic product with a high state of crystallinity composed of minute crystals set in a residual vitreous matrix left over from the crystallization process.

Bioglasses and bioglassceramics are therefore materials which, even though produced following the classical procedures of glass technology, nevertheless stand out for their biological functions associated wi th specific chemical compositions.

A typical product is the bioglass experimented with in the seventies by Hench and associated (1) (2) based on calcium phosphosilicate and with a chemical composition of Si02 45%, CaO 24,5%, Na20 24.5%. P205 6%.

As for all glasses it contains a vitrifying agent: silica, a melting agent: the oxide of sodium. a stabilizer: the oxide of calcium, but differs due to its phosphorus content and to its particular silica/alkali oxides and alkaline earth oxides ratio which render the glass hydrolitically unstable. Its bioactivity is in fact associated to a process of surface hydro li tic degradation of the vitreous phosphosilicate.

In fact, placed in contact with a solution of blood plasma, bioglass undergoes aseries of sequential physico-chemical transformations as foliows: 1) rapid exchange of sodium and potassium ions in the glass with hydrogen ions in solution; 2) migration into solution of soluble silica following the breakup of the siloxy Si-O-Si bonds and the formation of ionic silica groups ofthe SiOH or Si(OH)4 type at the glass interface; 3) condensation and polymerization of the ionic silica groups with of the formation of layers of silica in the form of a gel; 4) migration of calcium ions and phosphoric anions towards the surface of the silica gel layer; 5) formation of a film rich in calcium phosphate; 6) growth of the silica layer controlled by the interdiffusion of the alkaline ions exchange; 7) growth of the amorphous calcium phosphate layer with the incorporation of calcium and phosphorus ions from solution;

150

8) crystallization of the amorphous calcium phosphate film with the incorporation of OH-(C03)2 eF anions, forming a layered structure of the mineral fluoroapatite hydroxycarbonate; 9) agglomeration and production of a chemical bond between the apatite crystals and the collagenous fibrils and with other proteins produced by osteoblasts or fibroblasts.

The outcome of this reaction and the development of a crystalline layer of calcium phosphate set within a matrix of silica gel is that of a stable bond between glass and bone tissue.

According to Gross and Strunz (3) the bond between implant and bone tissue is controlled by the migration of monophosphates on the surface of the glass, whereas an opposite effect seems to result from the transfer of tri-tetra- and poly-phosphates from the siliceous matrix. Again according to these authors, the mineralization process of the bone tissue would be affected, if not inhibited, by the local mono/polyphosphate ratio whereas the soluble ionic calcium and silicon concentrations would be determining factors on the functions of osteocalcines and glycoproteins at the implant's interface.

Subsequent studies conducted specifically to define the general biocompatibility of glasses demonstrated that the preceding composition falls within the field of the ternary diagram silica-calcium oxide­sodium oxide characterized by a high surface reactivity.

As shown in tab. No. I, typical bioglass compositions are restricted to within weIl defined intervals for Si02(20-60), Na20(3-30), CaO (10-S0), P20S (0-18) and CaF2 (0-18).

TABLE 1

System Compos it ion

Na20-CaO-Si02-P20S Na20 CaO Si02 P20S

Na20-K20-MgO-CaO-Si02-P20S Na20 K20 MgO CaO Si02 P20S

MgO-CaO-Si02-PZOS MgO CaO Si02 P20S CaF2

Na20-K20-MgO-CaO-AIZ03- NaZO/K20 Si02-P20S-F MgO

CaO A1203 Si02 P20S F

24.S % 24.S 4S.0 6.0

S-10 wt% O.S- 3.0 2.S- S.O

30 -3S 40 -SO 10 -lS

4.6 wt % 44.7 34 16.2

O.S

r 8 wt % 2-21

10-34 8-1S

19-24 2-10 3-23

Mark

Bioglass@

Ceravital ®

A-W

Bioverit®

151

EXPERIMENTAL

The GSB study group, which has been operating in Italy in the bioglass filed since 1982, has prepared, in cooperation with the Stazione Sperimentale deI Vetro, different formulations of bioglass compounds and has refined experimental test procedures wi th the aim of developing glasses of different bioactive characteristics (3) (4).

Results of the Group's efforts include the development of a particular BIOVETRO composition which on the basis of in vive and in vitrio tests has been found to be biocompatible, biologically active and with having osteoconductive properties when associated to osteogenetic precursors. The vitreous material so obtained can be reduced to particles of a controlled grain size as needed, can be shaped to the form of cylinders or small bars, or alternatively transformed to a continuous filament of a few microns diameter via high temperature extrusion of the glass using a special crucible die.

In practice the research activity of the Gruppo Biovetri centres around an objective which is also the theme of the present conference: it deals specifically with the possibility of coating metallic prostheses of ti tanium composition with sprayed-on BIOVETRO layers of proven osteoconductive properties.

Given the early stages of the research, for the time being only limited preliminary results are available; these are being presented here because they are considered to be of interest for future developments of the research programme under way.

As is weIl known, there are several methods for producing ceramic coverings on metallic substrates; these include the eYD, sputtering, the SOL-GEL procedure, ionic implanting and lastly the plasma-spray process. The latter is the most suitable amongst those mentioned for the production of coatings starting with granular material.

In practice the principle of spraying powder using plasma consists in the immission of gas into achamber known as a torch across which a 30-40 KW electric arc is sparked, raising the gas temperature to 10,000-30,000 °K. Powders of different grain sizes, carried in a stream of gas, are then injected into the torch chamber. The powders melt and after accelleration are forced to flow through a nozzle and are shot energetically onto the substrate to be covered.

The particle velocity may vary from 100 to 350 rn/sec, so that the flight time is in the order millionths of a second. At the moment of impact the particles yield their thermal and kinetic energies to the substrate, they deform into lenticular shapes and solidify in 10-6 seconds.

By making the torch sweep over the piece being coated, one obtains a coating of particle upon particle in successive layers until the desided coverage is obtained.

The advantages of the plasma-spray technique include a high depositional velocity, a reduced alteration of the metallic substrate and a minimum dimensional tollerance.

Inconveniences may arise should the liquid phase of the material not be very stable, if the coating has cavities or micro-porosities, or if there is poor adhesion to the substrate caused by a limited interaction between the material sprayed and the metal.

As part of these preliminary studies on the possible applications of plasma-spraying of BIOVETRO onto metallic supports, operational conditions of the process were investigated by observing the morphology and nature of the glass deposited. The osteoconductive properties of these vitreous coatings were also evaluated be fore embarking on to the

152

more demanding experimental phase of series production of prototypes of BIOVETRO coated titanium prostheses.

RESULTS

Thus, after conducting aseries of experiments on the spraying of BJOVETRO onto metallic supports, the operational conditions necessary to give perfect bonding between bioglass and metal were developed, permitting the continuation of preliminary tests, the results of which are now described.

Figure 1 gives a clear indication of the morphological structure and nature of the BIOVETRO coating: it shows a SEM photomicrograph of part of the surface of sprayed-on BIOVETRO; figure 2 gives a view of the thickness of the layer applied to the metallic support.

It is possible to see that in both cases the structure of the depos i ted BIOVETRO does not appear to be homogeneous: there are both microcavities and more or less spherulitic grains welded together . Adhesion to the metal is not uniform and discontinuities are present.

Figure 1. SEM micrograph of the surface of a sprayed BIOVETRO. 200 X

153

Figure 2. SEM micrograph of the thickness of layer of a sprayed BIOVETRO applied onto a metal . 200 X

A Camebax 50 L electron microprobe was used to study the chemical homogeneity of the sprayed BIOVETRO ; X-ray wavelength dispersive microanalyses and area scans gave maps of the silicon, calcium and phosphorus concentrations can be visualized.

From such mapping and colour sequence it is possible to confirm that the distribution of the elements examined is reasonably homogeneous and that the presence of microcavities corresponds to the colourless zones, showing up white in the pictures in question.

These first data confirm that with plasma spraying it is not possible to obtain a continuous coating of uniform compositions as one would expect from massive BIOVETRO; this may give rise to doubts that during the spraying process the BIOVETRO may be subjected to such modifications , as a result of i ts high temperature excursion, as to alter its original properties.

These doubts were quelled by X-ray diffraction observations on powders of the sprayed BIOVETRO which confirmed the amorphous nature of the layer, free from obvious crystalline phases. However it was not possible to exclude microliquation phenomena; if the latter do exist, they are such that they do not modify the degradation process of the glass in contact with interstitial liquids.

In order to confirm that the osteoconductive properties of sprayed BIOVETRO remain unchanged, in vivo tests were carried out in the Animal Anatomy laboratories of the University of Parma. A sheet of titanium, only one surface of which had been coated with sprayed-on BIOVETRO, was inserted for 30 days into the bone tissue of a rabbit's tibia . At the end of that time, SEM observations showed a clear sequence of phases formed by the metal onto which the BIOVETRO had been deposited, followed by a zone of neoformed tissue and then the original layer of bone tissue. In the SEM photomicrograph of the interface metallbone tissue one can observe the sequence of the metal surface in contact with the neoformed tissue and the original bone tissue beyond.

Microprobe X-ray maps of the distribution of titanium. silicon and

154

calcium in the metal-BIOVETRO-bone tissue implant (fig. 3), and the Ti and Ca scans restricted to the metal-tissue interface of neoformed bone tissue (figure 4) both show the consistent presence of calcium in the neoformed tissue interfaced with the BIOVETRO .

LENGTH IN MICROHS

Figure 3. Scanning of Ti. Si and Ca concentrations in the thickness of metal-BIOVETRO-neoformed tissue-bone tissue layers as observed at the electron microprobe under the conditions reported in the text .

METAL SONE TlSSUE NEOFORMED TlSSUE > ~~~~~+----t: CI) z w ~ Z

LENGTH IN MICRONS

Figure 4. Scanning of Ti and Ca concentrations in the thickness of mE't.al-neoformed tissue-bone tissue layers as observed at the electron microprobe under the condi tions reported in the text.

155

One can presume, on the basis of the distribution of calcium ions in the strip of neoformed tissue in contact with the BIOVETRO, that the mineralization process under way can re ach a more advanced stage of development than could be expected from the tissue in direct contact with metal.

CONCLUSIONS

Certainly, the preliminary data obtained here will need additional investigations to obtain further details on the chemieal, physical and tribological characteristics of sprayed BIOVETRO. At the same time, before one can consider to have terminated these preliminary sets of investigations, it will be necessary to verify the glasses' biological characteristics, to comprehend the role of microcavities present, the mechanisms of degradation which affect the osteoconductive properties of the glass, and the criticality or otherwise of the "adhesion" property to the metallic titanium support in the presence of a biodegradable and bioabsorbable vitreous system.

Having made the appropriate reservations , one can conclude on a cautiously optimistic note that there are technological possibilities of using plasma-sprays to coat titanium with BIOVETRO, with the later retaining the same osteoconductive properties as that of its bulk form.

REFERENCES

[1] Hench, L.L., Splinter, R.J., Allen, W.C. and Greenlee, T.K., Bonding mechanism at the interface of ceramic prosthetic materials. J. Biomed. Mater. Res. Symp., 1971, 2, 117-141.

[2] Hench, L.L., Ceramic implants for humans. Advanced Ceramic Materials, 1986, 4, 306-324.

[3] Gross, U., and Strunz, V., The interface of various glasses and glass ceramies with a bony implantation bed. J. Biomed. Mater. Res., 1985, 19, 251-271.

[4] Locardi, B., Vetri e vetroceramici biofunzionali per applicazioni in chirurgia ortopedica. Riv. Staz. Sper. Vetro, 1990, 4, 163-168.

[5] Pazzaglia, U.E., Gabbi, C., Locardi, B., Zatti, G. and Cherubino, P., Study of the osteoconductive properties of bioactive glass fibers. J. Biomed. Mater. Res., 1989, 23, 1289-1297.

ACKNOWLEDGEMENTS

This work was carried out in the ambit of the research activity of the Gruppo Biovetri (GSB).

156

PLASMA SPRAY SYSTEMS FOR THE DEPOSmON OF MATERIALS FOR BIOMEDICAL APPLICATIONS

A. SALITO, G. BARBEZAT, H. FILMER, J. HOCHSTRASSER AND A.R.NICOLL

Plasma-Technik AG Rigackerstrasse 16, CH-5610 Wohlen, Switzerland

F. TROTTA Plasma-Technik AG

Via Nicolo V 13, 1-00165 Rome, Italy

ABSTRACT

Comp.0nent innovation, increased m-oductivity, reliability and lifetime are todaY's directions in engineering. Together with progress in mechanical engineerin~ and materiar technorogiesj these advances are being gained by the use OI coating technologies baseo on the concept that die substrate provides the shape and strength prpperties anp the coating the n~cess~ry surface enhal1cement. The Pfoductlon of .coatmgs for use on ~ngmeenng components m an economlc and conSlstent manner reqUlres proper eqUlpment selection in connection with the complete coating process.

Acceptance of thermal spraying has been forthcommg due to the resp'onse of the industry in meetmg biomedical requirements in solving applicational problems with re..gard to base materials and coatings.

Of the processes avaiIable plasma spray processing produces engineered layers for a variety of blOmedical applications. FOr example, hyaroxylapatjte, 1.'i and Ti02 sprayed onto Ti, Ti-6Al-4V, CoCr and plastic substrates usmg arr and vacuum plasma spraymg.

This pap'er re~orts on the status of thls process and in ~articular on the advanceö proouction features found in connection wifh individual proce~sing steps,. e.g. the quality. control . 9f pre-surface preparation techmques, powoer and plasma coatmg deposItIon.

157

INTRODUCTION

Increasing interest in the use of cementless medical implants has ensured ongoing development in techniques to optimise the surface properties. The plasma spray process is especially suitable, since the deposition of cost­effective coatings of the desired composition and structure is possible.

Plasma spraying is a line-of-sight process which involves the injection of powder into a high temperature gas plasma jet. The molten or semi­molten powder particles impact on the substrate and solidify, resulting in the formation of a coating. The coating properties (Fig. 1) are a complex function of the properties of the powder and substrate and the spray parameters. The plasma spray environment also has a decisive effect. For example, coatings produced by air plasma spraying (APS) may contain oxide phases as a result of reactions between the molten powder and the atmosphere. This effect is eliminated in vacuum plasma spraying (VPS), where spraying is done in an inert atmosphere of less than 250 mbar.

SUBSTRATE PROPERTIES POWDER PROPERTIES

Reaction to temperature Comlhosition, phases, size, changes mOf[> ology, Reaction to surface activation physlcal properties

I I SURFACE ACTIVATION POWDER MELTING

Grit blasting, ultrasonic Plasma Qarameters cleaning, transferred arc Powder feed parameters cleaning (VPS)

I I C@ATING DEPOSITION

Component s ape, movement parameters (robot)

I COATING PRQPERTIES

Microstructure, thi.ckness~ bond str~ngth, we.at: J;esistance, porosity, corroslon reslstance, blOCOmpatIblhty, etc.

Figure 1. Schematic diagram showing the most important factors which affect the properties of plasma-sprayed coatings.

158

QUALI1Y CONTROL IN THE PLASMA SPRAYlNG PROCESS

From Figure 1 it is clear that the reproducibility of plasma sprayed coatings is only possible if rigorous control is maintained over all aspects of the process. A high standard of reproducibility is characteristic of today's plasma spray technology, due to the use of computer-aided control systems and a high integrity quality control system for the raw materials and coatings.

Substrate Properties The properties of the base material must be taken into account when planning plasma spraying. The temperature rise during spraying can be limited by the use of cooling, or by allowing the part to cool between passes. These techniques can be used when spraying plastics (to prevent melting) and Ti (to prevent microstructural changes). Other aspects to bear in mind with Ti are the effect of grit blasting and coating on the fatigue strength [1] and the fact that hydrogen can not be used as plasma gas, due to embrittlement.

Substrate Surface Activation The treatment of the surface before spraying is extremely important, since it has a major effect on the adhesion or bond strength of the coating. Usually the surface is grit blasted using a high purity grit, such as pure white Al20 3• During the grit blasting, the grit breaks down gradually and must be renewed after a predetermined number of parts have been done. The grit blasting roughens the surface, thereby enhancing the mechanical adherence of the coating. The component is usually masked using a reliable method so that only those areas which are to be coated are grit blasted. The air pressure, jet angle and jet-to-workpiece distance must be kept within narrow tolerances and the entire operation kept clean to prevent contamination of the surface by foreign particles.

One of the most effective controls to monitor the grit blasting, is the measurement of the surface roughness of the grit blasted part. Table 1 shows measurements for two sizes of grit and illustrates the importance of selecting the correct grit.

159

TABLE1 Roughness of titanium plate after grit blasting with two different grit sizes

(average offour measurements)

Grit size (}lm)

63-150

600-1180

1.9

4.2

5.8

11.5

12.1

25.5

14.4

30.9

After grit blasting the parts are cleaned in an ultrasonie bath. The cleaning medium, frequency and time are kept constant. After cleaning the parts are ready for spraying. In VPS a surface are cleaning (so-eaHed "sputtering") ean also be done in the vaeuum tank just prior to spraying. All the are cleaning parameters are kept constant by computer control.

Powder Properties The powder must have a consistent composition, density, morphology and size distribution. In addition it must flow weH enough to be fed. Sampling and testing methods must be standardized [2]. In many cases international standards exist or are being developed. For example, the phase purity and ehemistry of hydroxylapatite powder eonforms to ASTM F1185-88. The grain size of eaeh powder lot is eheeked by laser light seattering methods (MICROTRAC) and is kept within narrow tolerances (Fig. 2). The morphology is eheeked in the seanning eleetron microseope (SEM) and compared to standard lots (Fig. 3).

Powder Melting and Coating Deposition An eleetrical are is generated in the plasma gun and the gas mixture is ionised by the are to form a plasma. The plasma gas is usuaHy a mixture containing several gases (Ar, He, N2 or HJ. The heat content, thermal conduetivity and velocity of the plasma can be controHed by the ehoice of plasma gas and spray parameters. In today's plasma spray systems eaeh parameter in the spraying process can be set as desired and subsequently kept constant. The plasma gun movement, for example, is controlled by a six-axis robot, allowing even complex shapes to be sprayed in a controlled, reproducible manner. This makes it possible to study the effect of varying one or more parameters, as weH as to produce coatings of consistently high quality.

160

AtvORY 6021 Laser light scattenng meaSlXements

100

~ 90

Ö 80 > 70

~ 60 .9

50 c m 40 .c .... ~

30

m 20

~ 10

0 0 20 40 60 80 100 120 140 160 180

Partlcle slze (um)

Figure 2. The size distribution of the 44-125 ~m hydroxylapatite powder (Amdry 6021).

40 lJ,m

Figure 3. The morphology of the 44-125 lJ,m hydroxylapatite powder.

X10 3 2 . 00 1.62 1.28 0 . 98 0 . 72 0 . 50 0 . 32 0 . 18 0 . 08 0 . 02

20 . 0

2 . 00 1.62 1.28 o 98 0 . 72 0 . 50 0 . 32 0 . 18 o 08 o 02

40 . 0

-,

N

25 . 0

45 . 0

161

30 0 35 . 0

Jf~ ..,

.lL . . ' 50 0

Figure 4. XRD pattern of the powder shown in Figs. 2 and 3. Symbol 0 represents CaO phase and symbol X

represents ß-tricalcium phosphate.

Coating Properties and Laboratory Testing

40 . 0

I ,

60 . 0

In addition to the tightly controlled raw materials and coating process, the standardization and documentation of laboratory test methods is equally important. All test methods, including the preparation of mounted cross sections, have been standardized, so that the properties of coatings can be meaningfully evaluated. Laboratory results are documented. The production of parts is strictly controlled by an internal quality assurance manual.

COATINGS FOR MEDICAL APPLICATIONS

Plasma-sprayed coatings have been applied to hip, knee and dental implants. Only the most commonly sprayed coatings, namely hydroxylapatite (HA) and porous titanium are discussed here.

162

Hydroxylapatite Coatings Optical micrographs of the APS and VPS coatings are shown in Figures 5 and 6. The coating properties are shown in Table 2. All coatings were sprayed on a titanium substrate.

TABLE2 Typical properties of hydroxylapatite (HA) coatings

Property VPS coating APS coating

with bondcoat without without bondcoat bondcoat

Porosity (%) 1-3 1-3 3-5 (Image analysis)

Bond strength range 50 -70 25 - 45 40 - 60 (MPa)

Roufuhness (pm) R - 9 R - 8 R - 9 1Sty us tester) Ra - 22 Ra - 21 Ra - 25 our measurements RP - 44 RP - 44 RP - 56

each, on four sampies RZ - 53 RZ - 54 RZ

- 65 t t t

The phase purity (Figs. 7 and 8) of the coatings is measured by X-ray diffraction (XRD) using the standard file card [3] for crystalline HA. Nevertheless, the results are only an estimate since a standard for pure ß­Ca3(P04)Z is lacking. It is more meaningful to compare spectra than to attempt absolute measurements.

A correlation between the degree of powder melting during spraying and the percentage crystallinity of the coating has been observed. Ideally, the powder particles should only undergo surface melting, in order to bond effectively in the coating, but should not be completely molten.

Porous Titanium Coatings These coatings (Fig. 9) are usually produced by VPS using Amdry 918-2 powder and have a high porosity (ca. 40-50%) and roughness (Ra - 33 ]lm, Rp-85 ]lm, Rz-212]lm and Rt -350 ]lm). Typical pore size distribution is 90% in the range 0-50 ]lm, 7% in the range 50-100 ]lm, 7% in the range 50-100]lm and 3% larger that 100 ]lm.

163

50 }ol m

Figure 5. Microstructure of an APS hydroxylapatite coating.

Xl0 3 2.00 1.62 1. 28 0 . 98 0 . 72 0 . 50 0 . 32 0 . 18 0 .08 0 . 02

20 .0

2 . 00 1.62 1.28 0 . 98 0 . 72 0.50 ~ 0 . 32 0.18 0 . 08 0 . 02

40 . 0

25 . 0 30 . 0 35 0

45 . 0 50 . 0 55.0

40 . 0

60.0

Figure 6. XRD pattern of an APS HA coating, produced with Amdry 6021 powder.

164

50 }.1 m

Figure 7. Microstructure of a ~S hydroxylapatite coating with Tl bondcoat.

Xl0 3 2 00 1.62 1 28 0.98 0.72 0.50 0 . 32 0.18 0.08 ............. .",

0.02

20.0

2.00 1.62 1. 28 0.98 0.72 0.50 0 . 32 0.18 0.08 0.02

40.0

25 0

45.0

30 0 35.0

. :1

J.t-:: 50 0 55.0

Figure 8. XRD pattern of a VPS hydroxylapatite coating.

'" .. ..

40.0

~

60.0

165

100 }Im

Figure 9. Microstrueture of a porous titanium eoating (VPS). The eoating thiekness is usually 200-300 }Im.

CONCLUSIONS

HA, Ti and Ti02 eoatings produeed by air and vaeuum plasma spraying have been optimised to a large extent although further work is envisaged. The necessary quality eontrol struetures for production are in place.

1.

2.

3.

REFERENCES

W. Winkler-Gniewek, H. Stallforth, M. Ungethüm and H. Gruner: Strueture and PrQPerties of VPS Coatings in Medical Teehnology. In ~ 19 PI~mfTbehnik Symposiullk Vol. 3, ed. H. Esehnauer et al., Plasma-Tee ni A , Switzerland, 19ö9, pp. 95-102

A.R. Nicoll, Retestin§ Plasma ~ Powders, Plasma-Technik AG, 1986, Publieation No. 6005E

~tawar\i ~ecifieation fur Comfi.&8ioQ Qf ~r~mi§ H~droxylapatite IQ!: _ urgtea Implants, ASTM:" 19 , eSlgnatlon 11 5-8

166

BIOCERAHICS FOR MAXILLOFACIAL APPLICATIONS.

J .G. C. WOLKE·, C.P.A.T. KLEIN·" and K. OE GROOT·. Biomaterials Research Group, School of Medicine, University of Leiden, Rijnsburgerweg 10, 2333 AA Leiden, The Netherlands.

* :Oepartment of Biomaterials, School of Medicine, University of Leiden, The Netherlands.

# :Oepartment of Orale Implantology, ACTA-Vrije Universiteit, Amsterdam, The Netherlands.

1. Introduction.

Surgical implants to repair or augment parts of the skeleton (bone, teeth, joints) can be produced from a number of materials, among which ceramies play an important role. There are two types of these specialty ceramies or bioceramies: bioactive ceramies and bioinert ones. Bioactive ceramies are all bioceramies that have at least a surface structure resembling that of the mineral phase of the skeleton, i.e. calciumphosphate. All other bioceramies (such as aluminumoxide) are bioinert. Oue to their similarity with (the mineral phase of) bone, bioactive ceramies bonds to bone in a natural way. This property renders them very attractive for surgical application. Unfortunately, mechanically they are very weak, especially in tension, the consequence being that clinical application of bulk bioactive ceramics is limited to sites in the human body where no tensile forces are present. Another disadvantage was that sometimes needed porosities, to simulate porous bone, so-called spongious bone, are difficult to introduce. with respect to apatite coatings, we have learned a number of important facts. First of all, contrary to our initial expectations apatite coatings are not biostable when implanted, but they degrade visibly within one year. Secondly, not less important, we showed that a coating is most important during the healing phase of the wound in which the implant is inserted. This means, that it is not necessary for a coating to be permanently present after wound healing and bony ingrowth has occurred. We have completely characterized the coatings in terms of crystallographic structure, bonding strength, thickness, and in

167

vitro and in vive degradation behaviour. In addition the sequence of bone reaction towards such coatings has been delineated. Furthermore, we have critically evaluated so-called push-out strength measurement of the bone-implant interface, both from an experimental point of view as weIl with the aid of finite elemental analysis. Animal studies have confirmed, that the widely used "push-out" test results are not representative for the true tensile strength of a coating und er in vive conditions. Biological effect have been shown by crystallographic changes after implantation. Most important, clinical results proved the coatings to enhance the performance of tooth implants and artificial hip joints dramatically. It is expected that in the coming years throughout the world the majority of tooth- , hip- , and possibly other bony implants will be coated with hydroxylapatite.

2. Chemistry and physics of calcium phosphate ceramics. l

The current knowledge of calcium phosphate bioceramics is best summarized in Figures 1-4. Figure 1 shows the possible phases of CaO/P20S mixtures. At temperature Tl there exists the equilibrium for hydroxyapatite (HA): CalO (P04 )6(OHh (or Ap) = 2Ca3 (P04 h (or Ca3P) + Ca4P20 9 (or Ca4P) + x IH20. Temperature T2 separates the phase HA + CaO from the phase HA + C4P. Temperatures Tl and T2 depend on partial water pressure, as figure 2 shows. Figure 2 shows the influence of partial vapour pressure (pH20) on the stability of various calcium phosphates as a function of temperature. The temperature T2 (see also Fig. 1), representing the equilibrium HA + CaO = HA + Ca4P - H20, decreases with increasing water pressure. The temperature Tl (see also Fig. 1) shows that at a given temperature, for example = 1250 0 C a variety of coexisting phase may exist: if log pH20 (mm Hg) <0, the phases o:C3P + C4P are stable: if 0< log pH20 (mm Hg) <1, the phases HA - C4P are stable. Figures 3 & 4 only differ in that the phases without water vapour (Fig. 3) are quite different from those with a vapour pressure (Fig. 4). It is obvious that control over temperature, Ca P ratio and vapour pressure during sintering gives one the ability to produce a wide range of well-defined calcium phosphate products. Even more important: if these parameters are not all weIl defined a less well-defined end product may result.

3. Plasma-sprayed coatings of calcium phosphate.

Figure 5. shows schematically the principle of plasma spraying. A DC electric arc is struck between two electrodes, while a stream of (mixed) gases passes through this arc. This results in an ionized gas of high temperature up to 30000°C, with a high-

168

speed approaching the speed of sound due to the large expansion resulting from this temperature increase. Hydroxylapatite powder is suspended in carrier gas stream which is fed into the plasma flame. Obviously, apatite is thermodynamically unstable at plasma spray temperatures (as discussed in the previous paragraph). Ideally, only a thin outer layer of each powder particle gets into the mol ten plastic state which unavoidably undergoes phase transitions. This plastic state is necessary to ensure dense and adhesi ve coatings. By choosing an optimum relation between particle size , type of gas (the he at content of a plasma, and thus the ability to increase the temperature of a particle, depends strongly on the gas used) , speed of the plasma (the longer a particle resides in a plasma, the higher its temperature), and cooling process of the coated surface, one obtains coatings with the desired calcium phosphate(s) and crystallinity.2 Hydroxylapatite coatings are most of the time a compromise between an amorphous and a amorphous crystalline phase. Figure 6 shows the influence of different plasmagases on the crystallinity of the coating. Plasmagas Argon mixed with Hydrogen gives a higher degree of crystallinity but without Hydrogen the powder particles cannot enter the gas. This is because the high velocity and viscosity of the Argon gas cause the particles to bounce back from the flame, instead of entering it. Figure 7 shows that the use of Nitrogen as the plasmagas gives a thicker coatinglayer compared to Argon. A important criteria is the formation of calciumoxide , when this is too much it can react with water and destroys the coating layer. Figure 8 shows that a higher content of Hydrogen results in more Calciumoxide than Nitrogen alone. This is because Hydrogen gives more enthalpy to the flame and a lower velocity of the flame so the particle undergoes more melting and decomposing. The best conditions for spraying HA is to use pure nitrogen as plasmagas.

3.1. X-ray diffractrometry.

We will discuss one other physicochemical aspect, namely the crystalline structure of the coatings. 3 However the crystallinity of the coating always differs from that of the powders fed into the plasma flame: ß-TCP turned completely into a-TCP, HA and FA showed an appreciable line-broadening of the peaks. These results are understandable in terms of the thermal stability: the transition temperature of ß-TCP is around 1200 o C, HA decomposes at 1300 o C, and FA is stable above 1300 0 C. Hence, the capacity of the plasma flame to induce phase transitions decreases in this order.

For plasma-spraying a Metco system was used. The spray conditions were:

169

Are current Are gas Powder gas Gun speed Powder transport Cooling

400 A nitrogen nitrogen 10 cm/sec 10 - 15 gr/min air

Prior to the plasmaspraying all sampIes were sandblasted, followed by ultrasonic cleaning and drying at 100°C. Figure 9 shows the X-ray diffraction pattern of sintered HA starting powder. Hardly any line-broadening can be seen, which indicates a weIl crystallized material. Figure 10 shows the XRD pattern of HA-125 sprayed with Nitrogen as the plasmagas.lt can be seen that plasma-spraying hardly alters the crystallographic structure, only some line-broadening is visible. Figure 11 shows the XRD patterns of HA-45, the structure of the coating consists almost entirely of an amorphous phase. We believe this to be caused by complete melting of the small particles. This also shows that melting of particles result in amorphous coating. However, figure 12, a FA coating sprayed with the same particle size distribution gives a crl,stalline coating comparable with the 1-125 micron HA coating. It shows clear that in this ca se plasmaspraying hardly alters the crystallographic structure, only the degree of crystallinity is decreased as compared with the X-ray of the starting powder and the ratio of the intensity of the peaks is changed due to the orientation of the crystals during the cooling process.( peak 25.7° 20 is increased) So it is clear that the temperature behaviour of FA is superior to HA. Figure 13 shows the XRD pattern of HA-45 after a heat-treatment for 1 hour at 600°C in an inert furnace to prevent oxidation of the metal.This heat-treatment increases the crystallinity and no amorphous phase is present any more. Figure 14 shows the XRD pattern of HA-125 after incubation in vitro for 3 months in Michaelis Buffer pH 7.2 at 37°C. One can observe an increase in crystallinity after this incubation time and no amorphous phase is present. This increase in crystallinity is also observed seen in an in­vivo experiment (Figure 15). A coated plate has been implanted in the subcutaneous tissue of a rabbit and after 6 weeks the XRD diffraction showed an increase in crystallinity and disappearance of the amorphous phase.

3.2. Tensile bond strength.

The coating thickness of the tensile specimens is 45 +/- 5 microns. The results are:

Strength Failure mode n Araldite AV 118 60.8 MPa 100% Glue/HA 5 3M 38.3 MPa 100% Glue/HA 5 Lee Insta-Bond 5.4 MPa 100% HA/Ti 3 Concise 11. 7 MPa 100% HA/Ti 3

170

By various coating thickness the tensile strength glued with Araldite AV 118 are:

strength Failure mode n 40 microns 66.8 MPA 100% Glue/HA 5 80 microns 60.7 MPa 100% Glue/HA 5

120 microns 45.3 MPa 100% HAITi 2 100% Glue/HA 1

An important phenomenon is that the orthodontic adhesives are able to pull off the coating from the titanium although they have a very low tensile strength, and while the epoxy adhesives have very high tensile strengths no coating is removed from the metal. An explanation for the results of the orthodontic adhesives could be that the liquid in the adhesive penetrates through the coating layer and destroys the physical bonds while setting. The epoxy adhesives gives a crack propagation in the interface between epoxy-adhesive and coating. So these experiments show that different adhesives give different failure mode and tensile strengths. Thus we must be careful to compare data from different laboratories, if we are not sure about the experimental conditions.

3.3. Push-out test.

In literature many investigators published various results for HA coatings. To analyze the biomechanical characteristics of this test, we performed a finite element analysis and calculated the interface stress distributions for varying the following parameters. s Figure 16 shows the parameters we examined:

x - Clearance of the hole in the support jig. Y - Young's modulus of the implant. t - cortical thickness. d - implant diameter.

Figure 17 shows the effect of the clearance of the hole in the support jig. The lines indicate the stress distributions during the push-out test. A small clearance results in high stresses at the site where the jig edge support the bone. The best result gives a clearance of 0.7 mm, which gives the most uniform distribution of stresses along the interface. The worst result is with a clearance around 0.1 mm or less. In literature, many researches have reported to have minimized their clearance. This is not a good approach and will lead to unrealistically low apparent push-out strength values. Varying the Young's modulus of the implant show again that the interfaces stress distribution is in most situations not uniform (figure 18). A low Young's modulus results in considerably higher stresses at the medial site of the cortex, from which the load is applied on the implant. But high modulus results in slightly higher stresses at the lateral site of the cortex where the jig edge supports the bone. Only in this situation the interface stress distribution is much more uniform compared to the situation of a low Young's modulus of the implant. Because in most

171

situations, the Young's modulus of the implant cannot be standardized we can conclude that it is not allowed to compare materials with different Young's moduli with each other in push­out test.The cortical thickness and implant diameter did hardly influence the interface stress distribution.

3.4. Biologieal properties.

The purpose of this study was to compare, in a transcortical push-out model in the femur of dogs, at 5 and 25 months, the interface behaviour of plasma sprayed hydroxyl apatite and grid­blasted Ti-6Al-4V titanium alloy implants. Figure 21 shows a BSE imaging analysis of a HA coating after 5 months implantation, with special attention to the fracture site from a push-out test. The bone apposi tion is good and the fracture is located at the titanium-HA interface, so the bone­coating interface is stronger than the coating titanium interface. Figure 20 shows the section of a HA coating implants at 25 months showing good bone apposition , and no coating left. Figure 21 shows a titanium implant after 25 months implantation, giving a fibrous tissue interlayer and there is no bone apposition against the implant. Why a HA coating, figure 22 shows clearly that titanium plugs implanted in the femur of a dog, have a lower push-out strength than a HA coated implant. Especially in the beginning (3 months) HA shows a higher bonding, because of the coating. The coating will degrade, which leads probably to a small decrease of the push-out data. When the coating will be disappeared the bone formation will adapted to this new situation and the bone bonding increased. Titanium shows always lower results probably of the fibrous layer and fewer bone contact.

4. Summary of the results.

with respect to apatite, we have learned a number of important facts. The nature of the plasmaflame gives a more or less amorphous phase, a very hot flame gives not only an amorphous coating but also formation of calciumoxide. Nitrogen as a plasmaflame, being less hot, results in a higher crystallinity of the coating layer. A HA coating sprayed with a particle distribution of 1-45 microns gives an entirely amorphous phase. The thermostability of fluorapatite at elevated temperature is superior to hydroxylapatite and produces a FA coating with a higher degree of crystallinity compared to a HA coating. A heat treatment afterwards increases the crystallinity. After incubation in vivo or in vitro the amorphous phase decreased. Various adhesives show different tensile strengths and failure modes of the coating. A finite element analysis of the so called push-out test gives

172

the best result with a clearance of 0.7 mm, which gives the most uniform distribution of stresses along the interface and it is not allowed to compare materials with different Young's moduli with each other in push-out test. After 2 year implantations in the femur of a dog HA coatings show always higher push-out strength compared to titanium implants.

S. Literatur ••

1. Groot de, K., Ann. N.Y. Acad. Sci.,523, 227-233 (1988) 2. Groot de, K., C.P.A.T. Klein, J.G.C. Wolke, J.M.A. de

Blieck-Hogervorst.: Chemistry of Calcium Phosphate Bioceramies, Handbook of Bioactive Ceramies, Vol. II, Ed. Yamamuro et al, C.R.C. Press, 1990, p. 3-17.

3. Koch, B., J.G.C. Wolke, K. de Groot.: X-ray diffraction studies on plasmasprayed calcium phosphate coated implants, J. Biomed. Mat. Res., 24, 655-667 (1990).

4. Dhert W.J.A., C.P.A.T. Klein, J.G.C. Wolke, E.A. van der Velde, K. de Groot, and P.M. Rozing, "A mechanical investigation of fluorapatite, magnesiumwhitlockite and hydroxylapatite plasma-sprayed coatings in goats, Accepted for publication by J. Biomed. Mater. Res., (1991).

5. Dhert W.J.A., C.C.P.M. Verheyen , P.M. Rozing, K. de Groot, Finite element analysis of the push-out test, Accepted for publication by J. Biomed. Mater. Res., (1991).

ClP+ C"P

TI CaO + C.p C.p C,P C3? ..

Ap + + Ap LIQ

T,

-C,?

CoO+ Ap + Cz?

Cao

Figure 1: Phase diagram cf CaO/P20S mixtures .

1700

1600

1500

1400

1300 CoO+Ap

Q.

13600 '! Cl. .. U CI

aC5P +

C2P 1200~ ____ ~ __ ~ __ ~~~ __ ~~

70 65 C,.P Ap 55C~ 50

-Wt%CoO

Figure 3: Phase diagram cf cao/p2oS mixtures .

173

T or , ~ 16eo 1500 1400 1:300

5.25 5.50 5.75 6.00 6.25 6.50

104'T, K

Figure 2: Phase diagram cf caO/P2oS mixtures shcwing the influence cf water pressure .

1700

1600

1500

1400 CoO + c.p

1300

-Wt"10 CoO

Figure 4: Phase diagram cf CaO/P20S mixtures . waterpressure 500 mm Hg.

174

~. --., • _' 0 __ 0 _ 0 =-:J

Figure 5: View of a plasma gun.

Figure 6: The influence of different plasmagases on the crystallinity of the coating.

Figure 7: The influence of different plasmagases on the thickness of the coating.

Figure 8: The influence of different plasmagases on the formation of Calciumoxide in the coating.

loo ;:.C1'.!:Y'I:;:laI=:'::::'nI:.:� !..Y ;;;ln:..:"=---_____________ -,

.. ..... -_ .... -_ .......... __ ...... ---

110

40

20

o~----~----~----~----~----~ o 2 4 e e ~

110

100

50

o

0" .. ' 2

,. HydfOQlln

- Hltrooen ... Nvon

000

00'

..... ..... _ ... -_ ... _ .. -- ........ 000

• ,. Hydrogen

e

Z5;"~C=.~~I=~==.=~~. ________ _________ _,

2 e ,. Hydrogen

-- IIU_" -+ . AI9O"

• ~

175

~ ~l :r r 1 : , F -- -1 1 11 -- ---'I--i • . 11 11 I--f----

, I I 1 1 " I I I 1 I 1_ f- i

I' .,.;, .. , .... ",. I 1 I [' I' -"'1~ .. r' . l I ._ .. _ .. ,. ... I I I

I 1-1 ,I. , I 'I I I , 1 •

1 I I" 1 L· ,I _I, '-- ,- ---l-

I I 1 I

I i ~l I·· ---1-1--1-

I I i .1" I I J 1 . • I·r --

• i!I \..'~ i '~~~~;l~ I I ~I ~. 1 1 I r' I, . 'r

Figure 9:XRD pattern of sintered Figure 10:XRD pattern of HA125 HA starting powder.(25-400 20) plasma Nitrogen .

. nMofG,... .. C'OOIIIICI powdet tJ. 1 ~ 41 .kfO .. tw. mag_ UD lc~

i 11 1

i 1 r-1 • J .

Figure l1:XRD pattern of HA45, Figure 12:XRD pattern of FA45 plasma Nitrogen. plasma Nitrogen.

. ...,... ........... , ... _0c". ~ . .. UJ

Figure 13:XRD pattern of HA45 treated for 1 hour at 600°C.

:: :

.................. - -­_1_'''' La. U

176

~ I 1-

Figure 14:XRD pattern of HA125 after incubation for 3 months in Michaelis Buffer pH 7.2

Figure 15:XRD pattern of HA125 after in-vivo experiment for 6 weeks in subcutaneous tissue of a rabbit.

177

F

+ + + _d_

C

t I Y

I

J J

Figure 16: Schematic drawing of the push-out test.

16 r

0 xld-O 02

o 3.xld-O 06 12 ...

;) o 5. '/d-O

'" " a. ::i:

· r •

~

o 2 3

y - coordinate along Interlace (mm)

E - ll0GPa 1- 3 mm d - 5 mm

Figure 17: Interface stress distribution for the implant-support jig distance (x) between 0.1 and 1.0 mm

178

2

I. r ~ 5

..J 0

.. 25 12

• 50

Gi' 9 100 c.. ~ 8 0 200

i= 0 400

" 4

o ~------------~------------~------------~ o 2 3

y - coordinate alon9 interface (mm)

x - 07 mm 1- 3 mm d - 5 mm

Figure 18: Interface stress distribution for the Young's moduli (E) varying between 2 and 400 GPa.

Figure 19: BSE imaging analysis of a HA coating after 5 months implantation

179

Figure 20: aSE imaging analysis of a HA coating implants at 25 months implantation.

Figure 21: aSE imaging analysis of a titanium implant after 25 months implantation.

180

MPa 50

40

30

20 / 10

-------=---------oL-______ L-______ L-____ ~L_ ____ ~~ ____ ~ ______ ~

o 5 10 15

months 20

-e- HA coatlng -- Tltanlum

25 30

Figure 22: Push-out strength of titanium and HA coated plugs implanted in the femur of a dog.

181

NUCLEATION AND GROWTH OF DICALCIUM PHOSPHATE

DIHYDRATE ON TITANIUM ALLOY SUBSTRATES.

PASCALE ROYER, MICHELE FRECHE, CHRISTIAN REY

Laboratoire des Materiaux, Physico-Chimie des Solides

38, rue des 36 Ponts; 31400 Toulouse, FRANCE.

ABSTRACT

The formation of calcium phosphate deposits on titanium in supersaturated solutions has been

studied. A constant composition crystal growth method is described. Several coatings have been made using

this technique and have been investigated by Scanning eleclron microscopy, Fourier transformed infrared

spectroscopy, X ray diffraction. Besides, the growth rate of different surfaces have been studied.

INTRODUeTION

Apatitic calcium phosphate coatings on orthopaedic prostheses give them bioactive

properties and allow to suppress the use of orthopaedic cements. Several coating methods

have been proposed (electrodeposition, pulverisation, Dip coating ... )(1-3) but most

developed is the plasma spraying technique. However this method presents several

drawbacks and limits [decomposition ofhydroxyapatite in the plasma (4), inability to have

homogeneous films thiner than 20-50 Ilm, inability to depose carbonateapatite closer from

bone mineral than hydroxyapatite or other calcium phosphate cons~dered as bone precursors

(5)].

In order to extend the possibilities of calcium phosphate coatings, we propose to

develop two new methods of deposition in wh ich bioactive calcium phosphate salts such as

apatite, dicalcium phosphate dihydrate (DCPD), octacalcium phosphate, amorphous calcium

phosphate (6) would directly nucleate and grow on the titanium alloy substrate:

182

- The first method of deposition, previously presented (7) consists of an induction

of calcium phosphate precipitation (amorphous or apatitic) on a titanium cathode by

electrolysis of a supersaturated calcium phosphate solution.

- The second method is a direct crystal growth. In this method the metal surface

plays an active role unlike in other processes already proposed. We focus in this first report

on the nucleation and growth of DCPD by the second method. This compound has been

chosen essentially because of its high growth rate. Besides, DCPD could reasonably be the

first neoformed mineral deposit, in vivo, at the surface of implants in slightly acidic

conditions resulting from surgical trauma.

METHOD AND MATERIALS

The method consists of the nucleation and growth of calcium phosphate crystals on

different seed materials from metastable, supersaturated solutions of calcium phosphate at

constant composition. This method has been described by Nancollas and co-workers (8) and

it has been widely used for testing crystal growth inhibitors, nucleation properties of calcium

salts and other compounds.

When seeds are added to a supersaturated solution, they may induce calcium

phosphate formation depending on their potentiality to initiate the formation of first nucleL

The nucleation and growth of crystals consume ions from the solution and lead to a change

of pR which monitors the addition of reactifs in order to maintain constant the pR, the

concentration of calcium and phosphate ions and the ionic strength of the solution.

The crystal growth experiments were made in a water jacketed double walled pyrex

cell at 37°C. The stable calcium phosphate supersaturated solutions were prepared at the

required pR, by the addition of O.IM potassium hydroxyde. When the seed material was

added, as soon as the growth of the calcium phosphate occurred, a pR decrease by as little

as 0.01, triggered the simultaneous addition of two calcium and phosphate solutions from

mechanically coupled, automatie burets. The concentrations of the solutions were such that

their addition exactly compensated for the changes in anion and cation concentrations arising

from precipitation. The concentrations of the ionic species in the solution were computed

from mass balance, electroneutrality, dissociation constants of phosphoric acid and ion pairs

by making activity corrections due to the ionic strength (9). Potassium chloride, an inert

electrolyte, was added to the supersaturared solution and the reagents to insure the constancy

of the ionic strength. Experiments were realized for atomic Ca/P = 1, constant ionic strength

(IS = 0.12 moLl-I), constant pH (= 4.9) and temperature (T = 37°C).

183

Crystal growth experiments have been perfonned on bulky sampIes with different

surface shapes: Polished and corundon gravelled TA6V4 pellets, electrochemically coated

TA6V4 surfaces and pure hydroxyapatite sintered pellets. For all sampIes, the surface area

exposed to the solution was 59.4 mm2.

Several coatings have been made using these two methods and characterized by X­

ray diffraction, infrared spectroscopy and scanning electron microscopy.

RESULTS

DCPD fonnation has been attested on all surfaces by X-ray diffraction and infrared

spectroscopy.

The kinetic curves obtained on the different pellets are presented in figure 1. They show an induction period 't, during which no calcium phosphate fonnation could be

detected. This period corresponds to the time taken for DCPD nuclei to be fonned on the

foreign surfaces. After the induction period, the DCPD crystals grow at variable rate R defined as R= ~v.C / ~t.S where ~v is the volume of reactifs added during the time ~t,

C=~m/~v is the effective titrant concentrations and S the surface area of the sampIe,

exposed to the solution. For the detennination of R, we have considered the part of the

kinetic curve corresponding to a growth less than 3.10-5 mole of DCPD.

3,-------------------------------------, v (mn

2

(minute)

80

Figure 1. Kinetic curves of growth of DCPD on foreign seed pellets.

v I~, ...... IIAr

.PoI ... TA6V4 .... "ti .1~"CKh ... _hallJ .-.... .. hit cakI ... ' ..... ph ••• c.."....., •••• ,.",.

184

The different induction times determined from the curves and the different calculated

growth rate are summarized in table I. Induction times were found to be very short in all

cases; not exceeding a few minutes. Although the calcium phosphate seed surfaces present,

as expected. shorter induction times, dicalcium phosphate dihydrate seems to nucleate easily

on the titanium alloy surfaces. The induction time seems to be shorter on the gravelled pellet

which probably offers a larger number of nucleation sites, due to his rougthness. However,

when considering the growth rate, it appears that DCPD grows quicker on polished T A6V 4

surfaces than on gravelled ones. The calcium phosphate-carbonate coated surface presents

the most interessant growth rate whereas the DCPD growth rate seems weak on pure

hydroxyapatite sintered pellets.

TABLE 1

Constant composition growth ofDCPD on titanium alloy seed pellets (*).

Polished T A6V 4

Sintered HAP Polished TA6V4 Sanded TA6V4 + electrochemical

calcium phosphate

carbonate coating

't (min) 2 6 4 3

R.1O-8 1.48 1.33 0.87 1.96

(mol.min-1.mm-2)

------------* [Ca] = [P] = 19 mM; Total initial volume = 100 ml; T = 37°C; IS = 0.12 mM; pH = 4.9; C = 36mM.

The morphology of the DCPD crystals on polished and gravelled TA6V 4 surfaces is

shown on figures 2 (a) and (b). Natural DCPD crystals fonn large flattened plates which, in

an cases, can be observed on the metal surfaces. For polished TA6V4 pellet, DCPD crystals

are more numerous and appear to overlap whereas, on the gravelled pellet, single dispersed

crystals are observed quite parallel to the surface of the pellet.

In order to increase the coating thickness, several cycles of growth experiments have

been realized. Morphology of the coating obtained after 4 cycles is shown on figures 2 (c)

and (d). A homogeneous highly porous coating is observed.

185 . a . . b .

xSoo

. c . . d .

Figure 2. Morphology of growed DCPD crystals on different surfaces (a) polished surface

(b) gravelled surface (c) after 4 cycles on polished surface (d) optical microscopy.

DISCUSSION

The surface of a permanent implant has two essential functions : anchoring to bone

tissue and insuring stable bone implant interface. In addition, however, it might promote

bone tissue formation during the weeks following implantation. The deposits which were

realized by one or the other of the two methods have poor mechanical propenies and are

only weakly bound to the metal surface; they are bioresorbable and they could, only playa

pan in bane ingrowth processes specially where plasma spraying technique is unusable

(internal Fibermesh surfaces, grids, porous implants ... ). Instead of the electrochemical

186

deposition which cannot deposit layers thicker than a few microns, the constant composition

method can, potentially, build layers as thick as desired. The deposition rates are quite fast

for brushite and they depend on the surpersaturation. The efficiency of brushite as

biomaterial has been recendy demonstrated (6). It appears an excellent apatite progenitor and

has been found to be more effective than stoechiometric hydroxyapatite in bone

reconstruction incalvaria of rats. It shall be emphavoized that the crystal growth method

which we evaluate is quite adaptable and might be used as weIl for growing other calcium

phosphates including hydroxyapatites provided that the surface has some propensity for

nucleating the chosen calcium phosphate in the domain of stability of supersaturated

solutions. It must be noticed however that the other calcium phosphate salts have lower

growth rates than brushite and that the development of layers will not probably be easy. The

development of the crystal growth method for the realization of calcium phosphate coating

necessitate frrst the obtention of a surface with high nucleation sites, which would improve

the interfacial attachment of the crystal layer to the metal surface and give high density

deposits. In addition to physical modification, such as gravelling, the multiplication of

nucleation sites on titanium surfaces may be attempted by chemical modifications

(nitruration,phosphatation, ... ) and these studies are now in progress.

The data which were obtained show that all potential biomaterials which have been

tested, have the capacity to nucleate brushite at medium supersaturation rates. Titanium

especially, appears as a good nucleator of brushite as it has been found for apatite (10).

Heterogeneous nucleation is generally related to the existence of a epitaxial relationship

between the nucleating surface and the nuclei : if the mismatch between the atomic pattern on

the surface and an atomic plane of substance to grow is minimal, nuclei may form and grow

on the surface (11). In case of titanium surface, or more precisely, titanium oxyde layer,

such relationships have not been demonstrated. The finding that polished titanium surfaces

are more efficient than gravelled surfaces for growing brushite crystals shows that physical

parameters have to be taken in consideration even in crystal growth behavior. These

treatments produce in fact surfaces with different textures, and the development of large

brushite crystals seems to be facilitated on smooth surfaces whereas it seems limited on

gravelled surfaces by surface irregularities. The stoichiometric hydroxyapatite pellets did

promote very efficient nucleation of brushite which can be explained by the existence of

epitaxial relationships (11) ; however, the growth rate is lower than for non-stoichiometric

apatite. This surprising result might be due in part to differences in surface composition. It

has been shown in labelIed experiment involving Ca exchanges that high temperature

hydroxyapatites do not have a surface equilibrated with the solution (12). The equilibration

of the apatite surface might then delay the growth of brushite nuclei and could explain these

unexpected results. On the contrary, with precipitated non-stoichiometric carbonate apatite

187

no equilibration delay are expected and this sampie seems to be the most efficient substrate

for brushite growth.

The constant composition crystal growth method offers, however, another interest as

a fast test technique for optimizing nucleating surfaces. In all instances bioactivity has been

related to the propensity of a biomaterial for promoting calcium phosphate formation on its

surface: these phenomena have been weH described in the case ofbioglasses (13) as weH as

in that of hydroxyapatite (14-15), and more recendy bioactive polymers. The optimized

biomaterial could be indeed a material which would have the capacity to nucleate calcium

phosphate at supersaturation levels existing in biological media and would keep this capacity

on aging and even after calcium phosphate resorption. Such materials would be naturaHy

integrated in the bone tissue and the formation, in situ, of calcium phosphate dense coating

would even prevent the material from releasing potentiaHy toxic ions, or substances. These

materials could be produced by stahle surface modifications leading to dense nucleation sites

and low supersaturation level of nuc1eation. Of course, the growth of large brushite crystals

at pR 4.9 might not be the best test; other studies on growth of different calcium salts

suggested as bone mineral precursors have still to be done in order to better understand the

nucleation and growth mechanisms, and then bioactivity of skeletal implants surface.

Acknowledgment : We thank BIOLAND company for the financial support of this study.

REFERENCES

1 . P. Ducheyne, W. Van Raemdonck, I.C. Heughebaert, M. Heughebaert, Structural analysis of

hydroxyapatite coatings on titanium. Biomaterials, 1986,7,97-103.

2 . H. Solnick-Legg, K. Legg, Ion beam and plasma technology for improved biocompatibles surfaces.

MRS Bulletin, 1989, 14, 27-30.

3 . M. Freche, R. Morancho, G. Constant, Mise au point de l'elaboration de couches minces de

phosphate de calcium apatitique par double dccomposition. Ann.Chim Fr., 1985, 10,549-559.

4 . K. Oe Groot, Effect of Porosity and Physicochemical Properties on the stability, Resorption, and

Strength ofCalcium Phosphate Ceramics, BIOCERAMICS ; Material Characteristics Versus In Vjvo

Behavjor. Annals of the New York Academy of Sciences. vol. 523, 227-233.

5 . M. J. Glimcher, The nature of the mineral component of bone and the mechanism of calcification. Ed

Griffin PP, American Academy of OrthoIlaedic Surgeons Instructional Course Lectures XXXVI,

1987,49-69.6.

6. O. Suzuki, M. Nakamura, Y. Miyasaka, M. Kagayama, M. Sakurai, Bone formation on synthetic

precursors of hydroxyapatite. Tohoku J. EXIl. Med .. 1991. In press.

188

7 . P. Royer, C. Rey, Calcium phosphate coatings for orthopaedic uses. Surface and coatings technolQgy,

1991. In press.

8. P. Koutsoukos, Z. Amjad, G.H. Nancollas, Crystallization of calcium phosphates. A constant

composition study. J. Am Chem Soc., 1980, 102, 1553-1557.

9 . C.W. Davies, Ion association (Butterworths, London), 1960,41.

10 . T. Hanawa, Titanium and its oxide film: a substrate for formation of apatite. Proceeding of " The

hone-biomaterial interface workshop", Toronto, Canada, 1990, Dec. 3-4 (in press).

11. P.G. Koutsoukos, G.H. Nancollas, CrystaI growth of calcium phosphates-Epitaxial considerations.

Journal ofcrystal growth, 1981,53,10-19.

12. A. Skoubani, C. Rey, M.J. Fauran, G. Bone!, Chemical evaluation of calcium phosphate

Biomaterials Interface in Aqueous Media. Proceedings of the 7th CIMTEC, Montecatini, Italy, 1990,

July 2-5 (in press).

13. L Hench, Bioactive Ceramies, Ann N.Y. Acad. Sei., 1988,253,54-71.

14. M. Heughebaert, R.Z. Legeros, M. Gineste, A. Guilhem, Hydroxyapatite cerarnics implanted in non­

hone-forming sites. Physico-chemical characterization. J Biomed Mat Res .. 1988,22,257-268.

15. N. Bonsfield, Z.B. Luklinska, High resolution electron microscopy of a hone implant interface.

Proceeding of "The hone-biomaterial interface workshQP", Toronto, Canada, 1990, Dec. 3-4 (in press).

189

Relationships between bulk and surface structure and biomaterial biocompatibility

A. Bertoluzza1, M.A. Morelli2, A. Tinti1

Centro di Studio Interfacolta' sulla Spettroscopia Raman, Universita' di Bologna (ltaly)

lSezione di Chimica e Prop. Biochimica, Dipartimento di Biochimica, Universita' di Bologna, via Selmi 2, Bologna (Italy)

2Dipartimento Chimico "G. Ciamician", Universita' di Bologna, via Selmi 2, Bologna (ltaly)

Abstract

The different aspects (physical, chemical and biochemical) influencing the biocompatibility of abiomaterial for prosthetic use are analyzed and discussed in relation to the bulk and surface structure of the material and to biomaterial-tissue molecular interactions.

Also the actual strategies ("new" and "composite" biomaterials) are analyzed considering the effects involved in obtaining optimum conditions as regards to the biocompatibility.

Finally, the synthesis of calcium and calcium-aluminium oligophosphate glasses with a controlled structure for orthopaedic use on the one hand and, on the other, the possibility to give stability to the surface water layer of hydrophilic contact lenses based on organic polymers for ophthalmological use are discussed as examples of the concepts enunciated in the work.

Introduction

In the study of the relationships between bulk and surface structure and biomaterial biocompatibility for prosthetic devices it is necessary to consider that an exercising prosthesis has to generally be under physical demands compatible with the bulk structure of the material or materials making up the prosthesis itself. At the same time the prosthetic surface interacts in a compatible or incompatible way with the surrounding tissues, in relation to the surface structure of its constituent materials.

Consequently, a prosthesis fully satisfies the characteristics of use for which it has been projected when the demands of volume and surface properties required by the implant are found to be compatible with the bulk and surface structure ofthe material, or materials, which compose the pros thesis.

This work has the established aim of specifying the relationships between these two aspects, at the same time showing how ~uch relationships forms the actual strategy of research and study of prosthetic biomaterials. In this respect two typical examples of structure­biocompatibility relationships are reported both in the field of bioceramics for orthopaedic use and organic polymers (hydrogels) for ophthalmological use.

Experimental

Sampies: calcium metaphosphate, calcium oligophosphates and calcium-aluminium oligophosphates have been prepared as reported in previous works (1, 2); hydrophylic contact lenses made of polyvinylpyrrolidone (PVP), containing a small quantity of butylmethacrylate and crosslinked by bifunctional methacrylates, are commercial devices.

Instrumental: Raman spectra have been recorded using a Jasco R-500 spectrometer with the

190

488 nm of a Spectra Physics Argon ion laser source; infra.red spectra have been recorded using a Jasco Ff/IR 7000 equipped with an A TR unit.

Results and discussion

The actual strategies for studying biomaterials for prosthetic use takes into account mainly: I) research into "new" biomaterials; II) research into "composites" biomaterials with suitable bulk and surface properties. The first strategy is limited by different factors, among which the principle one is that the

material must have a bulk structure compatible with the physical demands of the prosthesis. Rarely the same biomaterial is characterized by a surface structure highly compatible with the biomaterial surface-tissue interactions.

To give an example, in Fig. 1 there is reported the ideal case of a metal made up of a cubic crystal in which the atoms of the bulk form 6 metallic bonds with an hypothetical octahedral coordination. The volume properties of the material are strictly related to the bulk structure of the atoms from which it is composed, and in the case in question with the formation of 6 ideal octahedral metallic bonds which every atom forms inside the bulk.

On the contrary, the surface properties - always in the ideal case - are determined by the "in saturation" of the octahedric bonds (indicated in Fig. 1 by dotted lines) which characterize the metal atoms located at different positions on the surfaces.

It can therefore be seen that according to the location of the surface atom (on a face, along an edge, at a corner, etc.) the in saturation changes and so as a consequence the surface properties change.

a b

c d

Figure 1 - Different "insaturation" (dotted lines) of an atom with an hypothetical octahedral coordination depending on its position in a cubic crystal: a) in the center,

b) on a face, c) along an edge, d) at a corner.

Further variations in property can be induced in "real" situations for the presence of defects, fissures, etc.

The fact that the surface atoms of a crystal aquire electronic properties different from those of the "bulk" in the same crystal, has been noted some time ago and described also in quantum mechanic terms which explain the teories of the surface electronic states relative to metals (Tamm (3), Sockley (4», to ionic crystals (5, 6), etc.

191

As the bonds of the bulk structure are intensified in order to render the biomaterial more compatible with the physical requirements, the surface "unsaturations" are also intensified and therefore the interactions between the biomaterial surface and the tissues. These interactions can give origin to phenomena of tissue intolerance, as often happens, hence the first strategy ("new" biomaterials), pursuing the aim to prepare biomaterials with a better bulk structure necessary to render them compatible with the physical demands which are required, creates inevitable intolerance as a consequence of the exaggerated prosthesis-tissue surface interactions.

In additon, still with regard to the first strategy, abiomaterial of new conception can not be said to have biocompatibility until it has undergone an adequate follow-up of biocompatibility evaluation both at a biological and a c1inicallevel.

The second strategy ("composite" biomaterials with adequate bulk and surface properties) overcome the failings of the first strategy, in that the body of the "composite", with an adequate bulk structure and properties, is "coated" with a surface biomaterial which is chosen for its adequate surface properties. Generally, the chemical nature of the two biomaterials is different and it is not easy to obtain a chemical coating (in other words a chemical interaction between the two materials) without first making large changes of their interface structure and therefore inducing transformations and disturbances also in the bulk, which are difficult to eliminate.

The coating is therefore obtained in most cases by means of physical interactions (as an example by means of the plasma spray technique for ceramic materials and by melting for inorganic glasses), consequently it is not very stable and not always agrees with the required surface properties. Moreover, the physical treatment needed for the coating can induce structural surface modifications that in turn can cause intolerance.

In the following two meaningful cases of the previous considerations will be discussed, id est: i. the bulk structure of bioactive oligophosphate glasses for orthopaedic use and their elastic

properties; ii. the stabilization of hydro gels for ophthalmological use in order to limit protein deposits.

i. Calcium and calcium-aluminium oligophosphates. Calcium metaphosphate and calcium oligophosphate glasses containing also small

quantities of aluminium constitute a class of bioactive glasses of particular interest and importance in the orthopaedic field (7).

One of their peculiarities is to show a good biocompatibility, due to their chemical composition and structure very similar to that of the bone tissue; for this reason they have osteoactivity.

As it will be shown later, these biomaterials have the characteristic of, depending on their chemical composition, a different bulk structure and therefore different and controllable physical properties.

In Fig. 2 the Raman spectra of "controlled structure" bioactive glasses with composition xCaO'P205 where 1~x~1.65 are reported. The spectra elearly show that as x changes from 1 (case of calcium metaphosphate) to 1.65 there is a gradual increase in intensity of aRaman band at about 1040 cm- 1 attributable to VsP032- stretching mode of terminal chain groups. At

the same time, a gradual broadening and shift to lower wavenumbers of the band attributable to

the lateral chain vsP02-stretching mode (at about 1175 cm-1 for x=l and at about 1140 cm-1

for x=1.65) is observed. Lower modifications have been observed for the band attributable to

vsPOP stretching mode along the polyphosphate chain.

The spectroscopic results confirm that for the investigated glas ses there is a gradual demolition of the indefinite polymeric structure of calcium metaphosphate CaO/P205= 1

(deducible from the formation of P032- terminal groups) as the x=CaO/P205 molar ratio rises

and the formation of oligophosphate glasses with controllable chain length. In fact there is a elose correlation between the average chain length of the oligophosphate glass (defined from n, average number of phosphorous atoms) and the molar ratio x represented by the relation (8)

192

[ 1]

This relation has been experimentally confinned in the spectroscopic study of the similar vitreous system xNa20'P20S for which it has been confinned that when x=1.62S the value of n (deductable from [I]) is 3.2 and the Raman spectrum of the corresponding glass agrees with the spectra of the sodium tri- (NaSP301O) and tetraphosphate (Na6P 4010) crystalline products.

1400 1000 600 200 cm-1

Figure 2 - Raman spectra of calcium metaphosphate (x=l) and calcium oligophosphate (l~x~1.6S) bioactive glasses.

Tberefore it is possible to plan the synthesis of bioactive glas ses with controlled structure and properties, depending on the requirements due to the physical aspects of the biocompatibility. Tbe chain length n ofthese glasses and therefore theirphysical properties

193

depend on the x=Ca01P20S molar ratio. As an example Fig. 3 shows the behaviour of the elastic modulus (slope of the straight line)

of glasses marle of calcium metaphosphate (x=l) and mOI'OOver containing small quantities of alumina. The addition of alumina has been marle to improve the physical properties of the material.

The case discussed above constitutes a typical example of application of the first strategy ("new" biomaterials) in the research of biomaterials for prosthetic use with controlled bulk properties.

üi 0.5 c: o ~

~ Q)

c:

o c(

o ...J

o y= 0.14

o y=0.10

• Y =0.00

O;----.---r---.---,r---.---,----r---.------------~-o 0.5

PROBE DISPLACEMENT( ~~ %)

Figure 3 - Dynamic load thermomechanical analysis of calcium metaphosphate (y=ü) and two calcium aluminophosphate glasses CaO. y AI2ÜJ' P20S (the slope of the straight

line is the elastic modulus).

ii. Hydrogels based on PVP. Hydrogels are used as soft contact lenses with a long wear since their elastic propenies are

very similar to those of the cornea (9). A peculiarity of these biomaterials is their hydrophilicity that, in the case of the hydrogel

based on PVP, reachs the 70% w/w. The hydrogel surface is coated with a cushion of water where the molecules of the ftrst

layer interact with the surface hydrophilie centers of the polymer and the interaction is transmitted and dispersed by the hydrogen bonds of the water at the surface.

Measurements made by means of vibrational Raman spectroscopy performed in our laboratory on an hydrogellenticule enriched with D20 have shown, after implant in rabbit eye, that a turnover of water between the lens and ocular tissues takes place. In fact the Raman spectrum of the lens after implantation shows only the bands ofH20 (10).

At the same time other measurements (11) have shown the formation of protein deposits on

194

the lens smfaces after a follow-up with continuous application and also that the quantity of deposits correlates with the loss of water of the lens because of the turnover. The quantity of water has been evaluated by means of thermogravimetric analysis executed on the lens immediately after removal.

Therefore the lens-tissue turnover of water is the phenomenon that mainly limits the structure and surface reactivity and at the same time substantially affects the chemical biocompatibility of the prosthesis.

A modification of the lens hydrophilicity in order to make the cushion of surface water more stable is not practical since with hydrophilicity the physical properties of the lens will also change. In particular the elastic modulus - as we have shown (9) - will became different from that of the cornea.

Therefore interventions are meant to give stability to the smface structure of the hydrogel by means of the formation of "composites" (as for example an adequate surface coating) without any modification of its bulk: structure.

Thus hydro gels constitute a typical example of the application of the second strategy ("composite" biomaterials) to the research of biomaterials with suitable bulk: and surface properties.

References

1. Bertoluzza, A., Battaglia, M.A., Simoni, R. and Long, D.A.,Vibrational spectroscopic studies of a new class of oligophosphate glas ses of biological interest. Part 2, J. Raman Spectr., 1983, 14, 178-183.

2. Bertoluzza, A., Fagnano, C., Marinangeli, A., Simoni, R., Tinti, A. and Morelli, M.A., Vibrational (Raman-laser and infrared) study on Ca<r0:3)2-Al2~ glasses, Advances in Biomaterials, vol. 7 (Biomaterials and Clinical Applications), Elsevier, Amsterdam, 1987, pp. 511-516.

3. Tamm, I.E., A possible kind of electron binding on crystal surfaces, Phys.Z.Soviet., 1932, 1,733-746.

4. Shockley, W., Surface states associated with a periodic potential, Phys.Rev., 1939,56, 317-323.

5. Levine, J.D. and Mark, P., Theory and observation of intrinsic surface states in ionic crystals, Phys.Rev., 1966,144,751-763.

6. Mark, P., Chemisorption states ofionic lattices,J. Phys.Chem.Solids, 1968,29,689-697.

7. Pernot, F., Rogier, R., Zarzycki, J., Bonnel, F. and Baldet, P., Les biomateriaux vitreux et vitroceramiques en chirurgie orthopedique et osseuse, BuH. Soc.Chim. Fr., 1985,4, 519-522.

8. Bertoluzza, A., Morelli Bertoluzza, M.A., Fagnano, C., Battaglia, M.A., Indagine vibrazionale Raman e ultrarossa di vetri oligofosfati sodici di composizione xNa20'P205' con x compreso fra 1 e 1,625, Rend. Accad. Naz. Lincei, 1976, LXI, 269-273.

9. Bertoluzza, A., Monti, P., Simoni, R., Garcia-Ramos, J.V., Caramazza, R., Cellini, M., De Martino, L., Calzavara, A., Applications of Raman spectroscopy to the ophthalmological field: Raman spectra of soft contact lenses made of polyvinylpyrrolidone (PVP), Studies in Physical and Theoretical Chemistry, vol. 45, Elsevier, Amsterdam, 1987, pp. 595-604.

10. Results not yet published. 11. Bertoluzza, A., Monti, P., Simoni, R., Chemico-physical structural characteristics of PT

70 contact lenses before and after continuous wear, Advances in Biomaterials, vol. 7 (Biomaterials and Clinical Applications), Elsevier, Amsterdam, 1987, pp. 315-320.

195

EXPERDIElITAL STUnY ON THE PROPERTIES OF BYDROXYAPATITE COATED IKPI.ART5.

* * ** ** C. Gabbi , P. Borghetti , N. Antolotti ,5. Pitteri

*Facolt8 di medicina veterinaria delI' universit8 di Parma. Istituto di anatomia normale degli animali domestici.

** Flametal S.p.A, Biocoatings Division.

ABSTRACT.

Various experiments confirm the biocompatibility of plasma sprayed hydroxyapatite used for implant coating, its osteoconductivity properties and its capability of accelerating the reconstruction of bony tissue, including neo formations as in the medullar cavity, if in presence of osteogenetically differentiated cells. The optical microscopy observations indicate a good bone-hydroxyapatite­metal adhesiOß, and this is confirmed by pull out tests on cylindrical specimens which were implanted in the femurs or tibias of a dog for a duration of three months.

INTRODUCTION.

The biocompatibility and the osteoconductive properties of hydroxyapatite (Hap, chemical formula: Cal0(P04)6(OH)2) are two weIl established facts. with various applications in bone surgery ([1] to [7]). Previous works of biological characterization carried out in presence of bony cells after permanence in vitro cultures (8) and in vivo ([I], [5], [6], [8]) showed that hydroxyapatite coatings, plasma sprayed on titanium implants ([10], [lI]), favour and accelerate the growth and reconstruction of bony tissues, with excellent adhesion between bone and hydroxyapatite ([1], [5] to [11]). The following research deals with the biological response of a highly crystalline hydroxyapatite coating (Calcid, TM) when implanted in rabbits for per iods of one, two and three months, and in a dog for aperiod of three months; finally, a mechanical evaluation of these latter implants is carried out by pull out testing.

196

MATElUALS MD HBTIIODS.

The various implants introduced were based upon (6 mm x 10 mm x 1.5 thick) Ti-6AI-4V thin plates in the case of the rabbits, and upon (3.8 mm diameter x 7 mm long) Ti-6AI-4V cylinders in the case of the dog; some of these specimens were implanted as such for comparison while, by means of a modified plasma torch ([8] and [11]), the remaining were implanted, coated by a 200 pm thick layer of highly crystalline hydroxyapatite. The chemical analysis of the hydroxyapatite coatings showed trace elements weIl below the limits of the ASTH specification F 1185-88 (table 1). By X ray diffraction of the crystalline phase, its crystalline ratio was estimated at 72 % ([8], [12]), with a Ca/P ratio equal to 1.67 ± 0.01, and without traces of parasite peaks; the expected hydroxyapatite spectrum, free from impurities such as calcium oxide and tri or tetra calcium phosphates was also obtained by infrared analysis, and the purity of the coating was estimated at over 97 % .

TABLE 1.

Chemical analysis of the Calcid™ highly crystalline hydroxyapatite.

ASTM F 1185-88 highly cristal. Hap

As < 3 ppm < 2 ppm Cd < 5 ppm < 1 ppm Hg < 5 ppm < 1 ppm Pb < 30 ppm < 1 ppm

total heavy < 50 ppm < 30 ppm metals

CU < 1 ppm Mn <100 ppm Fe <100 ppm

calculated Ca/P ratio 1.67 ± 0.01 estimated crystallie ratio 72 % estimated Rap yurity > 97 %

Implants were introduced in the tibia metaphysis of groups of ten rabbits, where they remained for one two and three months, while other implants had a three months permanence in the dog hind legs in the locations shown by the radiograph of figure 1.

197

FIGURE 1. X radiograph on the implants locations in the dog.

One, two and three months respectively after the implant operation substances marking the deposition of bony tissue were injected, followed by the sacrifice of the animalsj osseous segments were treated according to the usual histological techniques for the optical observation of hard, non decalcified tissues. The specimens implanted in the dog were prepared for vertical pull out tests from bone sections at a constant speed of 1 mm/min, on a computerized Instron tensile machine.

RESULTS AN» OBSERVATIONS.

Optical microscopy on preparations obtained from the test specimens implanted in the rabbits for both one (figure 2) and three months (figures 3 to 5) evidences a "restitutio ad integrum" of the damaged tissue caused by the implantation and an integration between the bone and the hydroxyapatite coating without interposition of fibrous connective tissue, which would suggest a good bone-implant, strengthj while in spite of osseous reconstructions sometimes observed around uncoated titanium test specimens, these always display a thin layer of connective tissue at the interface bone-implant, which is detrimental for obtaining a good attachment. Neo depositions of bony tissues beside the implants in the region of the medullar cavity are on the contrary only observed after one month of implantation, which could be explained by the differences in the trabecular structure of each subject, or by the ceasement of the superficial reactivity initially stimulated by the amorphous phase.

198

FIGURE 2. Histological section with hydroxyapatite coating , on a rabbit after one month in vivo. B: hydroxyapatite. C: bone tissues. Polarized light. xlOO magnification.

FIGURE 3. Histological section without hydroxyapatite coating on a rabbit after three months in vivo. Polarized light. xl60 magnification .

lW

FIGURE 4. Histological section with hydroxyapatite coating. on a rabbit after three months in vivo. B: hydroxyapatite. C: bone tissues . Polarized light. x160 magnification.

FIGURE 5. Histological section with hydroxyapatite coating, on a rabbit after three months in vivo. A: titanium. B: hydroxyapatite. C: bone tissues. Polarized light. x40 magnification.

200

The tensile tests after three months of implantation presented in table 2 confirmed the histologicalmicrographs of figures 3 to 5 for what concerns the affinity between bony tissue and hydroxyapatite by comparison with uncoated specimens, and this is also evident on figures 6 and 7: typical tensile curves are shown in the figure 8, the highest bond strengths being obtained with the hydroxyapatite coatings which, as already indicated by the histological observations of figures 4 and 5, appear continuous, uniform and without variation from the initial thickness, suggesting an in vivo stability and preservation of cohesion and possibly adhesion to the titanium substrate.

6

FIGURE 6.

FIGURE 7.

TABLE 2

Pull out strength testing after three months in vivo.

locations femurs tibias Type of coating

uncoated bond strength (MPa) 4.8 3.0

highly crystal. Hap bond strength (MPa) 17.9 14.8

(the bond strengths presented are averages of two values)

7

Uncoated specimen pulled out from the dog femur after three months of implantation. Hydroxyapatite coated specimen pulled out from the dog femur after three months of implantation.

201

kN

2

2 3 4 nun

FIGURE 8. Typical pull out curves of an uncoated implant (U) and of an hydroxyapatite (e) coated implant, after three months in vivo.

Such behaviour would suggest that highly crystalline hydroxyapatite coatings can res ist dissolution and infiltrations by bone cells and body fluids during long term implantations.

CONCLUSIONS.

It must be emphasized that in order to correctly interpret the various results presented, and to repeat them, one must bear in mind the situation of mechanical stability of the implant, which is not directly submitted to external forces, but only in contact with bony tissues influenced in their growth and remodelling by the natural loads. The experiments and observations carried out in this research confirm the biocompatibility and the osteoconductivity of hydroxyapatite, plasma sprayed on implants for in vivo experimentation, with osteointegration between the hydroxyapatite coating and the bone j "restitutio ad integrum " of the surgical damage caused by the introduction of the hydroxyapatite coated implant are observed after both one and three month in vivo. Pull out testing on specimens after a three months in vivo implantation shows that the cohesion of the coating or the adhesion to the substrate, thus eventually the bond between bone and implant, maintain high levels when the specimens are coated by highly crystalline hydroxyapatite. These bond strength results should be statistically analyzed through the establishment of a standardized methodology considering species and age of the animals, and position, number and permanence of the implantsj however, they appear quite significant and their confirmation is expected. In this respect, the behaviour and evolution during long term histological experiments (over one year) of amorphous material formations observed at high magnification in the interfacial areas should be investigated. Such structures are compatible with the organization of the bony tissues, and should give valuable informations on the coatings stability and their relation with the hydroxyapatite crystallinity.

2m

REFERENCES.

1. C. Gabbi, L. Russo, L. Bo. Impianto transcorticale rivestito da idrossiapatite. Annali Fac. Med. Vet. Univ. Parma, VI, 47-54, 1986.

2. H. Alexander, J. R. Parsons, J. L. Ricci, P. K. Bajpai, A.B. Weiss. Calcium based ceramics and composites in bone reconstruction. CRC Critical reviews in biocompatibility, 4 , 1, 1987.

3. M. Ogiso, H. Kaneda, J. Arasaki, K. Ishida, T. Tabata. Epithelial attachment to hydroxylapatite ceramic implants. Journal of dental research 59, page 941, 1980.

4. K. De Groot. Bioceramics of calcium phosphate. CRC Press, Inc., Boca Raton, Florida, 1986.

5. K. De Groot.

6.

Hydroxylapatite as coating for implants. Interceram, 4, 38-41, 1987.

J. F. Osborn. The biological behavior of the hydroxyapatite ceramic titanium stem of a hip prothesis - the first histological human autopsy specimens. Biomedizinische technik 32 7-8, pages 177-183, 1987.

coating on a evaluation of

7. S. D. Cook, K. A. Thomas, J. F. Kay, M. Jarcho. Hydroxyapatite-coated titanium for orthopedic implant applications. Clinical orthopaedics and related research 32, pages 225-243, 1988.

8. C. Gabbi, P. Borghetti, A. Tiziani, G. Liborio, N. Antolotti, G. Cattaneo, S. Pitteri. Hydroxyapatite coating: mechanical tests and biological behaviour. Seventh Cimtec, June 24-30 Montecatini.

9. P. Ducheyne, W. Van Raemdonck, J. C. Heughebaert, M. Heughebaert. Structural analysis of hydroxyl-apatite coatings on titanium. Biomaterials 7, pages 97-104, 1986.

10. K. De Groot, R. G. T. Geesink, C. P. A. T. Klein, P. Serekian. Plasma sprayed coatings of hydroxylapatite. Journal of biomedical materials research 21, pages 1375-1381, 1987.

11. N. Antolotti, M. Casali, A. Tiziani, L. Giordano. La tecnica plasma spray nei riporti in aria e in vuoto. Giornata di studio, Centro di studio AlM, Milano, 12/10/88.

12. G. W. Hastings, D. Dailly, S. Morrey. Hydroxyapatite coatings. Bioceramics - Ishiyaku Euro America Inc., pages 355-358, 1989

W3

PHYSICO-CHEMICAL CHARACTERIZATION OP HYDROXYAPATITE OPUNKNOWN

MANUPACTURE

A. Ravaglioli, A. Krajewski, A. Piancastelli (IRTEC-CNR, National Council of Research, Paenza, Italy)

R.Martinetti (Finceramica s.r.l., Paenza, Italy)

INTRODUCTION

The literature speaks of Ca-phosphates as implant materials more or less readsorbable, highly compatible, and able to stimulate bone ingrowth close to this limit surface.

The knowledge, however, is insufficient, and there are differences in assessing the partial readsorbability of Ca­-phosphates on the basis of specific compounds.

We do not know which type of powder the different authors (researchers and/or surgeons) have used, or if the powders were made in their laboratories or came from trade (also because the literature is not always clear in this respect).

In our opinion the biochemical behaviour should not change too much if the compound is sintered as a dense or porous ceramic, while some authors underline that the exist­ence of porosity determines a higher readsorbability of the hydroxyapatite (HA). (Porosity should increase ion exchange).

Some doubts have arisen that the HA at our disposal might be a mixture of other compounds, more adsorbable than hydroxyapatite.

On the basis of what has been said, a statistical research was carried out on products utilized in the litera­ture, including those coming from trade.

MATERIALS !ND EXPERI~ENTALS

Various x-ray analyses, chemical analyses, and tests were

204

carried out in order, in parti cular , to verify what CalF ratios could be obtained on the basis of the data coming from roentgenographic tests.

Tab. 1 Chemical analysis and CatP ratio of trade hydroxyapatile (HA)

PRODUCTS A B C D E F· G· H· I· L· (bulk) 1500'C

IIA(%) 99 97 98 99 100 95 80 95 95 99 CaZPZ07 (%) . 2 2 - - - 20 (a) (a) -Ca/P (moles) 1.652 1.666 1.588 1.690 1.690 1.689 1.418 . 1.658 1.689

• = powder (a) = amorphous

M·

70

30

1.504

So it was ascertained (see TABLE 1) that many of the powders utilized are made up not only of hydroxyapatite as a chemical component, but of other Ca-phosphates as weIl.

Since the correlation between readsorbability of calcium phosphate and its chemical behaviour is directly linked to the solubility product, it follows that in a physiological-pH environment HA is the more insoluble compound.

The existence of other phosphate compounds may cause these to be solubilized more or less rapidly. Therefore the overall bulk of the ceramic body will decrease before re ach­ing the equilibrium determined by the primer of the calcifi­cation mechanism.

Furthermore, we have ascertained that the phosphate that accompanies the HA is generally calcium pyrophosphate, pre­sumably formed on preparation in order to become suitable for medical use, surely much more soluble than calcium tri­phosphate.

On the basis oi the overall Ca/F ratios it was ascertain­ed that in the different compositions there exist some per­centages of other phosphates, because the ratio is often less 5/3 = 1.6 in moles, corresponding to the one of HA.

Since with calcium triphosphate such ratio is 3/2 = 1.5 and with pyrophosphates it is 2/2 = 1, it is evident that in the case of a mixture the overall ratio will be.less than the HA value of 5/3.

An interesting case is that of "11" powders, announced as calcium tri phosphates, in which the Ca/P ratio is actually the ratio of calcium triphosphate. In reality they are con­stituted of a binary mixture of HA and pyrophosphate in an appropriate proportion to obtain such ratio (1.5).

As chemical reagents for organic and inorganic synthesis, "lIn powders are very suitable as simulators of triphosphates but not as suitable for application as bioactive ceramic.

205

CONSIDERATIONS AND CONCLUSIONS

On the basis of what has been said, it will be necessary to standardize the criteria to be adopted in experimentals.

Nithin our Italian group a debate has been opened as to the way of carrying out experiments on ceramics of guaranteed pure composition in order to have at disposal standard beha­viours as terms of reference.

Therefore the powders and the products obtained will have to be carefully analyzed be fore each implant, with the aim of avoiding any pollution from compounds different from those announced.

According to our plans, the second step will be repre­sented by experiences with carefully prepared mixtures of pure compounds for obtaining ceramics of well-fixed, mixed composition.

The aim is to verify if there is a linear correspondence between the composition of the mixture and the biological behaviour (readsorbability, bioactivity, etc).

The objective of these experiences is to verify the possibility of obtaining ceramics of controlled readsorbabi­lity such that it may be calculated.

206

APPLICATION OF CERAMIC COMPOSITES AS IMPLANTS:

RESULT AND PROBLEM

S.M. BARINOV, Yu.V. BASCHENKO High-Tech Ceramics Scientific Research Cent re Academy of

Sciences of the USSR, Ozernaya 48, 119361, Moscow

ABSTRACT

Bioactive calcium phosphate apatite ceramic has excellent biocompability for failed bone tissue replacements. However, its low strength and fracture toughness restrict application of porous calcium hydroxyapatite ceramics implants. Another important problem is the bioactivi ty and compatibili ty regulations. Solution of problem is composite structures formation, either with strong ceramic filler, or with active tissue components. Review is presented concerning activities in the field of developments the ceramic composites based on calcium hydroxyapatite. Formation on composite structures such as hydroxyapatite - zirconia improve the mechanical behaviour. The characteristics of composites with collagen have been investigated. Preliminary tests results and existing problems are discussed.

INTRODUCTION

Calcium-phosphate ceramics and glass-ceramic materials are biocompatible with the living organism tissues. However, their direct application in replacement and reconstructive surgery as implants is rather difficult. There are, at least, two problems to be urgently solved. The first of these problems is the low strength, fracture toughness and fatigue resistance inherent in calcium phosphates. This prevents the use of calcium-phosphate ceramics as load carrying bearing implants.

207

Moreover, for new bane tissue to be formed, the implant needs to be porous with 50-100 Nm diameter open pores. This porosity significantly decreases low mechanical properties of these ceramics. Another problem is related to the necessity for the control of implant's bioactivity. The latter means not only the ability to form strong bonding with bone tissue but also the osteogenesis stimulation and the property of biodegradation. It has not been decided whether the calcium-phosphate ceramic i tself is the material providing the formation of new bone tissue or any special additional conditions are needed for osteogenesis to be proceeded. In connection with the above-stated the present paper seeks to do a brief review of the main strends of the investigations in the field of bioactive composite materials carried out in the U55R.

SOME EXPERIMENTAL RESULTS

Composite materials of hydroxyapatite - zirconia and tricalcium phosphate - zirconia systems have been obtained at the D. I. Mendeleyev Moscow Institute of Chemical Technology. The purpose of the investigation was to obtain porous ceramics with specified pore size (50-300 Nm) and increased strength. Zirconium dioxide introduction into hydroxyapati te made i t possible to obtain ceramics with the bending strength of 50 MPa at the porosity of 45%. 5intering temperature of composite mixtures did not exceed 1200°C, i.e. it was less than hydroxyapatite decomposing temperature. Two three fold increase in strength has been achived due to introduction of zirconia in the amount of 50% both into hydroxyapatite and into tricalcium phosphate. Degree of porosity within the limits 5-50% and pore size 50-300 Nm are obtained owing to the application of special combustible additions. Preliminary biological studies testified to the biocompatibili ty of these composites with bone tissue and the formation of strong bonding between the implant and the bone. The use of composite implants based on glass-ceramic materials is expected to be promising. Osteocompatible glass-ceramic materials in the 5i02 - P20S - Al203 - MgO (CaF2 ) system that contain tricalcium phosphate and calcium apatite as the main crystalline phases have been worked out at the Technological

208

Institute in Leningrad. Experiments on the implantation of the

materials concerned into the front left bone distal epiphysis

in rabbit have been carried out. Osteogenesis indications were

observed two weeks later. On the basis of glass-ceramic materials

the compositions for the filling of bone tissue defects were

developed. In doing so, the purpose was to provide close contact

with the cortical bone, which was achived by the combination of

specified rheologie mass properties and the composition ability

to harden within 5-8 minutes of its supply into defect zone. The

composi tion concerned contains the powder of glass-ceramic

calcium-phosphate material as an inorganic constituent and

collagen and silicone rubber as weIl. Drugs can be introdueed

into the eomposition. On the basis of the above materials porous

implants can be produeed. Aecording to the preliminary

investigations the material were shown to display

bioeompatibility and the absence of eytotoxicity.

Aseries of eomposite materials with polymerie matrix eontaining

the powders of hydroxyapatite, bioglass or glass-ceramics in the

Si02 - A120] - Zr02 - CaO - P205 system as a filler have been developed in the Institute of Medieine Polymers. Polymer filling

degree amounts to 40-60% by volume. The materials are made in

the form of paste and can be used for tailor-made products of

complicated shape during the operation since they can be polymerized to the solid state at the temperature as high as 100·C in a usual desieeator. Polymerization by the light is

eonsidered to be possible too. Matrix looks like the mixture of

methaerylic oligomers having two and more double bonds in a moleeule. Polymerization oeeurs in the euring by the action of

peroxide initiators. The similar binders are widely used in the

compositions of stomatological materials. The merit of the above

materials is considered to be low volatili ty and water solubili ty

and, as a consequence, low toxicity. Composite materials have

the following properties: porosity of 0,14 - 1%, water absorbtion

of 8-860 mg/mm2, water solubili ty of 53-560 mg/mm2, tensile

strength of 20 - 44 MPa. The properties of the material can be

controlled by changing the nature and the treatment conditions

of the filler.

The properties of hydroxyapatite being dependent, largely, on

its thermal background, the investigations related to the

development of the methods of production of so-called "cold"

209

ceramics have been carried out at Moscow State University. The ceramics structu~e is formed at the synthesis stage with the use of protein collagen binder. The materials concerned pass the tests for hone defects healing to satisfaction. Aseries of materials based on collagen wi th fillers of tricalcium phosphate or hydroxyapatite have been worked out at the Academy of Medical Sciences in Moscow. A ratio between protein and inorganic constituents of materials was (30-35):(70-65). Composites with hydroxyapatite have the property of high water absorbtion and can be succesfully applied in stomatology as drug carriers. Composites with tri calcium phosphate are intended for the filling of bone cavities after different interferences, including combinations wi th bone transplants. When used for osteomeli tic foci. It can be saturated wi th antibiotic and antiseptic solutions. Biological properties of these materials are as foliows: a) the drug is subjected to bioresorbtion with the replacement by material bone, b) it has osteoinductive effect, c) it causes no inflammatory response. Drug sterilization can be performed by means of gamma rays. The drug has passed clinical test to satisfaction in therapeutic stomatology. As for the collagen contribution to the process of biological interaction with the bone tissue, there has been no unanimous opinion on the matter yet. However, the data are available that disprove the above idea. Experiments on ectopic stimulation of osteogenesis have been conducted at the Central Institute of Traumatology and Orthopedics by Dr. G.N. Berchenko. For this purpose, complex graft consisting of xenograft bone and autogenous marrow were implanted into the supraspinations muscle of rats. Several holes, 1 mm in diameter, were drilled through the xenograft bone prior to its treatment. After that, a "pure" bone was obtained by means of the xenograft bone treatment. Autogenous marrow was obtained of rat' s thig-bone. The quality of the bone marrow constituted up to 30 mg. Complex graft was subjected to histological verification during 30-245 days. Morphological verification testified to the formation of thin capsule of connective tissue around the xenobone without any signs of rejection. Formed inside the canals was the bone tissue containing bone marrow elements. Some time later, the implantation into vessel xenobone occured with subsequent

210

osteogenesis and bane marrow formation. New bane did not formed

in the animals of control group to which xenobone was implanted

without bone marrow.

CONCLUSION

Thus, the above results demonstrate, in the first place, considerable variety of possible component combinations in composite implants and, in the second place, insufficiency of our knowledge on biological nature of osteogenesis when artificial implants are used.

211

RADIOACTIVITY MEASURBMENTS O~ ZIRCORIA POWDBRS

* G. CAPANNESI, A.F. SEDDA, C. PICONI, F. GRECO ENEA, Casaccia Research Center, Rome, Italy

*catholic University - Bioengineering Center, Rome

ABSTRACT

Partially Stabilized Zirconia (PSZ) ceramics are proposed as an alternative material to the A1203 components now used in prosthetic devices, for their bett er toughness and wear resistance compared to the Alumina ones. Due to the fact that to Zirconium may be associated a low content of actinides, it is worthwhile to assess the total radioactivity of PSZ materials to be used in biomedical applications. Measurements were performed on four PSZ sampies , two coming from the market and two manufactured by Gel Supported precipitation (GSP) process by TEMAV SpA. Natural radioactivity of bone was measured and results are reported for comparison. Measurements were performed by HPGe detectors, coincidence connected wi th a NaI (30 • 5 x 30. 5) cm Compton suppressor, thus enhancing sensitivity to 5xlO-4Bq.

INTRODUCTIOR

Hearth crust is relatively rich in Zirconium, as it ranks

as eleventh in abundance scale, that is more than common

metals as Copper, Lead, Nickel, Zinc. Zirconium minerals

of industrial interest are the zircon (ZrSi02) found in

Madagascar, Brazil and in Florida, Australia and South

Africa beach sands, and the Badelleyte, (Zr02) extracted

from the Mountain Plateau in Brazil.

It is weIl known that some Zirconium ores may contain

radioactive isotopes, and this fact originated, in a

society more and more concerned with environmental

212

quest ions , attention about possible harmful radioactivity

levels of Zirconia items, like for instance, X-Ray

contrast media for bone cements [1].

Radioactivity of Zirconium is mainly due to the presence

of isotopes coming as decay products from natural Uranium

and Thorium associated to Zirconium bearing ores.

Zirconia ceramic precursor today present on the market

are mainly made for refractory use, as furnace liners,

for the manufacture of high-temperature structural

components, for protective coatings on alloys. All of

these applications do not need much attention to the

small radioactivity levels linked to the material, but it

is worthwhile a careful assessment of this quantity when

the material has to be used within the human bOdy, as for

the manufacture of prosthetic devices.

MATERIALS AND METHOD

Katerials

The gamma radioactivity of four PSZ powder samples, two

coming from the market (samples A and B) , two

manufactured by TEMAV SpA by GSP process (samples C and

0) as well as the gamma radioactivity of human bone were

measured (See Table 1). Sample geometry was a cylinder 50

mm in diameter, 15 mm high.

TABLE 1 PSZ samples characteristic

Sample A B C 0

Composition Y-PSZ Y-PSZ Y-PSZ Ca-PSZ

Manufacturer Tosoh Dyn. Nobel TEMAV TEMAV

213

Human bone specimens from surgeries were calcinated at

900 C for 8h and reduced to powder in a planetary ball

mill using agate jars.

Kethod

Gamma countings were performed by lead shielded Ge (Hp)

detector (ORTEC), with FWHM 1.68 KeV and efficiency

of 29% at 1332 KeV, coincidence connected with a NaI

(30.5 x 30.5) cm Compton suppressor enhacing sensi ti vi ty

to 5x10-4 Bq, coupled with a ORTEC PHA/MCA system. After

correction of spectra for low energy gamma rays self

absorption, background subtraction and multiplet

deconvolution, total activities were evaluated by

direct comparison with multipeak aqueous solution with

the same geometry of the sampies, to calibrate the energy

and efficiency seal es of the measuring system to absolute

units. Calculations were performed by ORTEC Maestro II

standard software package.

RESULTS

The measured total specific gamma activities are

summarized in Table 2, the main identified radionuclides

contents is reported in Tab!e 3.

A

70±12

TABLE 2 Experimental results (BqjKg)

B C D

7240±210 1220±140 1160±140

Bone

85±15

214

TABLE 3 Main radionuclides contents in the sampies (BqJKg)

A B c D Bone

210 Pb 428 69 83 212 Pb 505 75 35 214 Pb 7 764 295 115 5 212 Bi 556 114 95 214 Bi 7 674 258 115 5 219 Rn 107 36 57 223 Ra 96 29 49 224 Ra 482 69 83 226 Ra 1221 164 227 Th 102 31 56 228 Th 245 151 223 230 Th 1327 234 Th 319 235 U 1 27 208 Tl 154 23 8 228 At 429 96 69 13

40 K 52 43 104 60

DZSCUSSZOH

Measurements reported deals on1y with gamma activities;

alpha emitters are below the detection limit of our

system (-5 count/h), whi1e Beta activity was neg1igib1e

(-4 Bq) due to the limited range in Zr02. Total gamma

activity from Table 2, must be rescaled to the actual

mass of a 28mm dia. ceramic THP sphere, Le. 50 g, and

put in comparison with the total mass of the mineral part

of human skeleton (4 Kg) and of whole body

radioactivity,as reported in Table 4.

TABLE 4 Radioactivity data comparison (Bq)

PSZ sphere

3.5 - 360

Whole human body

3700

Skeleton

340

Resulting whole body doses from the above sources are

reported in Table 5 in comparison with other sourees.

215

TABLE 5 Whole body doses ( ~Sv/anno)

PSZ inner body full natural medical sphere sources skeleton background doses

2-1070 125 12 1800 800

Natural source of external irradiation are cosmic rays.

(In Italy at sea level the whole body dose is 0.5

mSv/year) •

A second source of external irradiation comes from the

natural abundance of radionuclides in the rocks and in

the ground and buildings. In Italy the whole body dose

is 0.5 mSv/year in the Aosta district in the Alps, and 2

mSv/year in the viterbo district, near'Roma.

Diagnostic radiology dose (X-rays) on Italy's population

is 0.8 mSv/year on bone marrow and 0.5 mSv/year on

gonads. As majority of bone marrow is deeply inside the

human body, this dose has to be accounted as whole body

irradiation.

In conclusion, the average external irradiation in Italy

accounts about 1.8 msv/year (180 mrem/year), 0.9 msv/year

due to natural background (cosmic rays, ground

radiations) , 0.8 mSv/year due to medical diagnostic and

the balance due to technological sources of radiations.

CONCLUSIONS

The examined zirconia powders show a wide range of total

specific radioactivity va lues , and a qualitative

difference in gamma spectrum of contained radioactive

isotopes. As shown in table 3, in some cases (sampie A)

only 40-K, 214-Pb and 214-Bi are present, while in other

sampies (sampie B) all decay products from 235-U, 238-U

and 232-Th are present in not negligible amounts.

216

From these data the followinq conclusions can be drawn:

i) zirconia can contain a wide variety of qamma emittinq

isotopes The associated radioactivity level hiqhly

depends from the source of minerals used in preparation,

and from the purification level of the chemical precursors of the oxide;

ii) to lower the radioactivity levels of PSZ up to

neqliqible levels is mandatory for materials selected for

prosthetic applications. This implies the need of

performinq accurate radioactivity measurements for the

selection of chemical precursors to be used in the

preparation of ceramic powders;

iii) the quality and the intensity of qamma spectra are

an useful tool for the determination of the chemical

puri ty level of materials. Our work is in proqress to

characterize potentially toxic elements in Zirconia and to evaluate their correlations with the identified

radionuclides.

REFERENCES

1. Hopf Th.,Scherr O.,Globel B.,Hopf eh., Verqlechende tierexperimentelle untersuchunq zur Gewebsvertraqlichkeit und Messumqen der Radioaktivitat verschiedenern Rontqenkontrastmittel. Z.Orthop.,1989,127,620-624.

217

MICROSTRucruRAL ANALYSIS OF BIOCERAMIC MATERIALS

BRIAN BOUSFIELD Technical Marketing Manager (BUEHLER EUROPE)

Chairman of Enropean Sodety for Microstructnral Traceability

Univenity of Warwick Science Park, Coventry CV4 7EZ U.K.

ABSTRACT

The compatibility of various materials for implants and other biomedical applications is often the subject of microstructuraI analysis. The 'fingerprint' of metallurgy is said to be the material microstructure. In biomedical studies tbe microstructure is important at the pre-implant stage and for cIinical trial analysis. Today's technology allows materials for surgical implants to be investigated as thin sections for radiography and transmitted light microscopy or by the top surface preparation technique for reflected light analysis. The object of this paper is to introduce this technology with particular reference to the top surface preparation for microstructural analysis of Hydroxyapatite and different species of coral. Unlike many traditional materials, these materials respond adversely to an incorrect surface preparation technique. The need therefore to define a logical and systematic approach to surface preparation, to ensure it is repeatable, and is also a faithful representation that can if necessary be audited to some form of standard will not be overlooked.

INTRODUCTION

Materials selected for implantation into the human body must satisfy specific requirements in what could be described as a hostile environment. Achieving the appropriate tensile strength, frictional resistance, corrosion resistance and ability to withstand shock loads, are all factors that can be calculated in advance. It is when we wish to encourage bone growth into the implant that more involved methods of investigation become necessary. The compatibility of various materials, the size of pores, the degree to which they are discrete, the effect of spray coating to metal substrates, the coating interface bond, the porous and non-porous surface depth, the bone to implant surface structure are all factors to be investigated and controlled prior to implant. It is also necessary to retrieve implants to investigate the interface between implant and host. Has there been a rejection, have the biodegradable characteristics of the implant rate been compatible with bone ingress? All this information is available from microstructural analysis by analysing tbe microstructure achieved through cross­sectioning, and careful surface preparation.

218

MATERIALS FOR INVESTIGATION

Dense or porous coatings can be applied to implant materials via the powder sintering or the powder plasma spray technique. Metal sintering has been used on screws and pins, plasma spraying being a favoured route for larger implants requiring a coating of hydroxyapatite (HA). The two examples chosen for investigation are plasma sprayed HA on a titanium 6Al- 4V substrate and two different species of coral.

The examples quoted are pre-implant conditions. There is obviously also a need to investigate the interfacial bond after varying periods of use. The rate of bonding and the strength and stability of the bond vary with the composition and microstructure of the bioactive material. The surface preparation techniques given will also apply at the clinical trial stage. Techniques using cross-sections in reßected light are recommended when investigating materials exhibiting brittle fracture mechanisms (bone, coral, hydroxyapatite etc).

SURFACE PREPARATION TECHNlQUE

Once we use the microstructure as the fmgerprint of material characteristics, we must ensure the finished surface preparation has faithfully revealed the true microstructure. Preparing surfaces by mechanical material removal techniques is by definition a stress related activity which creates induced structural - damage at every preparation stage. These induced artefacts will not necessarily be removed when the specimen is polished and analysed. It is for this reason that a systematic approach to sampie preparation (1) was published giving a guide to the correct selection of preparation surfaces and abrasives. It soon became apparent, in particular with plasma coatings, how easy it was to prepare the specimen surface in such a manner as to achieve the results expected, rather than necessarily those which represent the true unaltered microstructures.

(a) porosity can be either induced or 'lidded' in such a way as to mask the true porosity level.

(b) delamination can be caused by the preparation procedure.

(c) radial cracking could be a coating defect or a specimen mounting induced artefact.

Clearly, the preparation procedure on its own is incomplete. Tbere is a need to audit to statistical and visual standards; these standards to be traceable to a recognised authoritv. This subject was addressed in 'Progress towards a Metallographie Standard' 1990. (2)

Material Investigation

Example No. 1 : Hydroxyapatite

This material is non-toxic yet biologically active to create an interfacial bond between host and doner. The material in this example has been plasma sprayed onto a Titanium aIloy substrate. HA bonds to the substrate by mechanical interaction, not chemical bonding or diffusion, and therefore requires the substrate to be roughened or porous. In this example the interface surface has been grit blasted prior to coating. It could altematively have been titanium powder plasma coated. Tbe sprayed coating should make a good mechanical bond to the interface surface, the spraying parameters controlling the porosity level. Tbe requirements of the coating could be fully dense or open structured. Tbe micro-sectioning technique will reveal the integrity of the coating manifesting any characteristics or features undersirable to its performance in use.

219

Surface Preparation Technigue

(a) Sectioning: Use of a thin blade on a low deformation cut-off machine is required. Suitable blades would be the high speed abradable resin bonded silicon carbide or the low speed cubic boron nitride (CBN) wheel. Always section from the direction of coating into the substrate and do not clamp the portion designated to be the sampie.

(b) Mounting: It is necessary to encapsulate the sampie to protect the coating. Vacuum Impregnation using a low viscosity (250 cPs) epoxy resin is necessary with porous coatings. Compression mounting is not to be recommended for these materials unless a protective cold mounted film is first applied.

(c) Grinding: Directional preparation with a small abrasive on progressively reducing shock absorbing surfaces will reveal a faithful reproduction of the microstructure. The method chosen for this material is illustrated in TABLE 1.

AUDITABLE PREPARATION PROCEDURE EXAMPLE No 1

CLASSIF.lCQA'JTIEO~~~S~=IP"'M=ENr=-CTYP=""E ---=BlAD~=E--------L!!!~~~~CENfRA:!RIA~L'''!W~O:gN!!lRO;UXY[X6APrlAU1TrE!l&---------SECTIONING ISOMET PLUS RESIN BOND SiC TECHNIQUE,

METIlOD SAMPLS/MOUNT RATIO MOUNTING VACUUM IMPREGNATED EPOXY

~~:~~E~'M~O~UMreD~~D~I~~=O~N~ALL~Y~-----------------------------

ABRASIVE Z IMAGE

!~~? ~~~ FORCE HEAD AXlS I~VE ~~~RAPH ~~~IS REMARKS kPs ROTATION um

PlANAR GRINDING Sie P240 35 CONTRA PLANE STAGE WATER

ABRASIVE Z IMAGE SURFACE ~~ I~RCE HEAD AXIS REFLECrIVE I ~~~~RAPH ~~bsIS REMARKS & TYPE ROTATION 1= FACTOR

SAMPLE S No 4 9,.., OlL 35 CONTRA 24

INTEGRITY S "'0 I ,.- '1IGH DH 10 CONTRA

STAGE

ABRASIVE IMAGE SURFACE ~~NT FORCE HEAD TIME REFLECTIVE !~~~RAPH ~~bsIS REMARKS & TYPE kPs ROTATION SECS FACTOR

POLiSHIN"G ~~IDALSIU"" I VF.< P.S. No 7 17 CONTRA 120 STAGE

TABLE 1 - PREPARATION PROCEDURE/PLASMA SPRA YED COATING

Auditable preparation procedures will require completion of all monitoring parameters shown to the RHS of the thick central line. It is also assumed machine operating conditions such as surface speed will be optimised. Time is not inc1uded in the method since this will vary dependent upon the size of sampie, ratio of mounting resin, number of sampies in a specimen holder, efficiency of grinding parameters etc. Auditing, therefore, utilises a measured material removal, (Z-axis), reflectivity factor, photomicrographs, image analysis and such factors as scratch pattern at each preparation stage.

220

(A) 8RIGHTFIELD X2 (8) DlC X2

MICROGRAPH NO.l - OPTICAL OBSERVATION

r C: II """ -The analysis of the prepared surface requires that the surface be a true reproduction of the microstructure, presented in the most favourable optical condition. From micrograph No. 1 the coating has been displayed in (A) Brightfield to show the planar information and (B) 1st order grey DIC to give an in-depth picture of the same field of view. Notice the radial cracks, probably caused by differential thermal contraction or thermal shocking. The porosity level for this coating, judged on this image, is extremely low; nearly zero. Unmelted particles are also in evidence. Close inspection of the titanium/hydroxyapatite interface reveals a dark unresolved line. This line when imaged with a higher resolving objecdve (200X Apochromalic) proved to be undesirahle fractured particles.

Example No. 2 - Coral

Two different species of coral, porites (micrograph 2A) and acropora (micrograph 2B) were prepared as part of an investigation into microporosity.

(A) PORITES X25 (8) ACROPORA X25

MICROGRAPH NO.2 - MICROPOROSITY IN CORAL

221

The micrographs No. 2 A & B c1early illustrate the vastly different porous structures of the two specimens. Unlike the pores with interconnecting channels which occur in plasma coated materials, coral as it grows creates a honey comb structure with continuous tunnels slowly varying in size and shape. Black shapes within the micrographs, therefore, illustrate areas of poor resin ingress. This can be overcome by using a lower viscosity resin or increasing the vacuum pressure during impregnation.

AUDITABLI! Plll!PAllAnoN PROa!DUIU! EXAMPLI! No 2

'-BOUIPMENr TYP!! BtAD! COIlCllNfRATION

MI!1lIOD SAMPLl!lMOUNT RATIO ,!~: VAaJUM IMPlU!ONATED EPOXY

ABRASIVE Z IMAGB

~CE ~ ,~RCE '=11ON ~s ~ I~ =~ RI!MARKS

PLANAR GRINDING SiC HOO 35 COMP PlANE STAGE WATER

ABRASIVE Z IMAGB

~CE ~ ,~RCE '=11ON I~ ~ I~ =~L!i RI!MARKS

SAMPLI! SN. 1 ~~m. "" ..... 35 , CONI'RA

~turY 6

STAGE

ABRASIVE IMAGB

!~CE ~vr I~RCE 1~"nN I~ I~ I~RAPH I=~ RI!MARKS

PoUSHING P~. No 7 =mALSILICA 20 360 , YBS

~AGE

TABLE 2 - PREPARATION PROCEDURE/CORAL

The procedure for coral looks and is very simple. This however does not necessarily imply any alternative method would yield the same information. To use, as is traditional, a diamond abrasive in the integrity stages, is counter productive due in part to the highly abrasive nature of coral. The chemomechanical technique as used removes any residual damage from the silicon carbide preparation stage and if pursued into the polishing mode will reveal structural information relating to the coral.

CRYSTALLINE OR AMORPHOUS STRUCTURES

Just as with the question of high density or controlled porosity affecting interface bonding, so does the implant interface surface condition. The dissolution of the implant material, the tissue growth relative to crystalline or amorphous structures have also to be addressed. This information, the microstructure from clinical trials, will reveal.

222

Micrograph No.3 is used to illustrate the different structural phases from the interface with a 6O.um outer layer, with its circular 10 to 20 .um features, to the 300.um pore surrounded by a modified enlarged structure. The technique for optically revealing this information is to use the 1st order grey D.I.C. (3) as used in micrograph No lA; preparation method No. 2.

CONCLUSIONS

PHOTOMICROGRAPH o. 3 -CORAL TRUcruRE.

This paper has addressed the importance of surface implant preparation in the understanding and development of future material combinations. The need to relate the preparation method to some traceable reference, along with the ability to manifest information through chemomechanical and optical techniques is also recognised. It is hoped to have illustrated the importance of cross-sectioning technology to broaded our understanding towards implant receptibility.

References

(1) BOUSFlElD B. A Systematic Approach to SampIe Preparation. Metals & Materials 1988 - 4

(2) BOUSFlElD B. BOUSFIELD T. Progress Towards a Metallographie Standard Metals & Materials 1990 - 6

(3) BOUSFlElD B. High Resolution Photomicrography Metals & Materials 1989 - 2

Acknowledgments

(1) Buehler Europe Ltd. for permission to reproduce the photomicrographs.

(2) Professor D. F. Williams, University of Liverpool, for supplying the coral material.

223

YTTRIA AND CALCIA PARTIALLY STABILIZED ZIRCONIA FOR BIOMEDICAL APPLICATIONS

P.FASSINA, N.ZAGHINI, TEHAV, 40059 Hedicina (80) Italy A.BUKAT, C.PICONI, 00060 CASACCIA - ROftA Italy F.CRECO, S.PIANTELLI, CA11IOLIC UNIVERSITY, OR11IOPAEDICS DEPT., ROftA Italy

ABSTRACT

Yttria and Calcia Partially Stabilized Zirconia are investigated as materials to be used in the manufacture of prosthetic components as weIl as for coating of THP stems. Ceramic precursors were made by TEMAV SpA using the Gel Supported Precipitation (GSP) process. The GSP (R) process was demonstrated to be very suitable for high purity ceramic precursors production, for the impurities control that can be obtained in each phase of the process itself. Powders have shown very good sintering behavior. Typical final density values are more than 99% T.D. for Y-PSZ and more than 98% T.D. for Ca-PSZ. These results were achieved by low temperature pressureless sintering ( 1500·C ). The paper describes the results of the microstructural characterisation performed both on ceramic precursor and sintered sampies, and some preliminary results of the cytocompatibility tests. These tests. made at the Catholic University by Inhibition 3H Thimidine Uptake Test using human lymphocytes. show little or no inhibition of lymphocytes proliferation from Ca-PSZ. Results of cell adhesiveness tests, performed on sintered plates, show a good adhesion of 3T3 fibroblasts on Ca-PSZ. Tests on Y-PSZ are in progress.

INTRODUCTION

The search for an implant material has focused on .etals, ceraaics, and plastics. The main advantage of ceraaics over the other materials is their better biocompatibility. Nearly inert materials. such as dense Alumina. undergo minimal alteration of the implant surface: the tissue response is usually no fibrous capsule or only several cells layers thick adjacent to the exterior surface of the implant. Stabilized Zirconia, where the stability of the tetragonal and cubic phases after sintering is obtained by adding small quantities of additives such as yttrium or calcium oxides. has the potential advantage over Alomina of a lower Young's Hodulus. higher strength. better wear properties and higher fracture toughness [I. 2, 3]. To investigate the behavior of Ca-PSZ and Y-PSZ in the bio.edical field. a joint program was started in Italy by ENEA, the Catholic University of Rome and TEHAV SpA in the framework of an agreement for a Research and Development Project on the ceramic materials.

224

The purpose of the present work was to manufacture batches of Ca-PSZ and Y-PSZ powder, by a modified sol gel process, and sintered saaples for biocompatibility testing, to characterize the materials and to perfona the tests whose first results are reported in this paper.

EXPERIMENTAL PROCEDURES

Materials Ceramic precursors were .ade by Gel Supported Precipitation process, (GSP) (R), a modified sol gel process developed since 1970 by SNAM, Italy, for the production of U, Th, Pu oxides and carbides .icrospheres to be used in the HTR nuclear fuel [4, 51. The GSP process was modified to produce high performance ceramics as structural or biocompatible advanced ceranlic materials [6, 71. The GSP process has been described elsewhere [81. To achieve higher densities than the ones previously reported, the powders were treated by spray drying after the calcination step. The starting materials were:

Zr(N03)4 (20 ~ ZrOa ) fram Magnesium Elektron LTD . CaO from Carlo Erba . YZ03 from RhOne-Poulenc Chimie

Fo~ing and sintering Disks for adhesiveness tests were obtained by cold axial pressing followed by pressureless sintering at different temperatures to achieve different densities. Preliminary tests have been performed on the GSP-made powders to set up the manufacturing proeesses able to produee a final prosthetic component such as femural head.

Characterisation The fol10wing investigation methods were used in the characterization of the powders and sintered bodies:

Low-power mieroscopy for macroscopic visual inspection SEM and light microscopy for microstructural analysis EDS and XRD to determine chemistry homogeneity and crystal structure Spectrographic Analysis for impurities eontent BET speeifie surface area Image Analysis system (QUANTIMET 520) for particle size distribution NETZSCH Simultaneous Thermal Analyzer STA 409 and Dilatometer 402E for thermogravimetry (TG), differential thermal analysis (DTA) and dilatometry analysis to define the powder behavior during the sintering.

The green density and the final densities (geometrie and by water immersion) were determined on all the sampies.

Cytocompatibility tests The first tests were performed in-vitro using eell culture method, thus providing a rapid and cost-effeetive screening for the material [91, although it is clear that it is rather difficult to make an extrapolation from the in-vitro biocompatibility tests to the in-vivo situation [101. Ca-PSZ ceramic precursors were tested on human blood mononuclear cells coming from healthy voluntary donors to avoid differenees from species. Tests were performed using Inhibition 3H Thimidine Uptake Test during PHA induced human lymphoeytes proliferation [111. This test is used for the evaluation of the human illllllunoresponse and it measures through inhibition of thc blastic phase (3H Thimidine eaptation) the toxie effect of the ma teria I on the ce 11 s in contaet with it.

225

Tests on Y-PSZ cera.ic precursors are in progress. We also studied by SEM short-term in-vitro assays [12] for the adhesion and the spreading of 3T3 fibroblasts cultured on low density (D<70~ TD) Zirconia disks. As control, the same cells were grown on glass or plastic surfaces.

RESULTS AND DISCUSSION

PoMders The properties of the powders are shown in Tables 1 and 2. The data indicate that these GSP-.ade powders have a very highly specific surface area and a very low i.purities content, depending only on the che.ical precursor purity and not on the che.ical route .

TABLE 1 Physical properties of Ca-PSZ and Y-PSZ powders

Powder

Ca-PSZ Y -PSZ

118 98

Average particle size (0)

115 53

Phase (by XRD) t-ZrOz c-ZrOa

FIGURE 1. SEM analysis of Y-PSZ powder

Figure 1 shows the spherical shape of the particulates that allows an high flowability of the powder with a very good behavior during the pressing . Une scans, for Calcia and Yttria respectively, show a good homogeneous distribution of the elements (Fig . 2 for Ca-PSZ).

226

TADLE 2 Typical i.purities content of Ca-PSZ and Y-PSZ powders

Elellent Al Fe Hg I1g Pb Zn Y Si

Content <140 <300 <5 <20 <100 <20 <100· <500 <l1II/g) • 5.5% wt for Y-PSZ powder

FIGURE 2. Calciua line scan for Ca-PSZ

FIGURE 3. SEM analysis of Ca-PSZ sintered sampIe

Cl (eguiv) <1000

227

Sintered sa.ples SampIes with three different densities (low, medium and high) were obtained for in-vitro adhesion tests. The fraeture surfaee of the highest density Ca-PSZ sampIes shows a fully sintered strueture where the tetragonal phase (D:I-2 ~m) is assoeiated to bigger eubic grains (D=6-8pm) (Figure 3). The X-Ray diffraction confirms the presence of the solid solution C~O.15ZrO.a501 . 85 in the cubic phase with traees of tetragonalone. The metastable-cubic phase of the Y-PSZ powder was transformed into tetragonal during the sintering. The presenee of a metastable-eubie phase suggests that the powder has high compositional homogeneity . The tetragonal grain size is very fine (D < 0.9pm) (Figure 4).

FIGURE 4. SEM analysis of Y-PSZ sintered sa.ple

Cytoca.patibility tests As shown in Figure 5, the inhibitory effeet of Ca-PSZ on lymphoeytes proliferation never exeeeded 50~ even at the higher eoneentration tested (i.e. 1.2 g/100 roll. On the other hand, in comparison with Ca-PSZ, Ti02 was more toxie at all coneentrations (from O.06~ to 1.2~). As early as after 12 hours of culture, the great majority of the cells adhered on the Zireonia surface (see Figure 6). In addition, the cellular behaviour of fibroblasts cultured on Zirconia is Qualitatively similar to the behaviour on control surfaces. These preliroinary observations indicate that this new implant material is adhesive for cells and does not show any adverse reaction.

CONCLUSI ONS

Two materials were inQuired as alternative biocera.ics to the most used Alumina in orthopaedic application. The possibility to produce these materials by the GSP proeess with homogeneous additive distribution, low impurities content and high density structure with controlled grain size

228

has been demonstrated. The preliminary in-vitro tests showed a lower PSZ inhibitory effect on lymphocytes proliferation than TiOz as reference materials, and a good adhesive behavior of fibroblasts on the Zirconia surface without any significant adverse tissue reaction.

120 oCaPSZ

100

..a 80

I eo

'" 0 «> .. 20

0

0.0 0.2 0.4 0.6 0.8 1.0 1.2 1.4

" (./v)

FIGURE 5. Concentration dependence of Ca-PSZ and TiOz inhibitory effect on PHA-induced human PDMC proliferation

FIGURE 6. Cell adhesion on the Zirconia surface

229

1. P.Christel, A.Heunier, H.Heller, J.P.Torre, C.N.Peille, Hechanical Properties and Short-Term In-Vivo Evaluation of yttrium Oxide Partially Stabilized Zirconia. J.Biomedical Hat. Res. ,1989, 23, 45-61

2. R.C.Garvie, C.Urbani, D.R.Kennedy, J.C.Hc Neuer, Biocompatibility of Hagnesia-Partially Stabilized Zirconia (Hg-PSZ) Cera.ics. J.Hater.Sci., 1984, 19, 3224-28

3. P.Christel, Biomechanics, CUrrent Interdisciplinary Research, S.H. Perren • E.Schneider Eds., Hurtius Nijhoff, Dordrecht, Nt, 1985

4. G.Brambilla, P.Gerontopulos, D.Neri, The SNAH Process for the production of ceramic nuclear fuel microspheres: laboratory studies., Energia Nucleare, 17, (1970) 217

5. P.Gerontopulos, D.Neri, R.Renzoni, USA Patent N°3728274 (1973)

6. A.Bukat, P.Fassina, F.Greco, C.Piconi, N.Zaghini, Bioceraaics by Gel Supported Precipitation Process. 7th CIMTEC, Sat. Symp. 3, Hontecatini (Italy), 1990 (in print).

7. A.Bukat, P.Fassina, F.Greco, S.Piantelli, C.Piconi, N.Zaghini, GSP-made Ca-PSZ for Biomedical Appl icat ions , 3rd Symp. on Ceramies in Hedicine, Terre Haute (USA), 1990, (in print).

8. P.Fassina, N.Zaghini, C.Alvani, Influence of Gel Supported Precipitation (GSP) process parameters on the PSZ powder sinterability. 7th CIMTEC, Hontecatini (Italy), 1990 (in print).

9. T.Inoue, J.E.Cox, R.H.Pilliar, A.H.Helcher, Effect of the Surface Geometry of Smooth and Porous-Coated Titanium Alloy on the Orientation of Fibroblasts In Vitro., J.Biomedical Hat.Res., 1987, 21, 107-126

10. C.J.Kirkpatrick, C.Hittermayer, Theoretical and Pratical Aspects of Testing Potential Biomaterials In Vitro. J.Hat. Sci/Mat.Medicine, 1990, I, 9-13.

11. S.Piantelli, P.Tranquilli-Leali, G.Lorini, H.Piantelli, Effetti inibitori di alcuni metalli di interesse ortopedico sulla trasformazione blastica dei linfociti umani. Min. Ort., 1983, 34, 1

12. R.A.tizarbe, N.Olro, J.G.Gavilanes, Adhesion and Spreading of Fibroblasts on Sepiolite-Collagen Complexes. J.Biomedical Hat.Res., 1987, 21, 137-144

230

MECHANICAL PROPERTIES OF PLASMA SPRAYED CERAMIC COATINGS ON ORTHOPAEDIC IMPLANTS

1. RIEU*, H. CARREROT*, J-L. AURELLE**, A. RAMBERT**, G. BOUSQUET*** * Department of Mechanics of Biomaterials, Ecole Nationale Superieure des Mines,

158 Cours Fauriel, F-42023 Saint-Etienne cedex 2, France ** SERF, 85, Chemin des Bruyeres, F-69152 Decines Cedex, France *** Department of Orthopredics, CHRU Bellevue, Boulevard Pasteur,

F-42023 Saint-Etienne Cedex 2, France

ABSTRACT

There are many contradictory results in the literature concerning the mechanical properties of ceramic coatings, especially the adhesion strength, for which an order of magnitude may exist between the values given by differents authors, for the same kind of coating. Because of the numerous parameters governing the plasma technonogy, it is weIl known that coatings may differ from one laboratory or manufacturer, to others. Moreover, the adhesion strength is not clearly defmed and depends on the experimental methods used for mechanical testing. Our aim was to get experimental values on the Same coatin~, measured with several methods. The ideal testing method does not exist. In order to understand the reasons for the scattering on the adhesion strength values, eleven different methods were applied to characterise the same c1assical alumina coating, which has been developped by the SERF company, on the Bousquet's joint prostheses since 1975, on more than 15000 implants.

INTRODUCTION

The increasing development of porous surfaces on cementless prostheses is at the origin of the multiplication of coating techniques. These days, the applications of calcium phosphates generate a proliferation of small industries using the plasma coating process. The quality of the products varles widely because : (i) the ideal coated surface has not yet been specified, (ii) the plasma technique is difficult to control, (iii) the characteristics of the coatings are difficult to measure.

The mechanical properties of a coating can be correlated to its structure. In particular, we have demonstrated [1, 2] that the Young modulus of an alumina porous coating with 30-40% porosity can be as low as 20 GPa. It is then possible to get a low modulus adherent materiallayer on a high modulus metal. The mechanical compatibility is better, when such a coating is in contact with bone.

The present paper focuses on the adhesion and cohesive properties of plasma coatings, and on the different methods proposed for measuring them. Adhesion refers to mechanical, physical and chemical bonds between two adjacent materials, and is measured by the forces necessary to separate them. When the strength of the coating itself is lower than that of the interface, the cohesive strength of the coating is rneasured.

231

Several papers describe methods of testing the adhesion of coatings [3-5]. For orthopredic application ceramic coatings, many different values can be found. As examples, for a 330 micron thick alumina coating, with 2 to 11 % porosity, the adhesion strength, measured by a "standard tensile pull-test", was 26±1O MPa on 316L SS and 24±7 MPa on Ti-6AI-4V ELI [6]. For hydroxyapatite coatings, different values were reported: 5 MPa on 316L SS, and 60-70 MPa on Ti-6AI-4V substrates [7]; 18 MPa tensile and 17 MPa shear strength, for a 75 micron coating [8]; 85 MPa tensile strength on titanium substrate, for a 50 micron HAP coating [9].

MATERIALS AND METHODS

Alumina plasma sprayed 100 and 200 micron thick coatings were performed on TA6V4 titanium substrates, with the industrial METCO computed assisted plasma device set up at the SERF company.

The testing methods can be divided into several families according the state of stress at the interface : tension or shear. In most of them, a plug must be bonded on the coating to subject the assembly to a tensile or shear loading. The crosshead velocity was 1 mm/min.

The following tension methods, were performed: - bonded pin on a flat substrate BP (fig. l-a): 7 mm diameter pins were attached on the

coating, either with an epoxy gIue preparedjust before (type 1),.or with a special glue placed on the pins long time before the test (type 2). The sampies were 70x2Ox5 mm plates. For cal­culating the adherence strength, the load was divided by the fracture area. In some cases, the fracture area was the area of the pin, but in other cases, the fracture area was larger. A special technique with a diffusing die was then used to rneasure the actual fracture area.

- the ASTM method [10] (fig. I-b): 25 mm diameter cylinders, coated on one end, were bonded to loading fixtures of the same diameter, with a suitable adhesive bonding agent.

- a modified ASTM method (M-ASTM) (fig.l-c): in order to clarify the effect of the area of the coatings, 7 mm diameter cylinders were tested in tension.

- fracture energy or energy release rate method (OIc) [4] (fig.l-d): a double cantilever beam is composed of two elernentary beams. One of them was plasma coated and the other was bonded on it, on an area of 13Ox1O mm. An acoustic emission sensor was fixed on the beam. Measurements of the compliance were performed from 40 to 100 mm crack length. The crosshead of the test machine was stopped every time after the crack propagated some millimeters. A calibration was done on specimens with various lengths of cracks. A control was also made at one step of the crack propagation by introducing a colored die to print the shape of the crack at this step.

- tapered pins (TP) [3] (fig.l-e): tapered pins were accurately fitted to tapered holes in the substrate so that the top of the pins and the substrate were in the same plane for plama spraying. The pins were withdrown in tension. Either the coated layer separated from pins, or not. When it separated, the adhesion strength in tension can be calculated. When it did not separate, fracture occured at the rim of the pin by shear in a direction perpendicular to the coating.

The three point bonding test (3PB) [11] (fig.l-f) is intermediate in between tension and shear methods.

Shear stresses are generated in shear, torsion and bending tests. The surface roughness of the substrates may playa very important role: if shear occurs in the coating, the measured shear stress is that of the coating. When it occurs at the interface, it is difficult to discriminate between the part due to the bonding strength and that due to the roughness.

- shear test (ST) (fig.l-g): a shear test device TRIBOMINE [12] was used to perform shear strain at the interface between bonded coated and uncoated IOx15x7 mm sampies.

- torsion test (TI) (fig.l-h): the same 25 mm diameter cylinders as for the ASTM tensile tests were used in torsion, on an aligned device, with a torque cell. From the value of the torque at fracture, the maximum shear stress can be calculated.

- four point bending (4PB) (fig.l-i): very close to an external bearing, the bending moment is approaching zero, whilst the tangential component is equal to half the loading force. At this point, a shear state of stress is obtained on IOxlO mm cross section of the sampies .

232

(a) Bonded pm on fl:n subsD"3te

(b) ASTM (e) Modtfied AST 1

150 )

Encrgy rele:lSe rate

Three POUll bending (f)

(i) Four point bending

(g) Shear (h) TorsIon

Push-QUI Peehng

(j) (k)

Figure 1. The eleven mechanical testing methods performed for measuring the adhesion strength of an alumina eoating on a titanium alloy substrate.

233

- push-out test (POT) (fig.l-j): 10 mm cylindrical pins were partially coated on a length of 40 mm, and bonded with an adhesive inside a 10,3 mm cylindrical hole, before pushing them out. Cylindrical shear occured and shear stress was calculated. Three surface geometries were prepared: regular grit blasting, deep grooves and a helical fme threading.

- cylindrical peeling (CP) (fig.l-k): the same pins as for POT were pushed into a cylindrical hole machined in a hard material die, with a minimum clearance. The cutting edge of the hole peeled out the coating from the pins.

RESULTS

All the experimental results are reported in table 1. Oenerally, the fracture was always adhesive. However, some comments must be done. For the BP tests, and the 100 micron thick coated plates, four values from 2 to 6 MPa

were eliminated because of the quality of the glue layer or mis-alignement For the 200 micron thick coated plates, the fracture propagated either at the interface

(adhesive) or in the coating (cohesive). Two different behaviours were observed: cohesive fracture in the coating, or at the limit of two layers in the coating. This is because we chose the same displacement program of the plasma system, with a lower velocity for the thicker coatings. The temperature exchanges were different and may explain the delamination between the two successive layers, for thick coatings.

TABLEI Adhesion strength for different mechanical tests

N lUUnncrons N 200microns BP bondedpins 7 mm type 1 18 12 ±2 MPa 18 8±4MPa BP bonded pins 7 mm type 2 18 1U±3MPa 18 8±4MPa A:sTM~mmm 5 14.5 ±.3 MPa 5 10.5 ±1 MPa M-ASTM7mm 5 18.5±1 MPa 3 12.5 ±4 MPa OIc energy release rate 3 35-50 J/m2 4 25-35 J/m2 TP tapered pins 6 7±2MPa 6 4.5 ±2 MPa ST shear test 3 24±2MPa 5 25±2MPa TI torsion test 25 mm 3 36±3MPa 3 32±2MPa 4PB four point bending 4 29 ±6 MPa 5 25 ±6 MPa POT push-out grit blasted 3 30±3MPa 0 * POT push-out grooved 3 33±5MPa 0 * POT pus!t~ut th!eade 3 27±7MPa 0 *

I CP cylindrical peelmg 5 100w MPa ./ 0 * N : number of sam >Ies; * not measured. p

In the TP tests, fracture was never cohesive and occured at the rim of the pins. The load did not correspond to the adhesion strength, but to the shear strength of the coating, in a direction perpendicular to the surface. The fracture area was a cylinder with a diameter of 7 mm and a thickness equal to the coating thickness.

In the POT tests, the maximum compression load corresponded to the beginning of the push-out movement of the pin. This load was used to calculate the shear stress. In most of the tests, fracture occured at the interface hole-adhesive agent, and not in the coating or at the coating-metal interface.

The CP method could be interesting because there is no adhesive agent between the sheared components. However, it was not possible, at that time, to get a constant peeling load and the results were not reproducible. Shear stresses from 10 to 60 MPa were obtained.

234

DISCUSSION

When analysing all the experimental adhesion strength values, some general results can be established :

- Effect of thickness: in all the experimental methods, the adhesion strength of the 100 micron thick coatings was higher than that of the 200 micron thick. This can be due to the residual stresses created during the coating formation. In the plasma coating process, melted droplets of ceramic solidify in contact of the metallic substrate or in contact with the previous solidified droplets. The temperature conditions are not the same for the successive solidified droplets.

- Effect of size: the larger the size of the specimen, the lower the adhesion strength. This can be explained by the probability to find a defect with a critical dimension to give a brittle fracture. A second kind of size effect can be observed when performing a 7 mm BP test on a 25 mm ASTM sampie. Generally, the measured strength is then lower, whilst the area is smaller. Stress concentrations can be at the origin of this difference.

- Effect of mode of test: the shear strengths are always higher than the tensile ones. Several mechanisms interact, but the surface roughness of the substrate must play the main role. The push-out tests are as difficult to interpret, as for the bone-implant analysis.

CONCLUSION

This study is a first experimental approach of the complex question of the adhesion strength of ceramic coatings. Using eleven different experimental methods we demonstrated that the large scattering in the literature can originate, not only from the coating processes, but also from the experimental technique used by the authors to caracterise their coatings. A complementary mechanical analysis of stress is necessary. But, already, we would recommend to the authors to describe clearly their measurement method before giving any value of the adhesion strength.

AKNOWLEOOEMENTS

We wish to thank all the people from SERF company for the preparation of hundreds of sampies and controlled plasma sprayings; Jean-Marie Delecroix, from Ecole Nationale d'Ingenieurs de Saint-Etienne and Etienne Bouyer, from Institut National des Sciences et Techniques Nucleaires de Saclay, for their helpful contribution in this important experimental program, and Vincent Rieu, for the computed assisted figures and slides.

REFERENCES

1. Carrerot, H., Rieu, J., Rambert, A., Bousquet, G.and Girardin, P., Mechanical Properties of Porous Alumina Plasma-Sprayed Coatings on Metal Stems for Joint Prostheses. In Ceramic in Substitutiye and Reconstructiye Surw:y, Proceed. 7th ClMTEC, World Ceramic Congress, Montecatini, 27-30 June 1990.

235

2. Carrerot, H., Rieu, I., Rambert, Aand Bousquet, G., Alumina Plasma Spray Coatings on Stainless Steels and Titanium Alloys for Prosthesis Anchorage. In Bioceramics, vol.2, ed. G. Heimke, German Ceramic Soc. Publisher, 1990, pp. 211-18.

3. Oavies, O. and Whittaker, I.A, Methods of Testing the Adhesion of Metal Coatings to Metal. Metall. Rey., 1967, 12, pp. 18-25.

4. Colin, C., Boussuge, M., Valentin, O. and Desplanches, G., Mechanical Testing of Plasma Sprayed Coatings of Ceramics. I. Mater. Sei., 1988, 23, pp. 2121-28.

5. Ingham, Ir., H.S., Adhesion of Flamme-Sprayed Coatings. In Adhesion Measurements of Thin Films, Thick Films, and Bulk Coatings. In STP 640. ed. Mittal, K L., ASTM Publisher, 1978, pp.285-92.

6. Brown, S.O., Orummond, I.L., Ferber, M.K., Reed, O.P. and Simon, M.R., In-vitro and in-vivo Evaluation of Melt-sprayed Alumina Coatings on 316L Stainless Steel and Ti-AI-4V EU Alloy Substrates. In Mechanical Pro.perties of Biomaterials. ed.G.W. Hastings and D.F. Williams, lohn Wiley & Sons Ltd, 1980, pp. 249-64.

7. Gross, K.A., Haddad, G.N. and Berndt, C.C., Thermally Sprayed Coatings for Orthopredic.Applications. In Proceed. 12 th Internl. Conf. on Thermal Spra,yinl:. ed. I.A Bucklow, Abington Publisher, 1989, pp. 153-61.

8. Thomas, K.A, Kay, 1.F., Cook, S.O. and Iarcho M., The Effect of Surface Macrotexture and Hydroxyapatite Coating on the Mechanical and Histological Profiles of Titanium Implant Materials. I.B.M.R., 1987, 21, 1395-414.

9. de Groot, K., Geesink, R., Klein, C.P.T.A. and Serekian, P., Plasma Sprayed Coatings of Hydroxyapatite. I.B.M.R., 1987,21, 1375-81.

10. C-633 - 79 Standard Test Method for Adhesion or Cohesive Strength on Flame-Sprayed Coatings. In Annua,l Book of ASTM Standards, 1985,03.01, pp. 653-57.

11. Roche, A.A, Behme Ir., AK. and Salomon, I.S., A Three-point Flexure Test Configuration for Improved Sensibility to MetaVAdhesive Interfacial Phenomena. Internl. J. of Adhesion and Adhesive, 1982,249-254.

12. Lopez, A, Rieu, 1., Mechanical Behaviour of the Ceramic-Metal Interface in the Femoral Head Prosthesis Conical Fittings. In Bioceramics, vol.2, ed. G. Heimke, German Ceramic Soc. Publisher, 1990, pp. 166-71.

236

PLASMA-SPRAY COATING OF TI'I'ANIUJI SUPPORTS WITH VARIOUS CERAMICS:

ASnmYMnmI~MACE.

A. Krajewski, A. Ravaglioli, V. Biasini, A. Martinetti, A. Piancastelli

Istituto di Ricerche Tecnologiche per la Ceramica deI CNR - Faenza (Italy)

S. Sturlese, S. Fioravanti

Centro Sviluppo Materiali (C.S.M.) - Roma-Pomezia (Italy)

* § S N. Antolotti , C. Mangano , F. Trotta

* Flametal S.p.A. - Fornovo Taro, Parma (Italy) § Poliambulatorio Odontciatrico Venini - Colico, Corno (Italy) S Plasma Technik A.G.- Wohlen, Zurich (Switzerland)

SUMMARY

The covering of surfaces of pros thesis is today usual. This study takes into consideration ceramic coverings on titanium substrates to carry out a comparisoll among them about the characteristics of the interface cera­mic/titanium. Dental rootE cf titanium were utilized as suitable samplings to cover. The quality, or high-grade, of the covering expressed in terms of joining adhesion of the ceramic covering to the metallic substrate is connected to the thickness, to the porosity, to the joined area and to the thermodynamlc state or complexity (nwnber of crystallographic phases) of the ceramic layer.

INTRODUCTION

A glassy or a ceramic coating of metal or ceramic support is generally

designed to protect the support from chemical attack or mechanic friction.

But in specialized technological applications, including biomedical implant

ones, such co~ting ml.lst perfoI'll1 the additional task altering the chemi cal

and physico-chemical properties that a surface of a coated piece assumes

when it is interfaced with another system.

A coating is applied by disparate techniques. In each case the important

thing is to pre-treat the surface to be coated (1). Such preparatory treat-

ment in fact, in combination with the characteristics of the covering mate-

237

rial in itself and those of the bond which that material establishes with

the support, plays a part in improving the mechanical adhesion strength.

MATERIALS AND J1KTHODS

The supports to be coated were cylindrical sampies of grade-2 titanium

sized 2.5 cm in diameter and 1 cm in height whose flat sides were polished

after being rendered parallel to one another and at right angle with respect

to the cylinder axis. The coating was performed on the flat faces by plasma

spraying. Preparatory to coating, the faces were degreased by treatment with

a solution of 3% HN03 and 1% HF and they were subsequently sand-b1asted with

spheroidal Alumina particles (1). The surface pollution which resulted from

the presence of embedded Al ° was rated at about 3%. The plasma spray 2 3

applications were made in air using as carrier an Ar/H2/N2 gas mixture. The

utilized powders were A1203 PT 1001, Zr02 (stabilized by 8% Y203) PT 1086

and AKRA 15 biological glass j as far as hydroxyapati te was concerned, the

powders utilized were of white type and DAC-blu, both commercial products.

All sampies were treated in the same manner. Apart of them was subjected

to morphological examinations at the interface between coating and substra-

te. Preparatory to such examinations, the sampies were cut perpendicular to

their surfaces and the cut-off portions were then duly polished. The analy-

ses at the interface were conducted by scanning electron microscopy in as-

sociation with microprobing. The remaining specimens were put through

scratch-test to assess their tensile strength. Before use, all samp1es were

examined by X-ray diffractometry to locate the different crysta1 phases.

RESULTS

X-ray examination on the coverings gave evidence that: no significant

amounts of crystalline phases were to be seen in the bioacti ve glass j the

hydroxyapatite did not appear to change its nature, at least as far as the

determinable phase was concernedj the A1203 exhibited a considerable extent

of )' phase (about 70% )'-phase, 20% ,,-phase, 10% Q- and e-phases with

trace of X-phase). Careful examination (2,3) on the Zr02 covering did not

show it has undergone appreciable phase variation after depositionj it shows

presence of cubic phase wi th some uni t percent of tetragonal and trace of

monoclinic phases.

238

Scanning electron microscopy (Figures 1 to 4) detected in all sampIes a

widespread occurrence of close porosity and intergranular microcracks. The

clustered grains, generally roundish, were arranged in striated heaps. In

bioactive glass layer such heaping appeared to induce formation of bubbles

as a result of sand-blasting treatment and/or air trapping or development of

02 gas following an oxidation-reduction reaction.

From microchemical analysis of the layer already examined under a microprobe

i t appeared that nei ther the Al203 nor the Zr02 coverings had undergone

appreciable changes in composi tion. On tl'ie contrary, whi te hydroxyapati te

had become poorer in phosphorous content (from a ratio Ca/P = 1.67 in the

Figure 1 - SEM microphotograph of interface between A1203 and Titaniua substrate.

Figure 3 - SEM microphotograph of interface between Hydroxyapati te and Ti taniUIII substrate

Figure 2 - SEM microphotograph of interface between zr02 and Titanium substrate.

Figure 4 - SEM microphotograph of interface between AKRA15 bioactive glass and Titanium substrate (cOllbination of secondary and reflected electronsl.

239

Coworl .. Thic:t .... _Ity Joir ........ MusenlA:M .. t ..:10 material (!'ßI) (S) (S) lIreßIIh (N) ('10·""· ')

AllO) 220 14 90 45.8 (:1;7.6) I

zrO, 200 S 8S 6S.1 (:I;S.6) 3

HAIMC-blu 100 S SO 435 (:1;9.9) 4

HAwhite 2SO 12 14 29.2 (:1;9.2) 4

AKRA ISO 45 70 23.0(:1;85) 7

..:Ia - dillilrencc '*- thennallinear expansion codIicient VII ... 0( ccramic coverI .. .-rial .nd lilanium rapocti>ely.

in the powderes to a Ca/P = 2.1 in the sprayed layer) while the DAC-blu

hydroxyapatite had not (from a ratio Ca/P = 1.67 in the powders to a

Ca/P = 1. 68 in the sprayed layer), while the bioacti ve glass had become

poorer in its sodium content.

Table 1 exposes values concerning the porosity, the thickness, the propor-

tion of joined areas at the interface on the basis of a variety of scanning

electron microscopy micr-ophotographs and the scratch-test values . An intere-

sting microporosity was observed in the hydroxyapatite layers (Figure 5),

Figure 5 - Secondary electrons SEM micropho­tograph oC a magniCied section area oC white Hydroxyapati te covering where grain boundary microcracks are diCCUsively distributed.

Figure 6 - SEM .ü:rophotograph oC a deCect displayed at the interCace between DAC-blu Hydroxyapatite cerIlIlie layer (H) and Tita­niu. substrate (M).

especially the DAC-blu ones. Hydroxyapati te ceramic appeared gathered in

local clusters, with development of porous nests of roundish shape (Figu­

re 3) having diameter between 5 and 20 ~m (prevalently around 10 ~m) . Such

nests were spaced some 30 ~m from each other on average . When they occur

at the interface, they give rise to defects like that of Figure 6 . The po-

res are predominantly elongated, up to 15 ~m long and no more than 2 ~m

240

wide (similarly to that of Figure 6). As shown in Figure 3, the porous nests

are accompanied by large Dores sized between 5 and 30 ~m (prevalently around

15 ~m) in diameter.

Scratch-test value can be related to that of maximum tensile strengr.h at the

interface (as indirect measurement for interfacial tearing test); it gives

a cri tical load that is obtained connected wi th the first scale formation

that bares the metallic substrate. To have an idea about the meaning of the

obr.ained values. it is to be considered that for instance DAC-blu hvdroxy-

apatite covering gives a value of 50 (±5) MPa following ASTM 633/C standard

test.

DISCUSSIONS ARD CONCLUSIONS

Wi th reg!lrd to the defects noticed wi thin the body of the deposi ted

layers, a certain porosity end e presence of intergranular microcracks were

observed in all the ceramic layers of the four mRterials examined, despite

the fact that the sRmples were used in operational condi tions judged the

best in the circumstances. While in the more refractory ceramic layers ana-

lyzed (A1203 and zr02 ' biologically inert) there where no variations to be

noted in the chemical composition, bioactive glass and white hydroxyapatite

revealed a variation probably due to sublimation of apart of the more vola-

tile components occurred on hot application. Regarding the interface, all

sampIes displayed a number of defects similar to those developed inside the

ceramic layer. Certain loosenings of the ceramic coating from the metal sub-

strate might derive from an interplay of mechanical tensions which form in

restricted areas on cooling as a consequence of the different thermal expan-

sion coefficients (see Table 1).

Other factors playing a role in the formation of defects at the interface

are those connected with the presence of a third component on the metal sur-

face after sand-blasting. One such component is shown in Figure 7, picturing

a residual A1203 grain remained embedded in the metal surface after sand­

blasting. The figure exhibits a corresponding loosening about 80 ~m long at

the interface in the deposited material (in this case bioactive glass). The

cause of such loosening may be due to the surfaced Al203 grain towards which

the wettability of the deposited material during the molten stage is diffe-

241

rent from the wettability of the same

material towards titanium, a fact

which subsequently can also result in

crack expansion during the cooling

process. In the ZrO 2 covering there

can be noted (Figure 2) whi te and

dark stripes as weIl as dark areas.

Whi te striated areas are made up of

light lamellar grains containing Fe, Figure 7 - SEil microphotograph of' interf'ace

Ta and Ba, while dark areas, made up

of dark grains of roundish shape,

between biological glass and Ti taniUID sub­strate where a detachaent f'ailure is present in corrispondence of' a stuck A1203 grain.

contain varying quantities of Si, Y, Ca, Fe and Mg. The presence of Y, Ca

and Mg may originate from the separation of these elements from Zr02 grains

(within which they are phase stabilizers) as a result of the absorbed heat

on application of the plasma flame. With regard to hydroxyapatite, but only

the white one, the cathodoluminescenceimage (Figure 8) relative to the area

shown in Figure 5 displays regions where a certain ex te nt of separation

between Ca and P took place: bright areas have highest P content (propor­

tional to brightness). This is of course a proof of the (at least partial)

transformation of this hydroxyapati te into other calcium phosphates not

detectable by X-ray diffractometry and therefore occurring in an amorphous

form. Such separation is practically not shown on the cathodoluminescence

images of DAC-blu coverings. On

the other hand, Infrared analyses

show the expected spectra of hydroxy-

apatite, free from impurities in the

case of DAC-blu coverings and wi th

some signals coming from CaO and tri-

or tetra-calcium phosphates in the

case of white hydroxyapatite cove-

Figure 8 - Catodo-lUIDinescence SEIl micropho- rings.

tograph of' the SEIl area of' Figure 5 where Al though the quali ty of scratch-test calcium enriched zones (darker areasl are emphasized. analysis is influenced by the thick-

242

ness of the covering layer in a way that it plays a role the more unfavoura-

ble the wider it is, the obtained data are however clearly comparable among

them. From the obtained values it is shown that scratch-test values are at

the same time proportional to the percent of joined area, to the amount per-

cent of porosity within the ceramic layer and to the difference of thermal

expansion coefficient between ceramic and substrate. It remains unknown the

role played by adhesion forces that are established between covering mate-

rial and substrate. The lesser scratch-test value of A1203 ceramic coverings

as to what expected can be attributed also to the low microstructural homo-

geneity of the plasma-sprayed alumina that is a polyphase material.

The wide size of the associated error to the values of hydroxyapatite sam-

plings is explanable by the fact that i t is a far too softer material in

comparison with the diamond ti ps of the testing device. The examined cove-

rings of hydroxyapatite show a lesser scratch-test resistance in comparison

with the other ceramic coverings, but however the whole behaviour of DAC-blu

coverings is certainly better than that of white hydroxyapatite ones. That

may be attributed (besides its lesser porosity) also to its greater chemical

homogene i ty; i t is particularly of interest to observe that the chemical

stability of DAC-blu is far too higher than that of utilized white hydroxy-

apatite, believed to be due to the highest crystalline degree of its star­

ting powders. High crystalline ratios can be consequently reached in the

sprayed layersby optimizing both the hydroxyapatite powder characteristics

and the spraying parameters. By using a modified plasma toreh, which allow

to greatly reduce the degrading of hydroxyapatite, and adopting computerized

and robotized spraying plants, that insure the reproducibility of the pro-

cess, it will be pOGsible to obtain easily highly crystalline layers of pure

hydroxyapatite having Ca/P ratio equal to 1.67 0.01.

The authors thank particularly S. Guicciardi of IRTEC-CNR (Faenza) and V. Adoncecchi of ENI-Ricerche S.p.A. (Monterotondo, Rome) for their help in carrying on scratch-test measurements.

243

LITERATURB

(1) A. Krajewski, A. Ravaglioli, R. Martinetti and C. Mangano; "A Study of the Methodology for Treatment of Titanium Substrates to be eoated with Hydroxyapatite", reported in this book.

(2) P.A. Evans, R. Stevens and J.G.P. Binner; "Quantitative X-ray Diffraeti­on Analysis of Polymorphie Mixes of Pure Zireonia", Brit. Ceram. Trans. J. 83(2)(1984)39-43.

(3) R.C. Garvie and P.S. Nieholson; "Phase Analysis in Zireonia Systems", J. Amer. Ceram. Soe. 55(6)(1972)303-305.

MICROSTRUCTURE AND MICROANALYSIS OF BIOGLASSES AND GLASS-CERAMICS FROM THE MgO-CaO-Pz05-SiOz

SYSTEM WITH ZrOz

J.Ma.RINCON and P.CALLEJAS Instituto de Ceramica y Vidrio, C.S.I.C

Arganda deI Rey, Madrid, Spain

ABSTRACT

Glasses and glass-ceramics formulated from the MgO-CaO-Pz05--SiOz system depicting high crystallization trend of wollasto­nite and apatite phases have been obtained. In some cases CaFa and Zr Oz has been added in order to nucleate the fluorapatite crystalline phase and to improve the mechanical properties. The microstructure of original glasses and glass-ceramics resul ting from controlled thermal treatments has been observed by -TEM and SEM. Likewise, EDX microanalysis have been carried out. The Vickers microhardness and the stress intensity factor (KIe) have been determined.

INTRODUCTION

Since last years, the calcium phosphate materials are usually considered biocompatible and even bioactive, because has been demostrated that new living bone is generated by chemical resortion of the biomaterial. Therefore, glasses glass­ceramics and ceramics containing fluorapatite and hydroxiIapa­tite are adequate materials for substituing or repairing human body parts (1)(2). But, generally these materials though have excellent biomedical behaviour, they depict poor strength and brittleness which limit their use in load bearing medical applications. Several trends have been tested in order to improve the mechanical properties of these biomaterials, viz a) by additions of Alz03 and ZrOz to calcium phosphate glass-ceramics (3). b) by forming a glass-ceramic composite containing apatite + wollastonite crystalline phases(4). In the present study, glasses and glass-ceramics formulated from the MgO-CaO-Pz05-SiOz system and giving rise to wollastonite and apatite crystalline phases have been obtained.

245

In order to control the crystallization nucleation processes and to improve the mechanical properties, ZrO~ additions have been tested in these materials which were microstructural and microanalytically characterized.

EXPERIMENTAL

An original glass has been prepared from the nominal composition: 4.6 MgO-44,7 CaO-16.2 P~Os-34.0 SiO~-0.5 CaFl (wt ~). This glass was prepared from reagent carbonate grade cheminal products for MgO. The CaO and PlOS were added from CaR PO •. 2 R~O and the SiOl was amorphous silica. The components have been melted in zircon crucibles for two hours at 1500 0C in a superkanthal furnace. Zircon crucibles have been used to prevent the corrosion and contamination problems, which usually are produced in other type of crucibles, such as: platinum, alumina-silica ... Melts were poured in brass moulds and some of the batch was crushed and meshed until < 60 f,m. Powders of this original glass have been doped (2 wt ~) with pure partial stabilized zirconia (PSZ) from Hizirco, Hayashi Chem, Ind.

Fig-l. - Original bioglass observed by SEM and TEM.

The composite powder was ball milled in alcohol and then moulded in cylindric bars (5xl0 mm) by isostatic pressing at 200 MPa in aNational Forge Europe machine. Two thermal treatments at 9300C a~d 12000C have been tested. The final products: transparent original glasses and white opaque glass ceramics were observed and analyzed by a SEMfEDX (Zeiss/Tracor-Northem) equipment working at 20 kV. Sampies were coated with a gold layer to facilitate the SEM observation. Carbon replicas also were prepared in order to carried out TEM observations at low voltage in a Phillips EH-300 instrument. XRD was used to identifie the crystalline phases devitrified

in the glass. The Zeta potential was determined with a Zetameter J.Riddick de la Zeta-Meter Inc. by using two types of pB modifiers: BCl-NaOB and CB3-COOB-NB.OB(5). Fracture toughness determinations by Vickers indentations in a Leco Microdurometer has also carried out.

RESULTS AND DISCUSSION

The TEM and SEM observations at high magnification (Fig.1) show that the original glass is phase separated with interconnected droplets lower than 0.2 um approximate size.Some dispersed areas depict white contrast corresponding to they are ZrOz enriched zones. Likewise, the XRD results show an amorphous halo. The glass-ceramic obtained at lower temperature-time(930 oC/4h) thermal cycle is partially crystallized, showing surface crystallization and a large elongated pore in the sampie core. The surface shows directional crystallization, followed by an "explosive" crystallization and the core an homogeneous glassy-cristalline interconnected microstructure (Fig 2a,b,c).

Fig 2 .- a) Bioglass-ceramic obtained at lower temperature, (surface crystallization); b) boundary crystallization close to the glass core; c) the glass co re showing pores produced by etching.

According to the XRD results, this glass-ceramic is formed mainly by fluorapatite and a low percentaje of wollastonite. The A-W glass-ceramic obtained at higher temperature-time ( 1200 0 C/2h) thermal treatment is fully crystallized with also surface and bulk crystallizations. This glass-cera.ic contains a crack network (Fig.3a) wich can be produced by the mismatch expansion coefficient of the apatite-wollastonite crystalline phases Likewise, the (a->ß) wollastonite transformation produces a volume change giving rise to this

247

crack texture. The Fig.3b shows that crystalline microstructure constituted by interconnected apatite (hexagonal rounded),O.5um size, and wollastonite (elongated, 2-4pm size) crystals. The Fig.3c is a more detailed observation obtained by TEM in the same material.

- -.,.--- r SI

.. . l..-:r_'

6:0 0.0

* ----

keV

Fig 3. a) and b) SEM observations of the bioglass-ceramic (1200oC/4h) at lower and higher magnifications and c) TEM observation on the same material, d) and e) EDX spectra of the crystal}ine phases.

248

This microstructure agrees weIl with the XRD results which give a major percentaje of wollastonite with respect the fluorapatite phase. With respect the microanalysis characterization of these materials. it has been possible by the X-ray energy dispersive (EDX) technique to identify the apatite and wollastonite crystals(Fig.3c and d.)

The table 1 shows the results of indentations test; viz: Vickers microhardness (Hv) and stress intensity factor or fracture toughness of the bioglass-ceramics here investigated.

TABLE 1

Microhardness (Hv) and stress intensity factor(KIc) of A-W bioglass-ceramics.

Convencionally moulded materials Hv(GPa) K (HPa m~) IC

original glass ................. 4.46 1.02

glass-ceramic(+ZrO )(930·C/2h) .. 4.69 0.97 2

glass-ceramic(+ZrO )(1200·C/2h). 5.18 1.30 2

Sintered materials

original glass ................. ---- ----glass-ceramic(+ZrO )(930·C/2h) .. ---- ----

2 glass-ceramic(+ZrO )(1200·C/2h). 5.22 6.14

2

It can be seen that ZrOz addition improves both the microhardness and the toughness mechanical properties on glass­ceramie with respect the original and sintered glasses. though the increased values between sintered and original glasses are not very remarkable. These values are in the normal range for glasses here obtained and higher for the sintered glass­ceramie with respect the glass-ceramic obtained for conventio­nal processing.

The Zeta potential of the original glass powder. which can be an important parameter for the sintering processing, reachs 1.6 - 1.7 values in strong media ( HCI-NaOH )and 2.0 value in weak media ( CH3-COOH-NH.OH ), which means that it is possible to leach easily the surface of these glass-ceramics in order to facilitate the bone-biomaterial bonding.

249

REFERENCES

1. Rinc6n, J.Ma and Martinez, F., Los vidrios y materiales vitroceramicos como implantes quirurgicos, Rev .Esp .Cir • Ost., 1984, 19, 77-95.

2. Orgaz, F., Capel F. y Rinc6n, J.Ma., Materiales biocer4micos y biovidrios, Bol. Soc. ~., Ceram, ~, 1987, 26, 13-19.

3. de Mestral ,F. and Drew, R.A.L., Calcium phosphate glasses and glass-ceramics for medical applications. 1, European Cera.ic Society, 1989, 5, 47-53.

4. Kokubo, T.,Ito S., Shigematsu, M., Sakka, S. and Yamamuro T. Fatigue and life-time of bioactive glass-ceramic A-W containing apatite and wollastonite, 1. Mater. Sei., 1987, 22, 4067-4070.

5. Roque, D., Callejas, P. and Rinc6n, J.Ma ., Medici6n de potencial Zeta en un biovitroceramico utilizando diferentes modificadores de pH, Bol, Soc. Esp. Ceram. Vidr.,1991. (printing proofs).

ACKNOWLEDGEMENTS Thanks are due to E. Martin Molna, F. Martlnez and J.C. Saez from the Departamento de Traumatologla y Cirugla Or~oP6dica, Hospital Ramon y Cajal de la S.S., Madrid for.the k1nd co~pe­ration for in-vivo and biocompatibility exper1ments, now 1n progress, which will be published soon elsewhere.

250

THE COLOUR OF ALUMINUM OXIDE CERAMIC IMPLANTS

GERD WILLMANN Feldmühle AG, D 7310 Plochingen, Fabrikstraße 23-29

ABSTRACT

Ceramic implants are manufactured of aluminum oxide ceramies (Alz03) doped with magnesium oxide (MgO). The yellow or brown discolouration following to Gamma-sterilization does not constiute a material defect but rather a quality feature. This ~olour pffect ist explained in the following.

IMPLANT COLOUR"

Ceramies based on aluminum oxide (AI203) have been approved RS implant material since more than two decades. They are particularly used for components of hip joint prostheses and dental implants [1-4].

Alllminum oxi.de cf'ramics are available in various types[5]. Für medical Rpplications, a bioinert material in compliance with ISO 6474 [6, 7] (Thl. 1) is used, which is available from a smRl1 numher of manufacturers, only.

This matf'rial standardized in compliance with ISO 6474 consists of aluminum oxide doped with magnesium oxide (MgO). Acrnrding to the current state-of-theart the rate of MgO cnntained is ahout 0.2 % i.n weight. Doping aluminum oxide wit.h magnesium oxide provides for the grain growth to he controlled during sintering [8-10] and ensures a small grain size.

The rolour of implants produced in this proper way is not white hut rather a slight yellow. This colour is even intensified upnn sterilization of the aluminum oxide implants lI:;;ing GRmma-ray:;; and may turn into yellow-brown to light-brown.

This discolouration is frequently regarded as a material defect which is absolutely wrong. On the contrary, the f'ffe~t of dis~olollration Is a sign for high-quality ceramies which are in compliance with the applicable standard, and will he explained in the following.

~1

COLOUR OF ALUMINUM OXIDE

Natural aluminum oxide is available as so-called alpha-Al203 or corundum in a highly crystallized and pure form as white-coloured material. It is available as red ruby when featuring slight traces of chromium oxide (Cr203) and as blue sapphire when doped with iron oxide or titanium oxide. In their monocrystalline form these materials count among the group of precious stones.

Synthetic alpha-aluminum oxide (corundum) is used as starting material for high-grade technical ceramics. During processing additives are supplied to the material or impurities introduced unintentionally. The material properties -including material colour - are either intentionally modified by introducing additives or changed unintentionally due to impurities.

Tbl. 2 gives an example of the influence of various metallic oxides on the colour of an aluminum oxide ceramic material [12-15].

Experience has shown that MgO-doped aluminum oxide ceramies constitute the only bioinert material which is suited for implant application and is in compliance with ISO 6474. The current, advanced state-of-the-art is indicated in the last column of tbl. 1. MgO-doped materials generally feature a slight yellow colour after they have been sintered in air or oxygen [12-15]. This effect is exactly reproducible and does not affect the useful properties of the material.

252

EXPLANATION

The eolour effeet ean be explained as foliows:

under appropriate eonditions (partial oxygen pressure high enough, impurity potential low enough) the bivalent magnesium Mg2+ is integrated into the lattiee sites of the trivalent aluminum ion A13+, also refer to fig. 1 and 2. This provides for an unoceupied valenee whieh is able to interaet with eleetromagnetie radiation (light, UV, X- An~ -Gamma-radiation), i.e. a selective absorption takes plaee, whieh expresses itself by colouration. (For details refer to [12-15] and the referenees indicated there).

INFLUENCE ON USEFUJ, PROPERTIES

The eorrelation between the meehaniea1 strength G" and the median grain sizp d is

1 -1J:' (1)

This means that the mechaniea1 strength of materials featuring a small (large) median grain size is relatively high (low). Upon addition of MgO, fine grained structures, i.e. high mechanical strength, can be ensured.

In the absence nf sili~ates, alkaline oxides and alkaline earth oxides there i8 no vitreous phase formed on the grain bounda~ips if MgO i8 added. This is parti~ularly significant ~inee silicates are susceptible to disintegration in the "'''man hody.

Basically, MgO-dnped aluminum oxide ~eramics are suited for implant application if the density, grain size and mechanical strength offerpd by them is in compliance with the applicable standard.

Only MgO-doped materials free of silicates show a slight yellow or hrown disco10llrat.ion folJowing t.n Gamma-sterilizat.ion, i .e. t.he rolour effect indieates the 1).1\1'11 i ty of thp material and its comp1 iance wi th ISO 6474.

253

TABLE 1

Figures indicated in ISO 6474 for aluminum oxide ceramic implants.

Properties Units

Bulk density g/cm 3

Al203 contents weight " Si02 + alkali oxides weight " Si02 weight " Na20 weight " Si02+Na20 + CaO+ Fe203 weight " CaO weight " Fe203 weight " Grain size um Bending strength MPa Corrosion mg/m 2

ISO 6474

> 3.90 ~99.5

< 0.1

~ 7 ~400

~ 0.1

under BIOLOX* discussion

2- 3.94 3.95 99.7

< 0.01 < 0.01 < 0.01 < 0.01

< 0.1 < 0.01 < 0.015

< 4.5 4.3 L 450 550

< 0.1 ------------------------------------------------------------*BIOLOX is a product of Feldmühle

TABLE 2

Colour of aluminum oxide

Additive

no additive

Silicates (Si02)

Cr203 (0.5 - 5 ")

NiO

MgO

NiO + MgO

CoO

colour after sintering

colourless, white

white

pink to depp-red

green

ypllow (brown after radiation treatment)

gold-coloured

grey to dpep-hlar~

bItte

254

( gO)

A13+

0 2-

Figure 1: Unoccupied valences cf the aluminium, magnesium and oxygen ions

Figure 2: Bivalent Mg-ion in the lattice site of a trivalent aluminum ion. Resulting from this is an unoccupied valence of an oxygene ion (diagrammatic view).

255

REFERENCES

1 . Dörre, E., Erfl\hrungen mit dem Implantatwerkstoff Aluminiumoxid­Keramik - Der Zahnarzt 29 (1~8n) 267

2. Bläsil1S, K. and Schnf'idf'r, E., Endoprothesen-Atlas Hüftf' - Georg Thieme Verlag (1989)

3. Dörre, E. et al. Behaltf'n Kf'ramikkompnnenten künstlichf'r Hilftge]enkf' ihre Ff'stigkeit im mf'nschlichen Körper? Biomed. Technik 27 (19821 303

4. Semlitsch, M., 20 Jahre Sulzer-ErfRhrung mit Werkstoffen filr künstliche Hüftgelenkf' - Werkstoffe & Konstruktionen 3 (1989) p. 216

5. Morf' 11, R., Handbook of Properties of Technical & Engineering Ceramies Part 2: Data Review Rpetlon 1: High-Alumina Ceramies - London: Her Majesty's Rtat.ionary Office (1987)

6. rRO 6474 - 1981 Implants for surgery - Ceramic materials based on alumina - 1st ed. 1981-02-01

7. Willmann, G., Die Bedeut~ng der ISO-Norm 6474 für Implantate aus Aluminiumoxid - Zahnärztl. Praxis 41 (1990) 286-290

8. Dörre, E., Hübner, H., Alumina - Springer Verlag (1984)

9. Petzold, A., Ulbricht, J., Tonerde und Tonerdewerkstoffe - VEB Dt. Vf'rlag f. Grundstoffindustrie (1984)

10. Hausner, H., Oxidkeramik - in [11] pp. 166 - 203

11. Salmang, H., Scholze (ed.) Keramik Teil - Springer Verlag (1983)

12. Endl, H., Hausner, H., Einfluß von Spurenverunreinigungen und Sinteratmosphäre auf optische und elektrische Eigenschaften von Aluminium­oxid - Ber. Dt. Keram. Ges. 57 (1980) p. 121

13. Endl, H:, Hausner, H., Optical Properties of Pure and Doped Aluminium Oxide - Science of Ceramies 10 (1980) p. 697

14. Blum, J. B. et. al. Temperature Dependence of the Iron Acceptor Level in Aluminium Oxide - J. Amer. Ceram. Soc. 65 (1982) p. 379

15 .. Genthe, W., Hausner, H., Die Farbe von Sinterkörpern aus Aluminiumoxid - efi/Ber. Dt. Keram. Ges. 64 (1987) p. 292

256

DENTAL CERAMICS AND COMPOSITE RESINS AS RESTORATIVE MATERIALS

CARLO PRATI, EUGENIO TOSCHI, CESARE NUCCI, ROMANO MONGIORGI*, ANTONIO SA VINO.

School of dentistry, Department of operative dentistry, *Department of Mineralogical Sciences, University of Bologna, ltaly.

Via San Vitale 59, 40125 Bologna, ltaly

ABSTRACT

The aim of this study was to evaluate the marginal microleakage of different materials used to restore dass 11 MOD cavities. Materials selected were: conventional BIS­GMNurethane composite resin (Heliomolar RO), BIS-GMNceramic APC composite resin (P 50), ceramic (IVOCLAR-VIVODENT) and urethane resin (Isosit SR) (N=10). A bonding agent was used to improve the adhesion between materials and tooth structures. Materials were prepared according with manufacturer's directions. After the restorations, each sampie was immersed in dye solution and after 12 hours horizontally sectioned and ispected under optical stereomicroscope to evaluate the dye penetration along the interface tooth-material. ANOVA and Tukey tests were used for the statistical analysis. All materials showed microleakage along the restoration margins. Several materials showed more dye penetration than other, suggesting that early stress, gap, adhesive and cohesive failures influence the interface between material and tooth.

INTRODUCTION

Ceramic bonded composite materials (P 30 and P 50) and composite INLAY systems (Brilliant D.I., Isosit, EOS) have been recently developed to increase the wear resistance of posterior restorations. Althougth unclaired properties have improved, microleakage remains a clinical problem. The aim of the present study was to evaluate early microleakage at three different levels from using different filIing materials.

257

MATERIALS AND METHODS

The materials used in the study are Iisted in Table 1a and 1b. Freshly extracted third molars of healthy young patients stored in saline solution at room temperature for no more than one month were used in the study. Teeth were cleaned mechanically with a curette and wetted with apomice, air dried and coated with a nylon varnish exeept for the mesio-occlusal-distaI surfaces. Teeth were coated with varnish before cavity preparation to reduce time after restoration linishing before imersion in dye solution. Eighty Class n MOD cavities were prepared with cervical margins linished Just below the CEJ. All preparations were made with high speed diamond burs (Intensiv, Swiss) with water coolant. Hand Instruments were used to finish the approximal boxes to a cavo­surfaee margin at 9OD•

All Bonding Agents were prepared according to manufacturer's directions and applied on cavity surfaces with a small brush. All the materials were used and applied in cavity following the manufacturer's directions. GICs were used only with P 30 and P 50. Three to four minutes after polymerization, the restoration were finished with discs (3M Sof-Lex Pop -On,3M Dental Products Co., USA) of medium, fine and superfine grits. The teeth were immediatly immersed in 2% erythrosin B solution for 12 hours at 37DC and then transversally sectioned with a diamond-bladed circular saw at three levels. The section levels were made at 0.5 mm (gingival section) and 2 ± 0.2 mm (intermediate section). Microleakage score was evaluated along the dentine-restoration interface under a stereomicroscope and calculated using alinear graduated scale, similar to that described by Hammesfarh et al.(1987) (1). The graduated scales were for occlusal and intermediate sections and for gingival section. The microleakage value was then expressed as the percentage of dye against the total of the restoration .Mesial and distal microleakage and lingual and buccal microleakage scores were taken separately. The leakage data was evaluated by !wo examiners and compared until an agreement was achieved. Statistical analysis was performed with the Kruskal-Wallis one-way analysis of variance and the Mann Witney U test.

TABLE1a Dentinal Bonding Agents used in the study

Materials

Scotchbond Dual Cure # Scotchbond 2 # Vitrabond * Dual Cement & Duo Cement&

#) Dentinal Bonding Agent, *) Light-cured GIC,

Manufacturer

3M Dental Products, st. Paul, USA 3M Dental Products, St. Paul, USA 3M Dental Products, St. Paul, USA IVOCLAR, Leicthstein COLTENE, Switzerland

&) Light and chemical cured compositecement

Materials

P30 P50 Brilliant D.I. Isosit EOS Heliomolar R.O.

258

TABLEIb Composite resins tested in the study

ManuCacturer

3M Dental Products, St. Paul, USA Idem Coltene, Switzerland Ivoclar, Leicthstein Idem Idem

RESULTS

For convenience table 2 shows the combined mean of microleakage for the gingival and occlusal section of each group of material: - The CEJ sections presented the highest microleakage in aII the tested materials. -GIC + P50 and isosit showed the lowest microleakage incoming the best marginal seal.

Materails

P30+SBDC P30 + GIC P50+ SB2 Coltene-Brilliant EOS Isosit .p 50 + GIC

Leakage(%)

49.1 * 41.0 * 38.4 * 24.7 * 14.1 ** 12.7 ** 2.4 **

TABLE2 Results

S.D.

42.1 34.2 31.3 30.0 17.1 10.4 4.9

The materials with the same number of * are not statistically ditTerent.

DISCUSSION

These results confirm others intending to demonstrate that there is considerable marginal microleakage immediately following restoration, especially at CEJ level (2). During the composite setting reaction, the dentine-composite bond might break with a consequent marginal gap and microleakage (3). This is especially true for P 30 and P 50 used without GIC.

259

The present data agrees with other studies that have demonstrated considerable microleakage at CEJ in Class n restorations (4, 5, 6) and in Class V restoration (2 , 7).On the other band, all the materials tested showed a large number of gap-free restorations at the occlusal sction,suggesting that a good enamel-composite bond and seal is initially possible after restoration. This would imply that adequate enamel thickness could Is obtainnable in all composite restorations. It has been demonstrated that inlay technique showed less marginal microleakage than hybrid composite. It is probable that the better marginal integrity of the inlays, compared to direct restorations, is due to free polymerization shrinkage. The present study suggests that CEJ sections especially are highly sensitive to restoration technique. This is particularly evident in composite restorations with the highest marginal microleakage, such as P 30 and P 50 restorations. The present study suggest s that a suitable layering restorative technique may prove necessary to reduce or eliminate early initial marginal microleakage in large Class n restorations.

REFERENCES

1. Hammersfahr, P.D., Huang, C.T. and ShatTer, M.M., Microleakage and bond strength ofresin restorations with various bonding agents. Dent Mater., 1987,3,194-199.

2. Prati, C. and Montanari, G., Comparative microleakage study between the sandwich and conventional three-increment techniques. Quintessence Int., 1989, 20, 587-594.

3. Kemp-Scholte, C.M. and Davidson, LC., Marginal sealing of curing contraction gaps in Class V composite resin restorations • .I. DENT. RES., 1988, 13, 12-19.

4. Eakle, W.S., EtTect of thermal cycling on fracture strength and microleakage in teeth restored with a bonded composite resin. Dmt.~, 1986,2,114-117.

5. Darbyshire, P .A., Messer, L.B. and Douglas, W.H., Microleakage in Class n composite restorations bonded to dentine using thermal and load cycling . .I. Dent. Res., 1988, 67, 585-587.

6. Prati, C.,Early marginal microleakage in class n resin composite restorations Dent Mater, 1989, 5, 392-398.

7. Gordon, M., Plasschaert, a.j.m. and Stark, M.M., Microleakage ofseveral tooth -colored restorative materials in cervical cavities. A comparative study in vitro. Dent Mater., 1986,2,228-231.

260

REINFORCED SILVER GLASS-IONOMER CEMENT AND LIGBT-CURED B.E.M.A. GLASS-IONOMER CEMENT UNDER SILVER-AMALGAM

RESTORATIONS: A MICROLEAKAGE STUDY

CESARE NUCCI, EUGENIO TOSCBI, ROMANO MONGIORGI*, CARLO PRATI School of Dentistry, Department of Operative Dentistry

*Department of Mineralogical Sciences, University of Bologna, Italy, Piazza Porta San Donato 1, 40126 Bologna,Italy

ABSTRACT

The lack of adhesion to dentine of amalgam restorations is responsable for insutncient marginal seal and for percolation of bacteria and chemical-hydrolitic solutions along the interface. Several silver-glass and B.E.M.A.-glass bases have been proposed to obtain gap-free amalgam restorations. The aim of the present study was to evaluate the microleakage and the marginal adaptation of two different glass-ionomer bases placed under amalgam restorations. A silver-reinforced chemical cured glass-ionomer cement (Ketac Silver) and a B.E.M.A. Iight-cured glass-ionomer cement (Vitrabond) were used as bases in the study. After microleakage tests, a complete marginal seal, without microleakage was observed only in B.E.M.A.-glass-ionomer/amalgam restorations.

INTRODUCTION

Glass-ionomer cements (GIC) have received much attention due to several clinical advantages such as dentine and enamel adhesion, biocompatibility, fluoride release and radio-opacity (1,2). These properties suggest that GICs could be used under amalgam (a combination of silver, mercury, tin and copper), the most common restorative material, to reduce early microleakage and marginal gap. Bowever, at present, no clinical and in vitro studies have evaluated the GICs as bases for amalgam restorations, because early generations of GICs did not show sutncient mechanical properties (1-3) To improve the mechanical properties of GICs, inorganic and organic components have been included: ceramic-coated silver particles (4), aluminum oxide (5), various metal salts (6) and B.E.M.A. (7) The aim of the present study was to evaluate the early marginal microleakage of amalgam restorations used with silver-reinforced and B.E.M.A. GICs.

261

MATERIALS AND METHODS

Extracted human third molars stored in saline solution at room temperature were used. Two GICs were selected ror the study. The experimental materials are presented in table 1. Third Class 11 MOD (mesio-occluso-distal) restorations were prepared with a diamond bur used with a high speed hand piece. Ten sampies were prepared ror each group. After cavity preparation, GICs were mixed according to their manuracturer's directions and then applied only on the cavity noors. Vitrabond was then photo-cured using a Iight-curing unit (Visilux 11, 3M, USA). In the control group, no GIC was placed on dentine surface. After 5 minutes, high-copper amalgam (Dispersalloy, Johnson & Johnson, USA) was mixed, applied on cavities and condensed with a band instrument. After 5 minutes, a11 amalgam margins were tinished with hand carver. The restored teeth were coated with nail varnish, exeept ror the area around the restoration, and placed in a saline solution ror 24 hours at room temperature and then placed in 2% erythrosine basic solution ror 48 hours. Each tooth was then removed, washed with water and sectioned transversally through the centre of the restoration with a diamond disko Two horizontal sections were prepared ror each sampie. Sections were evaluated under stereomicroscope at x60 magnitication by one examiner. A microleakage score was calculated along the dentinal and enamel walls. Dye penetration a10ng the dentinal wall was considered as the length or microleakage a10ng the interface between the dentine/enamel and amalgam. The length or microleakage was expressed as the percentage or dye penetration with respect to the total length or the cavity walls. Student's t test was used to determine any signiticant differences between pairs or experimental groups. A level oe signiticance or .05 was used throughout all statistical tests.

TABLEI

Glass-ionomer cements used in the study

Material (manuracturer)

Vitrabond (3M, MN, USA)

Ketac Silver (Espe, Germany)

Compound

Fluoroaluminosilicate glass powder H.E.M.A. and methacrylate resin

Fluoroaluminosilicate g1ass powder Sintered silver powder

RESULTS

After microleakage tests, aII the control sampies showed microleakage at the two sections evaluated. Ninety % of Ketac silver and 45% or Vitrabond sections presented microleakage. Table 2 illustrates the mean microleakage values or length or microleakage ror the three groups. Vitrabond showed the significantly lower microleakage value (p<0.01). Figures 1 and 2 drown the typical aspects oe a Ketac silver-amalgam and Vitrabond amalgam restoration.

262

Figure 1. The picture shows the gingival section of Ketac Silver-amalgam restoration.

Figure 2. The pieture illustrates the Vitrabond-amalgam restoration

263

TABLE2

Mean microleakage expressed as percentage of microleakage with respect of total length ofrestoration wall. (mean, SD, N= 10)

Material Microleakage SD Range

Control 100.0 (0.0) 0.0 100.0-100.0

Ketac Silver 68.4 18.4 12.1-100.0

Vitrabond 19.7 17.8 0.0-59.7

ANOvA test p <0.01

DISCUSSION

The most important problem of amalgam restorations is the open-space (i.e. marginal gap) that is responsible for microleakage and its clinical consequences. The present study demonstrated that it is possible to reduce marginal microleakage using a GIC able to bond the dentine surface (8). Our results confirm the marginal seal of Vitrabond, previously tested with composite resins (8) and the low marginal seal (higher microleakage values) of Ketac Silver (9). The low adhesion of ketac Silver and the lack of adhesion with dentine of amalgam are probably responsible for the great microleakage with respect to Vitrabond restorations. The clinical implications of our studies are that Vitrabond may be a biocompatible cement-base under amalgam restorations. This finding suggests that a mixed restoration (amalgam and GIC) may unify the advantages of both materials: dentine adhesion, fluoride release for Vitrabond and mechanical properties and wear resistance for amalgam. Further clinical studies are necessary to evaluate the long-life performance of GIC-amalgam restoration.

REFERENCES

1. Atkinson, A.S. and Pearson G.J, The evolution of glass-ionomer cements. BI Dmt I., 1985,153,335-337.

2. Smith D.C. Composition and characteristics of glass-ionomer cements, I.A.D.A., 1990, 120, 20-22.

3. Mathis, R.S., and Ferracane, J.L, Properties of glass-ionomer/resin-composite hybrid material, Dmt Mate.r. 1989, 5, 355-358.

4. McLean, J.W. and Gasser D. Glass-cermet cement. Quint Int 1985, 5, 333-343.

5. Oilo, G. and Ruyter, I.E., The influence ofvarious admixes on the phisical properties of a polycarboxylate cement, I Dmt Rci 1983, 62, 937-939.

264

6. Crisp, S., Merson, S.A. amd Wilson, A.D., Modification of ionomer eement by the addition of simple metal salts, lnd ED& CIwn &:Dd Ba. Dfi 1980, 19, 403-408.

7. 3M Company, St. Paul, MN, USA.

8. Prati, C., Nucci, c., Davidson, C.L. and Montanari, G., Early marginal microleakage and shear bond strength of adhesive restorative systems, Dmt MaW: 1990, 5, 195-200.

9. Thorton, J.B., Retief, D. H. and Bradley, E.L., Marginalleakage of two glass -ionomer eements: Ketac Fil and Ketac Silver, Am. J Dmt 1988, 1,35-38.

265

TENSILE BOND STRENGTH OF DENTAL PORCELAIN TO DENTAL COMPOSITE RESINS.

ROMANO MONGIORGI*, CARLO PRATI**, EUGENIO TOSCm**, GIOV ANNA BERTOCCm***

*Department of Mineralogical Sciences, **School of Dentistry, Department of OperativeDentistry and *** Department of Pharmacological Sciences,

University of Bologna, ltaly, Piazza Porta San Donato 1, 40126 Bologna, ltaly

ABSTRACT

The aim of the present study was to evaluate the in vitro tensile bond strength of specimens of porcelain bonded to dual curing dental composite using two organo­silane materials: Scothcprime Ceramic Primer (3M, Co., St. Paul, Mn, USA) and Silanit (Vivadent, Schaan, Liecthstein). Bond stremgth was between 9.2 and 10.4 MPa. The failure was observed only in composite sampies. The study suggest that silane resin materials have adhesive bond higher than cohesive value of composite resins.

INTRODUCTION

Silane coupling agents have been widely used in plastic technology since the '40s, but their use in dental materials was introduced in 1963 (1) with a composite resin filled with silane-treated silica powder. Silane coupling agents have been used in dentistry for bonding porcelain teeth to acrylic denture bases (2,3), bonding hydroxyapatite to composite (4), bonding porcelain inlays to etched enamel using composite as luting agent (5), reparing fractured porcelain welded to metal restorations (6,7,8,9, 10, 11), bonding orthodontic brackets to enamel (12), bonding porcelain laminate veneers or inlay/onlay restorations to etched enamel using composite as luting agent (13 ,14, 15, 16). Several studies have demonstrated that silane treatement is highly sensitive to a number of factors which may affect the reability of the bond: hydrolisis, polymerization, shelf life of products, possibility of combination, exposure to wet enviroment and the quality commercial products (5, 9,13,14,17,18). All these factors may be determinal to the effectiveness of the treatement. Many in vitro studies have been made on tensile bond strength, shear bond strength, effect of moisture and/or thermocycling on bond strength, and fracture resistance of teeth restored by these means but the results vary greatly (19,20, 21, 22, 6, 23, 24, 25, 11,26).

266

Two different materials have been recently marketed and are now available for routine dentistry: Scotchprime Ceramic Primer (3M) and Silanit (Vivadent). The purpose of teh present study was to investigate the in vitro tensile bond strength of specimens of porcelain bonded to dual curing dental composite using these two organosilane materials.

MATERIALS AND METHODS

Thirty porcelain cylinders, 12 mm " and 10 mm h, were hand-condensed using Vita Porcelain mixed with distilied water (p/w = 1.5). These specimens were fired at 950°C on quartz fiber refractory material trays. Each specimen was undercut and embedded in a cold curing resin block, 25 mm w, 25 mm d and 15 mm h, so that only the upper surface of each cylinder was exposed. The porcelain surface were rermed wiyh a coarse diamond in a high speed handpiece. Twentytive specimens were etched with Austenal Etching Solution in an ultrasonic bath for 3.5 min and then washed with residue-free compressed air. The five unetched specimens were assigned. to Group A. They were silane treated with Scotchprime Ceramic Primer (3M) and a thin layer of Scotchbond LC Adhesive (3M) bonding agent was light-cured on each porcelain surface according to the manufacturer's instructions. A plug of Micropont (Kulzer) composite resin, 6 mm " and 2 mm h, was formed in an undercut removable teflon mold and ligth -cured for 60 sec on each porcelain surface. The mold was then removed. Five etched specimens underwent the same procedure except silane treatment. They were assigned to group B. The remaining twenty specimens were divided into two groups. Ten etched specimens were silane treated with Scotchprime Ceramic Primer (3M) and bonded to Scotchbond LC Adhesive (3M) and to Micropont (Kulzer) as previously done for group A. They were assigned to group C. Ten etched specimens were silane treated with Silanit (Vivadent) and a thin layer of Heliobond (Vivadent) was ligth cured on each porcelain surface according to the manufacturer's instructions. A Dual Cement (Vivadent) composite resin plug was formed in a removable teflon mold and ligth-cured for 60 sec on each specimen. They were assigned to group D. Sinse Silanit (Vivadent) is not a one-part system, like Scotchprime Ceramic Primer (3M), but a two-part system, the two solutions A and B were stirred in a mixing weil and allowed to hydrolize for 45 min. Both Micropont (Kulzer) and Dual Cement (Vivadent) are self/light-curing dental composites recommended for porcelain laminate veneers and inlay/onlay restoratin cementation. All these specimens were shaped to fit into a specially designed text rlXture clamped into the jaws of an Instron-type machine. Tensile bond strength test was made after a week at 20°C in air, crosshead speed set at 2 mm/min and 20 kilos full scale on chart paper until failure occurred.

RESULTS

The means of the tensile bond strength and the standard deviation values are shown in Table 1. Examination under optical stereomicroscope of the fractured interfaces showed that cohesive failure occurred in the composite bulk in all of specimens. The porcelain surface itself was never fractured. The fracture occurred at the same time for the specimens of Groups A, Band C, whereas the specimens of Group D fractured partialy in the composite before final failure. The tensile bond strength means values of Group D were higher than control Groups A, Band C mean values.

TBS

GROUPA GROUPB GROUPC GROUPD

mean values (MPa)

10.91 10.65 11.41 13.07

Student t test; p >0.05

267

TABLEI

Standard deviation (MPa)

1.75 1.58 1.73 1.54

DISCUSSION

Numberof specimens

5 5 10 10

Most studies have shown a wide range of bond strength values probably because of the sensitivity of the technique. Besides the ditTerent dental porcelains and composites used, the ditTerent water immersion times and thermocycling techniques must be considered when comparisons are attempted. 3M (22) reported a 19.02 MPa (sd 2.84 MPa9 shear bond strength for Scotchprime Ceramic Primer (3M) used in combination with dental composite Silux (3M) and a 14.11 MPa (sd 3.84 MPa) shear bond strength for Silanit (Vivadent). 3M (22) claimed that higer bond strength is achievable using Scotchprime Ceramic Primer (3M), at which the bond strength to porcelain is higher than the cohesive strength of the porcelain itself (22). Vivadent (21) reported a 10.4 MPa (sd 2.7 MPa) shear bond strength for Scotchprime Ceramic Primer (3M) and a 8.2 MPa (sd 1.4 MPa) shear bond stregth for Silanit used in combination with teh dental adhesive Beliobond (Vivadent) and the dental composite Beliosit (Vivadent). The shear bond strength test was performed after 24 hours water storage of all specimens. Bowever it is worth noting that after a week of water storage, the shear bond strength recorded for Scotchprime Ceramic Primer (3M) amounted to 9.4 MPa (sd 1.6 MPa) and to 16.0 MPa (sd 0.8 MPa) for Silanit (Vivadent). Other Authors (25) reported a 39.7 MPa (sd 10.5 MPa) shear bond strength for silanit (Vivadent) after 24 hours water immersion of specimens but bond strength decreased to 30.6 MPa (sd 10.0 MPa) after live months water immesrsion. Another study (23) reported a 4.4 MPa (sd 0.7 MPa) tensile bond strength for Silanit (Vivadent) after two weeks water storage of the specimens. Other authors (20) reported a 23.54 MPa (sd 9.03 MPa) shear bond strength for Scotchprime Ceramic Primer (3M) for dry specimens and a 21.78 MPa (sd 11.58 MPa) shear bond strength after a week of storage of specimens. Another study (26) reported a 13.44 MPa for Scotchprime Ceramic Primer (3M) for both dry and wet specimens and observed cohesive failure occurring in the porcelain bulk. Other authors (24) repoted a 46.84 MPa (sd 6.55 MPa) shear bond strength for Scothprime Ceramic Primer used i combination with dental composite P-30 (3M) and observed cohesive failure of porcelain. The present study showed that the porcelain to composite tensile bond strength exceeded both the composite to bonding agent bond strength. This ditTerence determined the cohesive failure of the composite, evidently the weakest link in the chair. Thes results are lower than expected and the composite failure also preventred the evalutation of porcelain to composite bond strength whici was expested to be higer. Therefore these results cannot be used to assess any ditTerence between the etTectiveness of 3M and Vivadent organosilane materials. Bowever, these results are considerably higher than the 7.5 MPa tensile bond strength of etched porcelain to composite, expected to prove adeguate in clinical conditions (17) and the Authors consider this result encouraging.

268

Our results seem to support the use of Scotchprime in clinical practice. Indeed, the two-steps clinical follow-up on 24 sampies (13 patients) showed no surface alterations or failures suggesting a good correlation between "in vitro" and "in vivo".

REFERENCES

1. Bowen, R.L, Properties of a silica reinforced polymer for dental restoration . .IA.D.A.,1963, 66, 57.

2. PatTenbarger, G.C., Sweeney, W.T. and Bowen, R.L, Bonding porcelain teeth to acrylic denture bases . .IA.DA" 1967, 74, 1018

3. Semmelmann, J.O. and Kulp, P.R., Silane bonding porcelain teeth to acrylic . .I.AJl.A., 1968, 76, 69-73.

4. Loebenstein, W.V. and Kumpula, W., A new method evaluating coupling agents bonding polymer to tooth mineral . .I. Dent.Res" 1977,56,1219-1226.

5. Rochette, A., Adhesion par polymeres et traitment de surface en odonto­stomatologie Actualites Odontostomatologiques, 1972, 172.

6. Bighton, R., Caputo, AA., Matyas, J. and Tashima, B., Incisalload transfer by porcelain laminates . .I, Prosthet Dent, 1986,-Ahstr.. 724.

7. Rogers, LB., Feiler, P. and Eames, W.B., Evalutation of a bonding agent for porcelain and gold repair . .I, Dent Res" 1977, Ahstr. 502.

8. Dent, R.J., Repair of porcelain fused to metal restorations, .I.Prosthet Dmt.,1979, 41, 661-664.

9. Newburg, R. and Pameijer, C.B., Composite resins bonded to porcelain with silane solution • .I,A,DA" 96, 288-291.

10. Barreto, Mt: and Bottaro, B.F., A pratical approach to porcelain repair . ...L Prosthet, Dent, 1982, 48, 349-351.

11. Nowlin, T.P., Barghi, N. and Norling, B., Evalutation ofthe bonding ofthree porcelain repair systems . .I, Prosthet Dent, 1981, 46, 516-518.

12. Johnson, R.G., A new method for direct bonding orthodontic attachment to porcelain teeth using a silane coupling agent: an in vitro evalutation. AmJ.. Ortbod" 1980, 78, 233.

13. Born, B.R., Porcelain laminate veneers bonded to atched enamel. Dent CHn, NorthAm" 1983, 27, 671-684.

14. McLaughlin, G., Porcelain fused to tooth-a new esthetic and recostructive modality., Comp, Cont Educ., 1984,5,430.

15. Born, B.R., A new lamination: porcelain bonded to enamel., New York State Dmml., 1983, 401-403.

16. Calamia, J.R., Etched porcelain veneers: the current state of art. Quint.

269

Internat., 1985, 1, 5-12.

17. Chen, T.m: and Brauer, G.M., Solvents etTects on bonding organosilanes to silica surfaces • .I. Deut. Res. 1982,61,1439-1443.

18. Culler, S.R., Krueger, D.D. and Joos, R. W., Investigations of Silane priming solutions to repair fractured porcelain crowns • ..J.. Deut. Res" 1986, Abstr., 193.

19. Bailey, J.B., Porcelain to composite resin bond strength using organosilane materials • .I. Dent. Res" 1987, Abstr., 804.

20. Cabalzar, D., Langzeithaftung mit silanit. Yivadent Documeutation 1986

21. Culler, S.R. and Krueger, D.D., Inveswtigations of silane priming solutions to repair fractured porcelain crowns. 3M Documentation 1986.

22. Bueber, B.P., Untersuchung von Baftermittlern zur reparatur von Metall -keramik. Dtseb, Zabnartztl, Z., 1983,38,26-27.

23. Lacy, A., LaLuz, J., Watanabe, L. and Dellinges, M., EtTects ofporcelain surface treatment on the bond to composite resin.J.Dent. Res., 1987, Abstr., 1108.

24. Notter, O.R., Belser, V.C., Porzellanreparatur materialen: experimentelle Untersuchungen der Baftigkeit verschiedener Produkte. Schweiz. Mscbz. ZabnbeUk" 1982, Abstr. 805.

25. O'Kray, K., Suchak, A.J. and Stanford, J.W., Shear strength of porcelain repair materials • .I. Deut. Res" 1987, Abstr., 805.

26. Calamia, J.R. and Simonsen, R.J., Tensile bond strength of etched porcelain • ..J.. Dent. Res" 1983, Abstr. 1154.

270

GLASS-IONOMER CEMENTS AS BASE FOR COMPOSITE RESTORATIONS

EUGENIO TOSCm, ROMANO MONGIORGI*, CARLO PRATI, GIOVANNI VALDRE'*, CESARE NUCCL

School of Dentistry, Department of Operative Dentistry, *Department of Mineralogical Sciences, University of Bologna, Italy,

Piazza di Porta San Donato 1, 40126 Bologna, Italy

ABSTRACT

Glass-ionomer cements (GIC) are able to react with dentin and enamel tooth surface reducing the marginal gap along the restorations. The used of GIC in association with composite materials has been called "sandwich restoration" and has been proposed to improve the marginal seal and marginal adaptation of composite resin and bonding agents. The aims of this study were: 1) to evaluate the marginal adaptation of Class V restorations under SEM; 2) to compare the marginal stress, the cracks and fractures a10ng the interface of GIC and composites. Class V cavities were prepared and tilled according to manufacturer's directions. After 3 hours, the restorations were observed under SEM. Only GIC composite restorations showed a good marginal adaptation without gap, probably due to the elasticity and ftow of GICs.

INTRODUCTION

Glass-ionomer cements weredeveloped and introduced into clinical practice about 20 years ago. The first formulation of GIC consisted of fluoro-aluminosilicate powder mixed with (poly) acryllc acid liquid. The mayor advantages of fmt generation GIC were the capacity to bond dentine, enamel, composite surfaces and the ability to realase fluoride (1). The cariostatic action of GICs has been recently demonstrated (1), confirming the first preliminary studies. The major problems of first generation of GIC were the low compressive strength, the cracks and fractures in thickness of materials exposed to air and case of handling (2,3). The tirst application proposed for GIC was as filling material, and a clinical study demonstrated valid clinical results in class V restorations. The color shade and the fractures limited the use of GIC as direct filling material. In 1985 McLean and co-workers (4) proposed the use of GIC as a base under composite resin restoration. This has been called "sandwich technique"(Prati et al.1989) or "biomimetic technique". The first in vitro studies demonstrated great ditTerences among the various types of GICs and dentinal bonding systems (DBS)(S) and suggested that insufficient dentine adhesion was the problem of sandwich technique.

271

At present, great attention has been given to dentine chemical treatments to improve adhesion, and a new generation of GICs have been developed. Table 1 swohs the development of GICs.

TABLEI Glass-ionomer cements developed for clinical use

r GENERATION

ASPA

Ir GENERATION

Ketac Bond Ketac Silver Ketac Fil Durelon 3M Fuji land U GC Iining cement

The aims of the study were to evaluate: 1) the dentine adhesion of GICs; 2) The marginal gap of dass V restorations.

MATERIALS AND METHODS

The GICs used are schematized in Table 2.

Material

Ketac Silver GC Iining cement lonosit Zionomer Vitrabond

TABLE2 Materials used in the study

chemical-cured chemical-cured Iight-cured Iight-cured Iight-cured

Manufacturer

ESPE, Germany Fuji-GC, Japan DMG, Germany Den Mat, USA 3M, USA

DENTIN ADHESION TEST

ur GENERATION

Base Line lonosit Zionomer Vitrabond

Fifty human teeth used for the study were stored in saline solution for 1 month at 4°C. They were mounted in a uniform mold and abraded from vestibular enamel using a diamond bur to obtain a flat dentine surface. GICs were mixed according to manufacturer's directions and applied on a teflon tube (4.0 mm internal diameter, 5.0 mm height and positioned on flat dentine surface. After 24 hours, the sampies were tested. Statistical analyses were carried out with ANOVA TEST and Student t test. SEM TEST

272

Twenty class V butt-joint restorations were prepared in human extracted teeth. No bevel was made at enamellevel. The cavities were prepared at enamel level (N =8), at dentine level (N=8) and at cementum-enamel junction (N=4). Two groups of restorations were prepared: one group was prepared using GIC as base, one group was prepared without GIC (control group). In the first group Vitrabond was applied as base on the noor of the cavities, and photocured for 30 sec. A layer of Scotchbond 2 (3M, Co, MN, USA) was than applied on an GICs and enamel surfaces and photocured. Finally, one increment of composite resins Silux (3M, Co, MN, USA) was inserted in the cavities and photocured.In the second group Scotchbond 2 was directly applied on the dentine and then composite resin. After finishing the composite surface, each sampie was prepared for SEM analysis. A Jeol (Japan) was used to evaluate the marginal pp around the restoration.

RESULTS

Table 3 illustrates the mean shear bond strength values and the standard deviation of the materials tested in the present study. SEM analysis demonstrated no marginal gap around the Vitrabond restorations. An open-space of 2-14 microns was observed along the dentine margin in the other group of restorations. The figures 1 and 2 shown the SEM aspect of one restoration with Vitrabond + Silux and with Scotchbond + Silux.

Ketac Silver GC Lining cement Ionosit Vitrabond Zionomer

TABLE3 Shear bond strength values (MPa ± SD; N=10)

1.4 3.8 ,. 3.9 ,. 6.8 ** 7.2 ,.,.

1.2 2.7 3.1 2.5 2.4

'I'he matenaIs wdh fhe same number 01 * are not stäbsbcally dlßerenL

DISCUSSION

The shrinkage of composite resins is the most important problem responsible for microleakage. To eliminate this inconvenience, cement bases have been used, although marginal gap remains a clinical problem. The low adhesion of the first generation of GICs was considered the major limit to their clinical use. To avoid gap formation between restorative and tooth, and thus produce a leak-proof seal, a permanent bond between tooth sustrate and GICs is necessary (6). In the present study, the newest generation of GICs (Vitrabond and Zionomer) gave the highest bond strength values. Several investigators have determined the performances of conventional and newest GICs. Mc Caghren (6) determined the shear bond strength ofVitrabond and reported bond strength of 9.3 MN.m-2 to dentin (the specimens were stored in distilied water at 37°C for 24h prior to testing).Similar results were described by Prati et al.(1990) (7-8).

273

Figure 1. Excellent marginal adaptation of Vitra bond restoration. No marginal gap is present along the interface between restoration and dentine.

Figure 2. Marginal adaptation of Scotch bond 2-Silux. A gap of 1-12 microns is visible around the restoration.

274

With the development of GICs, the problem of marginal leakage has been lessened. However, the results of previous stadies (5-7) indicate that the problem has not been completely solved. SEM analysis of Vitrabond restorations demonstrated a dosed marginal space, suggesting that dass V sandwich restorations may ensure a clinical marginal seal. A clinical stady (9) showed that the newly formulated GICs set rapidly and this is thought to be an advantage by general practitioners. In a recent stady Suzuki and Jordan (10) suggest that the combined ionomer-composite restoration provides a reUable chemical bond to dentin, micromechanical bonding of the composite to ionomer surface, and a highly acceptable esthetic result. On the other hand, Mount (11) noted that in the "sandwich technique", when using certain composite resins and their prescribed bonding resins, the failure under tensile stress was adhesive at the union rather than cohesive in the cement. Nonetheless, the results confirmed that the high adhesion of latest generation GICs means they may be used in clinical practke.

REFERENCES

1. Atkinson, A.S. and Pearson, GJ., The evolution of g1ass-ionomer cements.Jk DmU 1985, 159,335-337.

2. Oilo, G., Characterization of glass ionomer filling materials. Dent Mater, 1988, 4,129-133.

3. Oilo, G. and Evje, D.M., Abend test for measuring cement-dentin bond.Jknt ~,1988,4,98-102.

4. McLean, J.W., Prosser, HJ. and Wilson, A.D., The use of glass-ionomer cements in bonding composite resins to dentine. Br Dent .I, 1985, 158, 410-414.

5. Prati, C. and Montanari, G., Comparative microleakage study between the sandwich and conventionaI three-increment techniques. Quintessence Int., 1989,20,587-594.

6. McCagharen, RA, Retief, D.H., Bradley, E.L. and Denys, F.R., Shear bond strength of light-cured g1ass ionomer to enamel and dentin • .I Dent Res, 1990, 69, 40-45.

7. Prati, C., Nucci, C, Davidson, C.L. and Montanari, G., Early marginal microleakage and shear bond strength of adhesive restorative systems. Dent Mater, 1990, 5j 195-200.

8. Prati, C., Nucci, C. and Montanari, G., EtTect of acid and cleaning agents on shear bond strength and marginal microleakage of glass ionomer cement. Dent Mater, 1989, 5, 260-265.

9. Knibbs, P J. and Plant, C.G., An evaluation of a rapid setting glass ionomer cement in general dental practice. Australian Dent .I, 1989, 34, 459-465.

10 Suzuki, M. and Jordan, R.E., Glass ionomer-composite sandwich technique . .IADA. 1990, 120, 55-57.

11. Mount, GJ., The wettability of bonding resins used in the composite resin/glass ionomer "sandwich techoique". Australian Dent .1,1989,34,32-35.

275

SYSTEMIC CONTROL OF TISSUE AND CELL REACTIONS RELATING TO CERAMIC IMPLANTS

U.GROSS,C.MÜLLER-MAI*,C.VOIGT** Institute of patholo~ ,Department of Traumatology and

Reconstructive Surgery Klinikum Steglitz and Institute of Anatomy* of the Free University of Berlin,Hindenburgdamm 30,

D 1000 Berlin 45

ABSTRACT

On the basis of a number of animal experiments wi th implantation of metals,ceramics,composits, mainly into the distal femur of rabbi ts, some insight could be collected on the tissue response,the materials response and the biomechanical quality of the implantjtissue bond. Some principles for the anchoring of implants to bone were demonstrated. Additionally, surgically and by post-mortem examination retrieved implants and the surrounding tissue were analysed for structure and quality.Furthermore, in a limited number of cases the metal content in the soft tissue layer was chemically determined. In same ca ses there was a rather high content of metals in the soft tissue layer • The high metal concentration seems to be not compatible wi th normal live of cells occurring in the imflammatory and reparatory or regenerative product.

INTRODUCTION

After the implantation of surface reactive and inert materials there are material and host responses which are interrelated (1) • These responses are operative at various distances from the surfaces of implants (2,3).

Some reactions are strictly localized at the interface and related to changes of the implant (leaching, degradation, corrosion) (4) , and to changes in the tissue of the host (interfacial reactions or responses).The leading mechanisms are related to wound healing and complicated by the adsorption

276

and desorption of substances at the implant surface. Beside~, only the sequence of the main cellular events are known. The leading phenomena are macrophage and foreign body giant cell occurrence and development of organization and granulation tissue, and the regeneration of cells that are typical for the implantation site (1).

Other reactions are observable in the neighbourhood of an implant and related to the dissipation of soluble or particulate material coming from the implant, and, furthermore, to the adaptation of the host tissues to the change of the various conditions, e.g. biomechanical conditions,change in blood supply, in innervation and some other parameters that have not yet been determined sufficiently.The best known phenomenon of this kind is stress shielding and consecuti ve bone loss in the case of skeletal implants.The leading mechanisms of this neighbourhood response are known in general pathology as adap~ation. stress shielding induced by implants seems to be not principally different from atrophy of bone induced by inactivity, e.g.by a cast.

Other responses or reactions are even more remote and correlated wi th the release of soluble ( 4 ) and particulate substances from the implanted materials affecting remote organs via propagation by blood ,lymph, or intracavital dissipation. There are indications for preferential affection of organs due to their role in metabolism for resorption, turnover and excretion of substances(2,3) .These processes are addressed in the current literature as systemic effects after implantation of biomaterials.

In following the suggested ar~lmentation it can be stated various systems are invo1ved in the contro1 of tissue and reactions after the implantation of devices for replacement of structure or function of a diseased organ.

that cell the

The objectives of this study are:i.to give a survey on tissue structures at the. interface of metallic and ceramic implant cylinders with different surface roughnesses in relation to the distance of the implant when inserted in trabecular bone of the distal rabbit femur,and i1. to compare the results of the animal studies with the results of some retrieved specimens from patients with clinically successful and failed implants.

MATERIALS AHD METHODS

This consideration is based on a number of experiments that have been executed over the years wi th a variety of implant materials and designs .The materials were glasses,glass­ceramics,hydroxyapatite in different kinds of preparation ( 1 ) , calcium carbonate as natural coral , compesi tes made from metal and ceramics (5,6) or polymer matrix being carbon-fiber

277

reinforced.Different animal models have been studied and the results were compared with those gained after retrieval analysis of patient specimens removed either surgically or at post-mortem examinations.

The methods used in this investigation were histology of conventional or undecalcified sawed sections,scanning (SEM) and transmission electron microscopy (TEM)(6,7), computer tomography, morphometry,biomechanical testing for tensile strength,chemical analysis of the metal conte nt in tissue (Fe,Cr,Ni,co,Ti,Al,V),and statistical methods.The more detailed documentation of these experiments has been given to some extent in prior publications (1,5-7).

RESULTS

After 84 and 168 days postoperatively cylinders of titanium with a roughness of 1,20 and 50 ",m,and cylinders of Ti6Al4V with a roughness of 0.5 ",m were not bonded to bone.They displayed near the interface a bone frame which was integrated in the structure of the trabeculae of the distal rabbit femur.From the bone frame foot-like projections of bone developed to come into contact with the implant surface (Fig.1).

Figure 1.Cylinder of titanium with a surface roughness of 1 ",m 168 days after implantation in the distal femur epiphysis of an adult female rabbit.Sawed section.Von Kossa/fuchsin staining.Bar 1 mm.

278

Figure 2.Cylinder of titanium, surface roughness 100 #Am,168 days after implantation.Anchoring of the bone trabeculae without a bone ring as in Fig.1.Characteristic enlargement of the trabeculae near the interface.Sawed section.Von Kossa/fuchsin staining.Bar 1 mm.

In contrast,titanium cylinders with rough surface (RT 100 #Am) and cylinders made from surface-reactive hot isostatic pressed (HIP) hydroxyapatite (HA), roughness Rz 0.5, 20, or 50 #Am, and flame sprayed HA ,roughness Rt 50 #Am, were anchored to the bone trabeculae without this ring-like bone frame. The anchoring of the bone trabeculae was with continuity of mineralized extracellular matrix to the surface irregularities of the implant materials which were found in the TEM.The diameter of the trabeculae was increasing towards the implant surface,and the trabeculae displayed a surface profile corresponding to Baud's curves (8) (Fig.2,3). These bone­bonding implant cylinders could be loaded for tension with 1-4 N/mm2 • In contrast, the non-bonding implant cylinders did not reveal tensile strength at the interface.

Figure 3.Anchoring of bone trabeculae in the HA flame spray coating of a Ti6A14V cylinder 84 days after implantation. Enlargement of the trabeculae towards the interface.Shallow depression in the HA coating due to resorption.Sawed section.Von Kossa/fuchsin staining.Bar 100 #Am.

The percentage of bone at the implant surface was higher with some glass-ceramic composi tions than wi th hydroxyapati te compositions.The reason for this finding is not yet clearly understood.However, this observation indicates that there can be differences in the tissue response relating to the chemical composition of the various surface reactive implant materials.

279

The known balance of bone development and resorption could be observed in these experiments.There were, however, additional findings that indicated the preferential resorption of implant material at the edge of inserting bone trabeculae. Furthermore, the assumption that the resorption of implant material ,e.g. hydroxyapatite,is correlated with the establishment of the physiological relevant surface profiles was endorsed (Fig. 3) • In SEM and TEM the resorption was found to be correlated with the activation of osteoclast-like cells.This does not exclude other mechanisms of resorption or dissolution of soluble implant materials.

The development of surface profiles corresponding to Baud's curves at the edge of very different implant materials was also found when the implant cylinders were implanted into the diaphysis of long bones.An example with use of a cylinder of a glass-ceramic and titanium composite which was implanted in the humerus of a rabbit is given (Fig.4).This figure also demonstrates the effect of stress protection in the

Figure 4.Cylinder from titanium/glass-ceramic composite 84 days after implantation into the rabbit humerus.Surface profile of the bone regenerate resembling Baud's curves.Sawed section.von Kossa/fuchsin staining.Bar 1 mm.

cortical area near the implant cylinder which was bonded to the bone.Additional experiments with insertion of a tube made from non-bone-bonding material PTFE (Polytetrafluorethylene) in the humerus of rabbits also revealed the characteristic surface profiles.These structures contribute to an optimal

mechanical state to structures structures calculated

280

providing a homogeneous surface stress stresses.ln other words with these fracture is minimized.The analogy to

animals has been presented,

strength in avoid notch the risk for in plants and and used as a principle for construction (8).

Among aseries of retrieved total hip replacements (THR) (9) which proved to be clinically successful at postmortem examination of patients that had died from other diseases, the cemented metal stems were encased by hone cement and this polymer was penetrating between hone trabeculae providing a mechanical interlocking between the hone and the metal stem of the prosthesis (Figs. 5,6) • There was only a structural change of the bone trabeculae insofar as the hone trabeculae that were not encased in the hone cement displayed a smaller diameter, and thus some indication of resorption.lt is

Figure 5.Microradiography of a transverse sawed section of a retrieved cemented femural component demonstrating bone trabeculae encased in the cement and some smaller trabeculae outside the cement, and atrophy of the cortical bone due to stress shielding.Clinically successful implant for a number of years.Bar 5 rnrn.

speculated that this atrophy could be due to the stress protection.The trabeculae encased in the cement being more or less separated from their blood supply were not resorbed but attained to some minor degree from cellular attack and resorption.

281

Figure 6.Sawed section of the undecalcified and in PMMA embedded femur.Same specimen as in Fig.5.Giemsa staining.Sar 5 mm.

In contrast,a clinically failed and loose femoral component of a THR displayed histologically and in microradiography a seam of connecti ve tissue at the interface between the cement and the surrounding hone (9)(Fig.7,8).This bone developed a rather continuous layer simulating a hone ring or a bone frame very similar to the ring-like hone frame that was observed with the above mentioned animal experiments.The ring-like hone frame supporting the connecti ve tissue was anchored in thE! trabecular structure of the femur.The cortical hone was rather atrophie supposedly due to inactivity and stress shielding ami additionally due to osteoporosis.The most important finding is the soft tissue layer between the cement and the ring-like bone indicating mechanical instability between the implant and the hone.

282

Figure 7.Microradiography of a sawed section of a clinically failed cemented femoral component of a THR displaying a ring­like bone giving support to the soft tissue layer surrounding the cement and metal component. Trabecular and cortical bone with considerable atrophy.Bar 5 mm.

Figure 8.Sawed section of the same specimen as in Fig.7. Soft tissue layer between the cement and the ring-like bone.Giemsa staining.Bar 5 mm.

283

When analysing the metal conte nt of soft tissue surrounding failed femural or cup components of THR there was an astonishing number of cases wi th a high conte nt of metal in the reactive tissue product which was more or less found to be necrotic.For some cases that have been analysed by using atom absorption spectrometry, the values are plotted in Table 1.

TABLE 1

Metal content in soft tissue surrounding failed femur and cup components of THR removed surgically for replacements. Values in mg per kg dry weight. Number of specimens = n. Mean, maximal and minimal values.

Metal Fe Co Cr Ni

n 102 50 70 31 max 5629 2746 2646 739 min 101 4 0.7 3 mean 936 239 197 64

The necrosis of the reactive tissue was apparently slowly progressing and there were no signs of major cellular reaction against the necrotic tissue.Immunohistochemical staining of the macrophage population revealed some differences in the expression of proteins that are considered as representing the cellular outfit of macrophages.There was a number of macrophages that did not express the proteins normally present in these cells.Whether this result is due to the progressing necrosis of cells and tissue as mentioned ,or if this is a sign of disturbed macrophage metabolism and outfit cannot yet be answered. It seems at least that this problem should be analysed.

In conclusion,retrieved implants from patients displayed findings that follow principles found in animal experiments. There is a variety of different systems that control the tissue und cell reactions at various distances from the implanted materials.

ACknowledgements:Supported by the Ministry for Research and Technology,Bonn,Ol VG 8603,FRG.

284

REFERENCES

1. Gross,U.M.and Müller-Mai,C.,Hard materials-tissue interface: General considerations and examples for bone bonding and for epithelial attachment.In: Handbook of Bioactive Ceramics. Vol. 1, Bioactive Glasses and Glass­ceramics, Eds. T.Yamamuro, L.L.Hench, J.Wilson, CRC Press, Boca Raton, 1990, pp.25-39.

2. Brown,S.A.,Farnsworth,L.J.,Merritt,K.,Crowe,T.D.,In vitro and in vive metal ion release. J.Biomed.Mater.Res., 1988, 22, 321-38.

3. Smith,G.K. and J.Black, Estimation of in vivo Type 316 L stainless Steel Corrosion Rate from blood transport and organ accumulation data, ASTM STP 859, ed. A.C.Fraker and C.D.Griffin,American Society for Testing and Materials, Philadelphia, 1985, pp. 223-47.

4. Hench, L. L., Bioacti ve glasses and glass-ceramics: A perspective. In CRC Handbook of Bioactive ceramics, Vol. 1, Bioactive Glasses and Glass-ceramics, Eds. T.Yamamuro, L.L.Hench, J.Wilson, CRC Press, Boca Raton, 1990, pp. 7-23.

5. MÜller-Mai, Ch., Schmitz, H.-J., Strunz,V., Fuhrmann,G., Fritz,Th., Gross,U.M., Tissues at the surface of the new composite material titaniumjglass-ceramic for replacement of bone and teeth, J.Biomed.Mater.Res. 1989, 23, 1149-68.

6. Gross,U., MÜller-Mai,Ch., Fritz,Th., Voigt,Ch., Knarse,W., SChmitz,H.-J., Implant surface roughness and mode of load transmission influence periimplant bone structure. In Clinical Implant Materials, eds. H.Heimke, U.Soltesz, A. J . C. Lee, Advances in Biomaterials , Vol. 9, Elsevier Science Publishers, Amsterdam 1990, pp. 303-08.

7. MÜller-Mai, C. M. , Voigt, C. , Gross , U. , Incorporation and degradation of hydroxyapatite implants of different surface roughness and surface structure in bone. Scanning MicroscQPY 1990, 4, 613-24.

8. Mattheck,C., Engineering Components grow like Trees, Mat.­wiss.u.Werkstofftech. 1990, 21, 143-68.

9. Sokiranski,R., Metallartefaktreduzierung (MAR) in der Computertomographie am Beispiel der Hüftendoprothese. Inaugural-Dissertation, Freie Universität Berlin, 1990.

285

IN VITRO CYTOCOMPATIBILITY AND TISSUE REACTION TO CERAMICS

A. Pizzoferrato, E. Cenni, G. Ciapetti, S. Savarino, S. Stea

Laboratorio di Biocompatibilim dei Materiali da Impianto - Istituto Ortopedico Rizzoli

Via di Barbiano 1/10 - 40136 Bologna - Italia

ABSTRACT Ceramics are a class of materials since long used in prosthetic surgery, mainly because of

their in vitro biological compatibility. Our researches have been aiming for years at the compatibility evaluation of implant

materials by rneaI'l.s of in vitro tests, and both experimental and clinical implants. In vitro tests, that mainly resort to cell cultures, allowed both morphological and

functional reactions of various cells in contact with cerarnics to be evaluated. Experimental implants have allowed the study of the tissue re action to the implant, both

static and dynamic, of inert, porous and resorbable ceramics, as weil as the evaluation of the wear reaction to such materials.

The evaluations carried out on removed clinical implants permitted to find out the behaviour of some ceramics in the real conditions of their final use, that is when the prosthetic components, submitted to weight and motion, can be wom out.

Some conclusions can therefore be drawn from our global experiences on the biological compatibility of the ceramic materials.

INTRODUCTION The first studies on the use of ceramics in the construction of prostheses to be used in

surgery started about 20 years ago.

Initially much importance was given to ceramics evoking scarce or no tissue response;

subsequently a different approach was tempted, looking for ceramics able to interact with the

living tissues to establish an intimate bond.

From a chemical point of view, ceramics are macromolecules made of ions distributed in a

regular structure.

The bioceramics used in clinical applications can be divided into main groups on the basis

of their reactivity to the physiological environment, according to the following scheme:

286

BIOCERAMICS FOR CLINICAL APPLICA TIONS

NEARLYINERT

-alumina

-carbons

SURFACE-ACTIVE

-hydroxyapatite

-bioglasses

RESORBABLE

-TCP

-microporous apatites

A class apart is constituted by the composites in which the ceramics are in combination

with other materials. Ceramic coatings on metals or polymers can be considered in the same

way as the composites.

The main feature of ceramics is their biological compatibility due to the fact that their

constituting elements are normaIly found in the living organism (Ca, P, Mg, Na) or however

weIl tolerated by the organism (Ti).

From the mechanical point of view they have a high hardness, but their brittleness does

not always allow their clinical use.

Bioinert ceramics, such as alumina, have a great resistance to hydrolysis and undergo

no significant chemical alteration in the physiological environment. These ceramics get in

contact with the living tissues through a mechanical-type connection, as the tissue shapes and

models itself on the implant surface. They are mainly used in the construction of articular

components of orthopaedic prostheses and dental implant coatings, as weIl as the mandibular

crest augmentation in edentulous athrophies.

The class of carbons encompasses various types of carbon (LTI carbon or pyrolithic

carbon and UL TI carbon or 'vapor-deposited carbon') differing in the various fabrication

processes. Their drawback is a difficult shaping. The L TI carbon has however been used

successfuIly in the manufacture of heart valves for cardiac surgery. The carbons, very

promising but then scarcely used in the clinical application, have been experimented in the

construction of artificial tendons, of osteosynthesis plates and in the manufacture of implants

for odontostomatologic implants.

The ceramics with active surface or bioactive ceramies are able to trigger

controIled reactions in the physiological environment.

The bio~lasses with active surface were introduced in the late 60s under the hypothesis

that these materials, releasing ions in the matrix surrounding the implant, could stimulate on

their surface the growth of bone tissue, capable of bearing the physiological loads. The

bioglasses are a family of amorphous structure ceramics capable of a remarkable reaction with

the environment; the best known and most used is the Hench bioglass®, on the basis

formulation of which the other bioglasses have been developed.

The bioglasses bind directly to bone through:

- ion release in the surrounding matrix

2J37

- production, at the interface, of a kind of organie matrix that is later mineralized.

Both in vitro and in vivo, after an implant of the Hench bioglass (and generally of

bioglasses containing 45-55% by weight of silicium oxide),-a layer rich in silicium diffused

from the bioglass develops, as well as a layer rich in Ca-P, coming both from the bioglass and

ion deposition in extracellular fluids (1). The thickness of these layers can vary according to

the in vitro or in vivo conditions. Some years after the discovery of the bioglasses it was

demonstrated that also hydroxyapatite could evoke a similar behaviour.

Active hydroxYllPatite can be divided into adense and a macroporous form (pore

diameter> 100 Ilm). Both of them can be considered as osteoinductive materials, i.e. able to

stimulate the formation of bone tissue in areas where normally bone stern cells are present.

Bioactive HA thus permits a direct contact with the bone and, when macroporous, favours the

bone ingrowth inside the pores.

The tissue response to active surface ceramics results in a chemical bond between the

tissue and the implant. The active surface ceramics are used in the replacement of the ossicules

in the middle ear, as coating of orthopaedic prostheses and in dental and maxillo-facial

implants.

The resorbable cera mies have the property of being resorbable because they are

constituted by elements normally present in the organism.

However the degradation needs these material to be microporous, with pore diameter< 5

11m as de Groot in 1980 demonstrated (2), after having thoroughly reviewed several studies on

the dense, micro- and macroporous ceramics, that the dissolution process of "degradable"

ceramic occurs in two steps: extracellular dissolution of the necks among sinterized particles

and intracellular phagocytosis of the particles isolated in this way. The first step becomes

impossible in dense ceramics and very difficult in the macroporous ones: in this way porosity

can affect the degradation process.

The most used materials are some Ca-P salts with a particular Ca/P ratio and a particular

structure (Ca/P ratio between 1.5 and 1.67 and apatitic-type structure) such as

tricalciumphos.phate ITCP) and microporous HA.

Tricalcium phosphate, or TCP, belongs to the ceramics group based on calcium and

phosphate, and it is used in orthopaedics and odontostomatology as biodegradable substitute of

the bone.

The resorbable materials are mainly used as temporary support or filling material and

however for a limited lapse of time, being then replaced by living tissues. Typical application

of these materials are the filling of alveolar cavities and cystic bone cavities.

The coatings of various bioceramics on metals and polymers allow to obtain a tight bond

between implant and bone thanks to the chemical affmities of the coating.

288

Dense alumina, calcium phosphate ceramics (hydrooxyapatite and TCP) and bioglasses

were used as coating on various metals and polymers. A critically important issue is the long­

term integrity of the bond. When a ceramic is coated on a joint prosthesis, the purpose of the

coating is the fixation of the device: thus the coating must be stable. When a thin bioglass film

is applied on a porous metal surface, its function is to enhance bone ingrowth in the post­

operative period. If this happens, a shorter revalidation period is needed for the patient. On the

long term no ceramics coating persistence seems to be needed, as the bone ingrowth will

procede even in the absence of the ceramic lining. In such case a ceramic-to-metal bonding is not

important.

The coupling of HA and autologous bone permits to solve the quantitative lack of bone

chips in the fillings of wide bone cavities. Glass ceramics have been mingled with PMMA to

obtain bioactive cements; bioactive glasses have been coupled with metallic fibres to

manufacture bone-ingrowth prostheses able to enhance the development of adiacent bone. The

use of polylactic acid, degradable by the organism in a relatively short time, with various types

of fibres has led to the production of artificial tendons and ligaments and fixation plates with

bond characteristics with the surrounding tissues, avoiding at the same time unfavourable

reactions such the usual phragmentation in the carbon fibres that are not part of composites.

Three major schools of thought can be identified in the bioceramics composite field (3).

First, maintain the excellent biological behaviour of ceramics, whether inert or bioactive, and

mitigate the relative lack of mechanical properties by reinforcing it with an appropriate second

phase (e.g. active bioglass composites and polyethylene-hydroxyapatite composites). Second,

create materials which act as a temporary scaffold for tissue development; thus use bioresorbable

ceramics but provide them with mechanical and elastic properties which meet the functional

requirements of the application till the normal tissue is regenerated. Third, use ceramic fibres or

particles to create composites with a combination of mechanical and elastic properties which no

other material presents (i.e. carbon-fibre-reinforced composites, such as carbon-fibre-reinforced

carbon or carbon-fibre-reinforced polysulfone).

Ceramics biological reactivity is studied in our laboratory according to an experimental

protocol which includes :

- in vitro researches on bulk sampies or particles

- in vivo researches on bulk sampies, on coating, on porous granules and on particJes

- researches on tissue sampies from removed clinical implants

289

CERAMICS CYTOCOMPA TIBILITY The screening of the biological response to ceramics. is performed by means of in vitro

tests with cell cultures.

For some time in our Laboratory citocompatibility tests on various types of cells

have been carried out, aiming at verifying the primary toxicity of biomaterials (4).

The HeLa cells adhere to asolid alumina sampie and proliferate on its surface, in contrast

with what happens by using a sampie of toxic plastic.

Murine fibroblasts C3T3), if cultured near asolid sampie of pure alumina, gather around

and proliferate on the material (Fig. 1), while around a copper sampie, a zone without cells

develops, because of the degeneration and death of the cells closer to the meta!.

FIGURE 1. Mouse fibroblasts 3T3 growing as a monolayer in contact with big panicles

of pure alumina (Giemsa stain, x 40).

The growth of L929 fibroblasts is not hampered by the presence of solid hydroxyapatite

sampie; the viability of the cells is expressed by the intense coloration of their cytoplasm,

because of the uptake of the vital neutral red stain.

290

Even if powdered, thus with a much wider contact surface than the solid form, pure

alumina does not alter the functions of the cells belonging to the immunitary system, such as

periphera1 bl00d lymphocytes and peritonea1 macrophages (5).

The lymphOGytes stimulated with mitogen continue to proliferate; the maCTQpha&es easily

phagocytize large amounts of small-sized alumina particles (1-10 J.Un) without any altemtion in

their chemotactic activity.

By transmission electron microscopy the phagocytized alumina particles appear

incorpomted in vacuoles as if they were to be digested. Some bioglasses "doped" with various oxides (Cr203,Fe203, Cao, Al203 and others) to

enhance their adherence capability to metal substmtes, resulted in the formation of many 1tiaru ~when experimentally injected into mouse peritonea1 cavity.

Finally, ~nulocytes chemokinesis is not altered by the presence of pure alumina extmcts:

such cells maintain the orientation induced by the chemoanmcting substance (EAS).

TISSUE RESPONSE TO CERAMICS

Our experience on ceramie tissue eompatibility panly derives from the study of

experimental implants and partIy by the analysis of failed prosthesis in humans (6).

The experimental implants were made by using the ceramic in various forms.

The early experiments consisted in etherotopie (in the rabbit paravenebral muscles) or

onhotQpie implants (in femural epiphysis) of dense alumina sampies, and their histologie

observation at various intervals showed a good eoexistenee between alumina and tissues. Sueh

eompatibility is documented by the slight inflammatory reaetion and by a frequently oceurring

of continuity between alumina and tissue.

In the past we have earried out also implants of metal dental roots on animals with alumina

~ and isotropie earbon coatin~. {The roots were made of titanium with a 150 11m thiek

alumina eoating, or of vitallium eoated with a very thin layer (5000 A) of LTI earbon (Low

Temperature Isotropie Carbon»). The evaluation earried out after various load periods has

shown the approaehing of the tmbeeulae to the surfaee without fibrous tissue interposition.

More reeent experiments were earried out on porous surface ceramies, whieh permit the

device fixation to the bone tissue.

Some macmporous hydroxyapatite sampIes were implanted in rabbit femura! metaphyses.

They were 0.4 cm-sided eubes, and had a pore number per surface unit of about 12 eaeh mm2

(pore size ranging 22 to 862 11m).

The results showed that after 4 months no tissue ingrowth had occurred in the pores with

size < 200 11m. In the pores with size mnging from 200 to 300 11m conneetive vascularized

291

tissue rieh in eells was observed. In pores with diameter exeeeding 300 11m, the presenee of

newly-formed bone trabeculae ofvarious thiekness was notieed.

Granules of resorbable microporous HA with two different porosity pereentages, used as

filling material of bone defects, were implanted in rabbit diaphyses to evaluate their biologie

reaetion. The good tolerability was demonstrated by the intimate eontaet between HA and bone

tissue.

The measurements earried out on the two types of granules before and after surgery

showed that the HA with higher porosity (sintered at 1000°C) starts being resorbed after the

8th post-operative week, while the same material with lower porosity pereentage (sintered at

1 150°C) starts being resorbed only after the 12th post-operative week.

The powder, obtained by milling, of a eomposite material of earbon fibres and epoxy

resin. to be used for fraeture fixation plates, eaused a strong giant eell reaetion, in spite of the

good tolerability of the same implant material as a block in rabbit paravertebral muscles.

Sampies of an alumina-bioglass eomposite with porous surfaee (mean porosity of 400 11m

with a 290-510 range) made by a eore of dense alumina eoated by various layers of alumina

miero-beads (size 500-750 11m) were implanted in the distal metaphyses of rabbit femurs and

observed at different intervals (1,4;6,8 and 18 weeks after surgery) (7).

After one week it was possible to notiee around the implant the presenee of loose

eonneetive tissue, rich in eells, in osteoinductive attitude, interposed between the pre-existent

tissue and the implant surface. After four weeks the growth of mainly fibrous tissue occurred,

while inside the pores of bigger size (exceeding 200 11m) it was possible to notice how some

beads were already trapped by the bone tissue.

The microhardness analysis, carried out in order to quantify the mineralization rate and

thus the maturity of the bone ingrowth inside the pores, showed that the newly-formed bone

tissue had reached a hardness comparable to that measured in the bone far from the implant

site. The results, expressed in Vicker's grade, are reported in the Table 1.

Implant life

1 week

4 weeks

6 weeks

8 weeks

18 weeks

TABLE 1

Bone Tissue Mierohardness expressed in Vickers Degrees

Inside the pores

23.5

59.3

58.3

57.0

61.6

Far from the implant

54.3

55.8

57.3

60.7

59.2

292

In order to evaluate the tissue response to ceramic wear materials, scalar amounts of

alumina powder, obtained by means of a joint simulator, were inoculated in pig joints. The

joint tissue reaction, two months after the injection, was limited to a scarse development of

fibrous tissue with slight hypertrophy of the synovial membrane.

Our experience is completed by the examination of tissue sampies from removed hip prostheses made with ceramic components. The examined prostheses were of the ceramic­

ceramic or ceramic-plastic type. Among those with ceramic-ceramic components, some had an

alumina coating on their stern.

The study carried out on such removed prostheses (16 ceramic-ceramic and 7 ceramic­

plastic) enabled us, on the one hand, to examine the ceramic wear debris and the tissue

reaction; on the other hand to evaluate the behaviour of the bone tissue at the interface with

ceramic. Peri-implant tissues removed during revision surgery presented variable amounts of

ceramic material with different aspects according to the wear particles size. If their size is less

than 1 Jlm they look like amorphous powder; if they are about 3-4 Jlm they display a

polyhedric crystalline, sometimes needle-shaped structure; if they reach a bigger size (10-15 Jl)

they look like splinters.

The chemical nature of the particles we observecl. by light microscopy was confirmed by

means of microanalytic study carried out with scanning electron microscopy. Such method

permits to analyze the X-ray pattern, characteristic for each element, emitted by a sampie as a

consequence of the electronic bombing. By transmitted light microscopy the wear powder has

a dark brown aspect with yellowish shadings; the crystals have a slight amber colour with

brownish corners; the needle-shaped particles are not easily visible; the splinters look slightly

amber-coloured, but they do not display brownish outlines.

Under polarized light the small- and intermediate-sized wear particles partially and non­

homogeneously deviate light, creating various aspects; usually the splinters totally deviate the

polarized light acquiring a homogeneous lucency similar to that determined by polyethylene.

However the ceramic splinter differs from the polyethylene one because it is weIl visible also in

transmitted light.

The cellular reaction to ceramic material is not particularly sharp when the wear particles

have the crystal or splinter aspect and size. ActuaIly in these cases a modest histiocytic reaction

occurs, and most of the material is found in the extracellular spaces, without ever determining

giant-cell reactions (Fig. 2).

The histiocytic-type cell reaction is on the contrary extremely sharp in the presence of big

amounts of wear powder made of particles with size not exceeding 1 Jlm.

293

In such cases, quite rare, wide areas of cell necrosis were noticed, which are sometimes

bordered by fibrous tissue layers and assurne a lobular aspecl. Such areas are ofter adjacent

to weil preserved areas of intense histiocytic reaction.

FIGURE 2. Alumina powder, worn out from a hip joint prosthesis, is partially ingested

by histiocytes and partially free in extracellular spaces (HE, x 40).

These observations confirm once more the hypothesis that the extent of tissue response

to a foreign material depends only partlyon its chemical composition, while it is strongly

influenced by the physical characteristics (shape and size) of wear particles.

In the prostheses wh ich were removed because of pain in spite of the stern stability, we

observed (in two cases) in ample areas a tight contact between the bone and the ceramic

coated surface, without fibrous tissue interposition.

CONCLUSIONS

The living tissue response to the presence of the bioceramics we examined can be

summarized in basic observations:

- the biocerarnics assayed in vitro in bulk or particulate form did not show any toxicity;

294

- direct contact with the bone is noticed, without interposition of connective fibrous tissue,

along wide areas of the implanted ceramic materials;

- the hone ingrowth process is regulated by the pore size and porosity percentage;

- the resorption of microporous ceramic is conditioned by the porosity percentage;

- wear particles determine a sharp histiocytic reaction and necrosis only in the rare cases in

which they are small-sized and are produced in great amounts. However they usUally have a

larger size and induce scarce histiocytic reaction and never determine giant cell reaction.

This study was supported by Consiglio Nazionale delle Ricerche, grant n°

9O.01364.CT11 and by Ministero della Sanita, Ricerca Corrente IOR, Area 1.

REFERENCES 1. Wilson, J., Pi gott, G.H., Schoen, F.J. and Hench, L.L., Toxicology and

biocompatibility of bioglasses, 1. Biomed. Mater. Res., 1981,15, 805-817. 2. de Groot, K.,Bioceramics consisting of calcium phosphate salts, Biomaterials, 1980,

1, 47-50. 3. Ducheyne, P., Bioceramics : Materials Characteristics versus In Vivo Behavior, L

Biomed. Mater.- Res., 1987, 21, 219-236. 4. Pizzoferrato, A., Vespucci , A., and Ciapeui, G., Valutazione della compatibilita

biologica dei materiali da impianto mediante colture in vitro, Minerva Ortopedica, 1982,33, 801-807.

5. Pizzoferrato, A., Vespucci, A., Ciapetti, G., and Stea, S., Biocompatibility testing of prosthetic implant materials by cell cultures, Biomaterials, 1985, 6, 346-351.

6. Pizzoferrato, A., Evaluation of the tissue response to the wear products of the hip joint endo- arthroprosthesis, Biomat. Med. Dev. Art. Org., 1979,7 (2), 257-262.

7. Pizzoferrato, A., Toni, A., Sudanese, A., Ciapeui, G., Tinti, A., and Venturini, A., Multilayered Bead Ceramic Composite Coating for Hip Prostheses : Experimental Studies and Preliminary Clinical Results, J. Biomed. Mater. Res., 1988,22, 1181-1202.

295

BIOCERAMICS IN OR1HOPAEDIC SURGERY: KNOW HOW STATUS AND PRELIMINARY RESULTS

S. Giannini, A. Moroni, G. Coppola, L. Ponziani, A. Ravaglioli*, A. Kraiewski*, A. Venturini**, M. Pigato**,D. Zaffe.***. Orthopaedic Dept University of Bologna, Rizzoli Institute, Bologna * IRTEC CNR Faenza ** Surgica1 Veterinary Dept. University ofBologna *** Anatomy Institute, University of Modena

ABSTRACT

Bioceramic materials for onhopaedic surgery can be distinguished in two groups: bioinen and bioactive ceramics. The bioinert ceramics provoke a minimal tessutal reaction; the bioactive ones stimulate the bone ingrowth and a bone-materiallink capable to support physiological loads occurs. Alumina (Al) is an inert bioceramic which presents an excellent mechanical strenght and a very low wear. For this reason Al is used in manufactoring of acetabular component andfemoral head of total.hip.arthroplasties (THA) Al is also used as coating of metallic hip prosthetic stems and for small joint prostheses. As bioactive ceramics we analyzed bioglass and hydroxyapatite.(HA). Bioglasses are actually considered very promising materials but they have not still a clinical application. HA is the more used bioactive material in orthopaedics. The dose interface link with bone tissue allows the use of HA in hone cavitiesfilling ofvarious skeletal diseases and as coating ofmetallic prostheses.

INTRODUCTION

About 20 years ago the first researches for the use of bioceramies in orthopaedic surgery.were began. Initially only bioinert materials were studied but recently several Authors started the study on bioactive ceramics. In the last 20 years the research in this field has notably increased. Now bioceramics represent an important attempt to solve various problems related to orthopaedic surgery: prosthetic implants, bone cavity filling, internal fixation devices.

The purpose of this study is to report our preliminary experimental and clinical results and to analyze the literature data on the utilization in orthopaedic surgery of the bioceramic materials.

BIOINERT CERAMICS

Among bioinert ceramies, AI represents the more studied material. Prom several years Al has been considered a very ductile material for its bioirtness, great

296

hardness, mechanical strength and very low wear. Experimental trans-cortical implants in bone, demonstrated AL is almost completely

surrounded by bone tissue from which Al remains separated by an average 100 microns thiek space.(1).

Poli-crystalline Al has been utilized from several years for acetabular components and femoral heads of THA.For this kind of employment it is mandatory the use of granules with an average size smaller than 3 microns; in this way it is possible to obtain prosthetic components with very smooth surfaces.

At the present time long lasting clinical follow-up regarding cemented and uncemented T.H.A. with Al components are available.

Clinieal results are positive (failures were due to design errors or mechanical weakness of the material). Nowadays technology is able to produce Al prostheses with the necessary mechanieal properties.

Moreover Al was used as HA on prosthetic components coating to improve the bone­prosthesis uncemented fixation. This kind of coating was obtained by the plasma-spray technique.

On 1976, at the Istituto Ortopedico Rizzoli, a THA with Al femoral head, Al acetabular component and a titanium stern coated with Al was produced. Clinical and radiographie results are very encouraging 13 years later.(2)

Al was also used for lenghtning of metatarsus (3) and for metacarpo-phalangeal arthroplasties (4).

On 1988 a metatarsus-phalageal total joint prosthesis was developed at the Rizzoli Institute. The main characteristics of this prosthesis are: anatomical design and uncemented fixation. Both the metatarsal and phalangeal components are Al made. The clinical results, at a two years average follow-up , are positive. (5)

Moreover Al is suitable to make small devices for the intern al fixation of bone fractures (Fig I).

BlOACTIVE CERAMICS

Bioactive ceramies are materials able to stimulate bone-ingrowth and an interface link able to support physiologicalloads.

The interaction between bioactive ceramics and bone is due to a complex combination of biologie, mechanical, physical and chemical processes. Various factors influence these processes such as the deformation of organic molecules, the discharge of chemical elements by the bioactive ceramics and the phagocytosis processes of particules released by the implanted materials. These phoenomena are not necessary costant during time and equal in every implant site.

The main factors conditioning the result of the implant are: implantation site, tessutal trauma, structural properties, material surface and disgregation of material at the interface with bone.(6-7-8).

Bioglass is probably the more experimentally studied bioactive ceramic material (6-7). When bioglass is implanted in bone, the bioglass alkaline ions filter at the surface of the

material, forming two layers: a first deeper reactive layer, rich of Silicium and a more superficial second one, rich of Calcium and Phosphorous. This second layer is fundamental for bone-implant link and its formation seems to be due to ions coming both from bioglass and body fluids (9).

Bioglasses, like all the bioactive ceramics, are not osteoinductive or osteogenic materials, butjust osteoconductive. Thus when implanted in soft tissues they do not provoke the formation of calcified tissue but facilitate the osteogenetic process when they are implanted in the bone. These materials were proposed as coating of metallic prostheses in order to improve the bone-implant interface but they are not still utilized in clinical orthopaedics.

HA is another bioactive material suitable for employment in Orthopaedics. Many Authors demonstrated that experimental bone implants ofHA form a close link

with the surrounding bone. (10-11-12).

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Figure 1 a-b) ankle fracture in a 26 year-old male. c-d) post-operative X-rays 6 months later the internal fixation with Al screws.

There are different types of HA according to chemical, physical and structural differences. The structural properties are able to modify the results of the HA implants. The more important characteristics are: kind of surface, size and shape of grains, structure and size of crystals, size, shape and distribution of pores. A size porosity more than 100 microns allows the penetration of bone (9).

The possibilities to employ HA in orthopaedic surgery are: bone cavities fllling, devices for internal fixation, coating of metallic prostheses.

HA granules, associated or not to bone grafts, can be used to fill bone defects in different diseases: bone cavities following tumor removal, bone stock loss for fractures, non­unions or revisions of failed prostheses.

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Moreover we use HA granules in the vertebral arthrodeses due to scoliosis (Fig. 2). 14 cases of bone defects due to the removal of benign bone tumors were operated on at the Rizzoli Institute (bone cysts, aneurysmatic bone cysts, fibrous dysplasia). The average follow-up is 16 months. Healing of the defects was observed in all the cases. (Fig. 3-4)

The present follow-up does not allow to determine which is the fittest physical structure of the granules to obtain the best osteoconductive effect and if the association with bone grafts is really necessary.

On the last years we started also the use of HA in association to hydrocortison and fibrin glue in the treatment of bone cysts.

Moreover we enphasized HA can be also usefully utilized to coat metallic prostheses allowing a firm fixation of the uncemented arthroplasties due to areal link between bone and the prosthesis surface. The experimental and the preliminary clinical results related to this utilization of HA are very positive and further studies on this topic are warranted.

Figure 2 a) C4-C5 vertebral dislocation in a 32 year-old male b) vertebral arthrodesis with HA granules

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Figure 3 a) huge non-ossified fibroma of the lower third of the femur in a 17 year-old female b) post-operative X-ray 4 yrs. later the HA filling

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Figure 4 a) aneurysmatic hone cyst of the humerus in 14 year-old male b) post-operative X-ray 2 yrs.later HA cavity filling

REFERENCES

1 Moroni A., Giannini S., Zaffe D., Ravaglioli A., Kraiewski A., Venturini A., Pompili M., Pezzuto V., Trinchese L. "Comparative histological and chemical-physical analyses on Alumina, bioactive glass, bioglaze and hydroxyapatite sheep hone implants" Proceedings of the 2nd International Symposium on Ceramies in Medicine,Heidelberg, Germany, September 1989.

2 Trentani C., Montagnani A., Vicenzi G. "Ten year. follow up of uncemented total Hip prosthesis with alumina acetabulum and titanium femoral stern by plasma jet technique" Proceedings of the Sixth European Conference on Biomaterials, Bologna, Italy, September 14-17, 1986.

3 Yonenobu, K. Tada, K. Tacaoca, Y. Tsuyuguchi, K.Ono "Elongation of brachymetatarsy with ceramies implants:a roentgenographic evaluation of its utility"J.B.M.R. 20, 1249-1256, 1986

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4 M. Minami, J. Yamazachi, K. Sadatoshi, I. Seiichi "Alumina ceramic prosthesis arthroplasty of the metacarpo­phalangeal joint in the rheumatoid hand" The Journal of Arthroplasty 3, 157-166, 1988

5 S. Giannini ,A. Moroni, L. Trinchese, M. Pompili, V. Pezzuto, A. Ravaglioli, R. Martinetti, C. Farina. "Alumina total joint replacement of the first metatarso-phalangeal joint" Proceedings of "Bioceramics and the Human Body" Faenza, Italy, April 2-5 , 1991

6 L.L. Hench, RJ. Splinter, W.c. Allen, T.K. Greenlee "Bonding mechanisms at interface of ceramic prosthetic materials" J. B. M. R. Symp. 2,117-141,1972

7 L.L. Hench, E.C. Ethridge "Biomaterials, an interfacial approach" Academic Press New York, 1982

8 T. Nakamura, T. Yamamuro, S. Igashi, T. Kokubo, S. Ito "A new glass ceramic for bone replacement: evaluation of its bonding to bone tissue" J. B. M. R. 19,685-698, 1985

9 P. Ducheyne "Bioceramics: material characteristics versus in vivo behavior" , 21, 219-336, 1987

10 J. Lemons "Orthopaedic uses of calcium phosphate ceramics, taskgroup report" Bioceramics Materials Characterization versus Acad. Sci. New York 1987

11 W. Van Raemdonck, P. Ducheyne, P. Demeester "Calcium phosphate ceramies" Metal & Ceramic Biomaterials Vol. 11 CRC Press Boca Raton 143-166, 1984

12 A. Ravaglioli, A. Kraiewski "Bioceramica e corpo umano "Faenza Editrice 1984

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OSSEOINTEGRATION OF HYDROXYAPATITE-COATED AND UNCOATED BULK ALUlIINA IMPLANTS IN THE FEMUR OF GOTTINGEN MINIPIGS

IlECHANICAL TESTING OF BONDING STRENGTH

JOACHIM ORTH, SERENA MACEDO, AXEL WILKE, PETER GRISS

Orthopaedic Department of Philipps-University, Baldingerstraße, 3550 Marburg, Germany

ABSTRACT

The bonding strength of uncoated, glass-coated and hydroxyapatite-coated cylindrical alumina implants was tested after implantation in the femora of Göttingen minipigs. HA-coatlng resulted in detachment of the coating during push-out test already 4 weeks after implantation at app. 4 MPa. A com­parable bonding strength was reached In uncoated implants after 8 weeks and never was reached by only glass-coated implants.

INTRODUCTION

Blocompatability of an implant can be defined by mechanical and biological performance. MaterIals for tooth replacement must possess a high rate of biocompatibil1ty because of their intimate contact to mucosa and bone.

Alumina-ceramlc Is a proven biologically acceptable substance wlth ex­cellent tribological propertIes (3, 4, 5, 7, 8). Its high mechanical strength allows the appllcation as a dental Implant (2, 6). As a dental implant alumina is characterized by lack of bleeding and plaque formation (10).

Not only bone ingrowth but also an efflcient gingival sealing is decisive for the success and the stability of a dental implant. The appearance of epithelial downgrowth in contact to alumina-ceramlc implants is well known. In contrast to that, a new gingival attachment apparatus and a elose bone contact is developed when using prestressed dense sintered hydroxyapatIte (10). On the other hand poor mechanical propertIes of hydroxyapatite ceramies don't allow their appl1cation under loaded conditions. Thus It could be reasonable to combine the advantages of both materials.

Plasmaspray-coating of implants in general presupposes a rough surface, whlch in case of alumina ceramic implants is producable only by use of expensive diamond materials. The use of a bond co at of glass between alumina and HA-ceramic Is an alternative possibility.

OBJECTIVES

The purpose of our study was to investigate the effect of a porous HAP-

303

ceramic-coating applied onto alumina implants on the mechanical bon ding strength to bone in comparison to uncoated and only glass-coated implants.

MATERIAL AND METHODS

Implant characteristics: Cylindrical rods of alumina measuring 4.5 mm in diameter and 15.0 mm in length were prepared. Identical implants were coated with glass (70% SiÜ2. 16% Ah03. 12% Na2C03. 5% CaC03. 4% MgO) by a melting process.The third type of implants were additionally coated with hydroxyapatite-ceramic by plasmaspraying technique. The hydroxyapatite coating (degree of purity )98%. high cristallinity. app. 20 % microporosity) had an average thickness of approximately 200 micron.

Surgical technigue: Under sterile conditions five plugs were inserted into monocortical drill holes of each femur of 12 mature. female Göttinger miniature pigs. To study only the bone-bonding properties each drillhole was exactly sized. allowing easy insertion without any undue laxity.

The animals were sacrificed 4. 8. 12. 24 weeks after surgery. After har­vesting both femora and X -ray investigation. they were prepared for further mechanical testing aad histological analysis. Each femur was cleared of soft tissue. The distal and proximal part of the femur containing implants for histological examination were remooved.

Mechanical testing: The remaining part of the femur was cut into three sections. After cutting the sections longitudinaily to expose the intra­medullary end of the plug. the surrounding cortical bone was milled in order to achieve a flat surface directly around the implants. Oue to this pre­paration the bonesegment could be placed in a rectangular position to the specimen holder without the risk of slipping sideways. Thus the position allowed accurate lining of the loading axis with the long axis of the plug (fig. 1). The implants were pushed out from the surrounding bone using an Instron test machine wlth a crosshead speed of 0.5 mmimin.

F

J

0= SONE D=STAMP

Figure 1. Test setup during push-out tests .

The force needed to loosen the implant was determined from load-displace­ment curves. To calculate the shear strength of the interface. the extraction force was divided by the total bone contact area gained from light micro­scopic semiautomatlc picture analysis (Leitz ASM 68 K) . It was calculated as: A=3.14xdxh. where "d" is the diameter of the plug and "h" is the cortical thickness ot the femur in contact to the implant.

Histologlcal evaluation: Two of the five implants inserted in each femur

304

were subjected to histological evaluation. After formalin fixation and alcohol dehydration the implants were embedded in pOlymethylmethacrylate and pre­pared by glueing-sawing and grinding technique. The 100 micron thick sec­tions were stained with toluidinblue solution and were examined by trans­mitted light microscopy.

Scanning electron microscopy (SEM): Retrieved implants and polymethyl­methacrylate embedded bone specimen after push-out test underwent SEM analysis after sputtering with gold in order to detect the failure mode.

Statistics: Mean and SEM (standart error of the mean) are listed together. Statistlcs were carried out with SPSS-program. After multiple re­gression and tkorr -test we proceeded with MANOVA. Signlficances were proved by aposteriori Scheffe-test. All results were significant at a level of p<0.06.

RESULTS

RadiographY: We observed intramedu11ary new bone formation around a11 HAC-coated implants already 4 weeks postoperative. The implants were weIl integrated into the surrounding bone. Contrary to that, a11 glass-coated implants showed radiolucent llnes along the surface until the end of the experiment. Simllar radiolucent lines around the uncoated alumina-lmplants dissapeared 8 weeks after implantation.

Mechanical testing: The results of the push-out tests are summarized in table 1. The significant differences of interface shear strength depend on the type of material and time after implantation.

TABLE 1 Mean(MPa) and SEM of shear-strength related to time after implantation; n = number of sampies, Ab Oa =alumina, HAC=HA-coated, G=glass-coated

(weeks) Ah03 n HAC n G n

4 0.80±0.45 3 3.96±0.32 5 0.06±0.03 4

8 4.84±.0.71 6 8. 20±,1. 50 6 0.46±0.16 6

12 5.88±,0.92 6 10.29±1.63 6 1. 15±.0 • 44 5

24 5.01±,0.99 6 5.78±1.17 6 0.70±0.12 6

4 weeks after implantation there was already a signifacantly higher shear strength of 4.0±0.3 MPa for hydroxyapatite coated plugs compared to uncoated alumina cylinders (fig. 2 a).

In contrast to these findlngs a11 glass-coated specimen could be moved by very small forces, with a maximum interface shear strength of 1.2±0.4 MPa. The dlfference in shear strength in comparison to uncoated plugs after 8,12 and 24 weeks was slgnificant (fig. 2 b).

Uncoated implants showed little shear strength after 4 weeks and an increase after 8 weeks. They reached an ultimate shear strength of 6.8±0.9 MPa after 12 weeks. The hydroxyapatite-ceramic coated plugs always failed at the interface to glass and not to bone (fig. 3). The apparent decrease of shear strength from 12 to 24 weeks for each material was not significant.

SEM findings: With SEM it was obvious that the localisation of the rupture during push-out was the interface between HA - coating and gl ass coating (fig. 3). On the rough surface of the glass coating hydroxyapatite

305

particles could be recognized. The intimate contact of new formed bone tissue to the hydroxyapatite-ceramlc surface was very impressive (flg 4) .

2

o

-7.

• p( O~

2

o -2 .r; - 2

tn

• •

7.

o

-2

• p(O.05

-4L---------~------~----------'-4 4 8 12 7.4

-4L----4------~8~-----,2~----~2-4~-4

weeks oIe,. -mplon olion weeks ofte,. Implonlo ion

Figure 2 a. Results of statistical analysis of shear strength. HAC -coated versus uncoated implants

Figure 3. SEM photograph showing HAC-coated implant surface after push out. Left side showing re­mained coating on im plant surface. right side showing the glass-Iayer after disconnection of HAC-coating

(Magn. lOOOx).

Figure 2 b. Results of statistical analysis of shear strength. uncoated

versus glass - coated implants

Figure 4. SEM photograph of bone segment after push out. Remaining HAC-coating (top) in intimate contact to weIl matured lamellar bone (bottom) 8 weeks after implantation

(Magn. 89x)

Histology: At 4 weeks the histological seetions showed proliferating woven. still unmineralized bone tissue both along the hydroxyapatite coating and along cortical bone. There was a very close bone contact to the implant

306

without ftbrous tissue interlayer. We observed a time dependend maturation of bone tissue, which was nearly ftnished 12 weeks after implantation.

The uncoated alumina implants after 4 weeks were nearly completely surrounded by ftbrous tissue with only a few direct bone contacts. The thickness of this ftbrous interlayer decreased over the time. After 24 weeks it was still present, the rate of direct bone contacts had increased.

All glass-coated plugs showed an interposition of fibrous tissue between bone and glass surface at 4, 8, 12 and 24 weeks and an inflammatory tissue response up to 12 weeks after implantation.

Biodegradation of the hydroxyapatite coating by macrophages could be observed in contact to ftbrous tissue. Direct contact seemed to l1mitate this process. Some HA partlcles were completely integrated into new bone.

DISCUSSION

The differences between uncoated and HA-coated implants were striking. Mechanical testing proved an increase of bonding strength by factor 3

to 4 at 4 weeks after implantation through the presence of HA-coating. The interface shear strength of plasma sprayed apatite implants was 3.96±0.32 MPa at 4 weeks. HA coated plugs always falled at the substrate-coating interface. This impl1es that the bonding of HAC to bone is even stronger.

Besides that, it was remarkable that the substrate bon ding strength of apatite coating was measured 2 to 3 times higher in vitro (20 MPa) than in vive (4-10 MPa). This points to an existing infiuence of physiological environment onto the substrate-coating interface.

The shear strengths we report for glass-coated specimen were much lower than those in the Uterature for glass coated implants (1). Glasses are manufactured in various compositions (9, 11, 12). Each component influences biocompatib1l1ty. Oxides of alumina or titanium for example stab1l1ze a glass but also decrease biocompatibllty as happened. The chemie al substance pz Oll leads to better integration into bone, but was not a component in the glass used in this study. The obvious inflammatory reaction upto the 12th week close to the glass surface should be taken into consideration when judging the biocompatib1l1ty of glass as a bond coat.

Our current results indicate that a larger amount of bone invades faster in the presence of HA coating and a reliable early fixation is guaranted without any fibrous interlayer.

Glass-coated implants have not led to any implant fixation in bone.

CONCLUSION

Hydroxyapatite coating can enhance and accelerate the anchorage of alumina implants in bone tissue. Further long-term studies are needed in order to analyse the substrate-coating interface and the effect, the physiological environment has on it. All in all the existing results in our opinion are encouraging for further development of methods for coating of alumina implants with HA. In our implant modification the HA-coating-alumina inter­face bonding was exposed to high shear-strength, which caused fracture bet­ween coating and substrate already after 4 weeks. This problems could probably be diminished by improvement of the geometrie design. This could be achieved by screw type implants or other comparable surface structures.

REPERENCES

1. Blencke,B.A., Brömer,H. und Deutscher,K.K., Experimentelle Untersuchun-

307

gen über die Verankerung von belasteten glaskeramischen Implantaten im Femurschaft, Biomed Techn, 1979, 24, 294-302.

2. Bolz,U., Hüttemann,R.W., Kirschner,G. and Sturm,J., Indication for the cHnical use of combined auto-alloplastic tooth reimplantation with alu­miniumoxide ceramic, Dental implants, ed G. Heimke, C. Hanser Verlag, München, Wien, 1980, pp 81-87.

3. Boutin,P., Arthroplastie de la hanche par prothese en alumine fritt(~e, Rev. Chir. Orthop., 1971. 68, 229-246.

4. Dawihl,W., Mittelmeier,M., Dörre,E., Altmeyer,G. and Hanser, U., Zur Tribologie von HÜftgelenksprothesen aus Ab 03 -Keramik, Med Orthop Techn, 1979, 99, 114-118.

5. Dörre, E., Aluminiumoxidkeramik als Implantatwerkstoff, Med Orthop Techn, 1976, 4, 104-105.

6. Ehrl,P.A. and Frenkel G., Experimental and cHnical experiences with blade-vent abutment of Alz Ü3 -ceramie in the shortened dental row situation of the mandible, Dental implants, ed. G. Heimke, C. Hanser Verlag, München, Wien, 1980, pp 63-67.

7. Griss,P., Silber,R., Merkle,B., Haener,K., Heimke,G. and Krempien,B., Biomechanically induced tissue reactions after A12 Ü3 -ceramie hip joint re placement. Experimental and early clinical results, J Biomed Mater Res 1976, 7, 519-528.

8. Heimke,G., Beisler,W., von Andrian-Werburg,H., Griss,P. and Krempien,.B, Untersuchungen an Implantaten aus A12 Ü3 -Keramik, Ber Dt Keram Ges Nr.l, 1973, pp 5-8.

9. Hench,L.L., SpHnter,R.J.M., Allen,W.C. and Greenlee,T.K., Bonding mechanisms at the interface of ceramic prosthetic materials, J Biomed Mater Res, 1971. 2,117-141.

10. de Lange,G.L., de Putter,C., de Vos,R. and de Laat,R. Perumucosal dental implants, The relations hip between bone anchorage and the quality of surrounding gingival tissues, Biological and Biomechanical Performance of Biomaterials. ed. P.Christel, A.Meunier and A.J.C.Lee, Elsevier. Amsterdam, N. Y. Tokyo, 1986, pp 519-524.

11. Strunz,V., Bunte,M., Stellmach,R., Gross,U.M., Kühl,K., Brömer,H. and Deutscher,K., Bioaktive Glaskeramik als Implantatmaterial in der Kiefer­chirurgie, Dtsch zahnärztl Z, 1977, 32, 287-290.

12. Ungethüm,M. and Fink,U., Bioaktive Werkstoffe. Eine kritische übersicht, Z Orthop, 1988, 126, 697-708.

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THROMBORESISTANCE OF Ti 6AI4V, COATED WITH A THIN FILM OF TURBOSTRATIC CARBON, FOR CARDIOVASCULAR APPLICATIONS.

M.A. GATTI, E. MONARI, M. DONDI, G. NOERA, G. FATTORI Laboratory of Biomaterials,

University School of Medicine, Modena, I

F. VALLANA, S. RINALDI, E. PASQUINO Sorin Biomedica

Cardiovascular Department, Saluggia, I

ABSTRACT

The pyrolytic carbon is weil known to have the best haemocompatibility hence it has been used in manufacturing cardiac valve components. Unfortunately the high temperature operative conditions in which the PVC (pyrolytic carbon) is obtained restrict its applications to a few types of graphite substrates. Sorin Biomedica has developed a new method of PVD (physical vapour deposition) which allows to coat, with a carbon film having the same turbostratic structure (Carbofilml. temperature­sensitive materials like pOlymers and metals without modifying their physical and morphological properties. In the present study we evaluated the haemocompatibility of a Carbofilm coated Titanium alloy (Ti6AI4V), that is going to be used for the manufacture of a new bileaflet mechanical valve. The Gott's ring test was applied by mean of placement of short tubular sam pies of candidate metals in four sheep jugular veins. Six Titanium and two Stellite 25 Gott's rings (i.d. 7.8 mm; lenght 4.5 mm) were machined and the inner surface was carefully smoothed and polished. The Stellite rings were left uncoated while the Titanium rings were treated as follow: two were Carbofilm coated on the inner surface, two only partially coated and the others were left uncoated. The reason of partially coating some sam pies with Carbofilm is to create a carbon- metal couple that could be present in weared valve housing areas. This couple could induce in priciple a galvanic phenomena with localized thrombotic phenomena. After the implants blood flow throughout the rings was measured for three hours and macroscopical and microscopical evaluations were performed after the explants. The results confirm the good haemocompatibility of the uncoated metal rings due to the high quality surface polishing and quality control checks. The totally and partially coated Titanium rings showed the same results with very low protein and blood cells adhesion to the metal surface; no pulmonary infarctions due to thrombotic occlusion were observed. No significant blood flow reductions have

309

been recorded and the microscopical examinations of the border line, between the uncoated metal and the coated one, did not show any localized thrombotic deposition confirming the absence of galvanic phenomena between Titanium alloy and high density turbostratic carbon ICarbofilm).

INTRODUCTION

Pyrolytic carbon is a high-density material with turbostratic structure composed of

either pure or silicon-alloyed carbon microcrystals.

These properties distinguish it from the other materials consisting of carbon, such

as graphite, diamond and glassy carbon [1,21. Its properties stem from a unique

combination of physical and chemical characteristics, such as chemical inertness,

isotropy, low weight, compactness, elasticity and high strength [2,3,41.

The characteristics of this manufacturing process have so far impeded the more

extensive employment of pyrolitic carbon for prostheses [51.

The temperature required (> 1200 0 C), for example, means that the choice of

substrate is confined to a few types of graphite.

Since 1985 the Sorin Biomedica's laboratories have perfected an original process

for the deposition of thin films 10,2 - 0,3 pm) of turbostratic carbon·ICarbofilm) on

heat sensitive substrates, such as metals and polymeric fabrics already widely

employed in biomedical field [61.

With this key technology Sorin Biomedica had the possibility to realize a new

bileaflet valve IBicarbon) with a prosthetic housing made of Ti 6AI4V that

guarantee a higher mechanical reliability, safety and best fluidodynamic

characteristics. The operative result consists in a titanium made housing with a

Carbofilm coating that confers to the housing surfaces the same

haemocompatibility of pyrolytic carbon.

The thromboresistance test with Gott's rings, implanted into venous blood vessels

without anticoagulant therapy, represents one of the severest methods to

investigate the potential employment of a biomaterial for cardiovascular

applications.

Objective not less important has been the confirmation that the galvanic potential,

negligible and not able to promote pitting [7,8,91. at the Carbofilm-Ti 6AI4V

interface is not responsible of localized thrombotic processes.

310

MATERIALS AND METHODS

The in vivo evaluation of the Ti6AI4V thromboresistance has been realized

following the method suggested by Gott and others researchers. They evaluated

different bio materials realizing rings and testing them in vivo into the arterious or

venous vessels [11,13,14,15,171. This techniQue was applied to study the

haemocompatibility of many polymers [11,12,15,161. but rarely of metals. They

usually evaluated the thrombotic potential of the tested biomaterials by mean of

blood flow variations du ring the in vivo test and after the explants with the

standard macro- and microscopic examination techniQues [11 I.

Eight60tt's rings were manufactured (i.d. 7 .8 mm and 4.5 mm long) using both

Stellite 25 and Titanium (Ti6AI4V) alloys.

After the mechanical machining the rings were internally levigated applying the

Sorin Biomedica's techniQues usually employed for the polishing of the metal valve

components.

Figure 1. Gott' s ring manufactured with Ti 6AI4 V. The inner side presents four

sectors Carbofim coated.

311

Particular ca re was used in the quality controls for the smoothing of the metal

surfaces in order 10 identify also minimal imperfections.

The Titanium alloy (Ti 6AI4 VI was used for the manufacturing of 6 rings (2 totally

coated with Carbofilm, 2 only partially coated and 2 controls were left uncoatedl

while the Stellite 25 alloy was employed for the manufacturing of two uncoated

rings to be used as negative controls (figure 11.

Four Merinos sheep were clinically selected with a homogeneous weight

(60 ± 5 Kgl and with anormal haemocoagulative profile in order to exclude

individual interferences of the animal model on the test.

After the general anesthesia and the preparation of the operative field, the jugular

veins were surgically exposed throughout a median incision of the ventral surface

of the neck.

The flow pattern of both jugular veins was immediately assessed, before to

introduce the Gott's rings, by an electromagnetic flowmeter (Carolina Inc. - USAI.

The rings insertion, by mean of a special holder, was realized throughout a

longitudinal incision of the jugular vein after a proximal clamping of the blood flow.

The rings after their positioning were fixed externally with a 2/0 silk suture

(Ethicon - USA) while the longitudinal in cis ions of the jugular vein were sutured

with 6/0 polypropilene (Ethicon - USA). For each animal both the jugular veins were

checked for three hours, during the test, in order to investigate the flow pattern

modifications.

TABLE 1

Experimental scheme of thromboresistance test with Gott's rings.

SHEEP N° LEFT JUGULAR VEIN RIGHT JUGULAR VEIN

Ti 6AI4V Stellite

2 Stellite Ti 6AI4V

3 Ti 6AI4V + Carbofilm Ti 6AI4V + Carbofilm partial coating total coating

4 Ti 6AI4V + Carbofilm Ti 6AI4V + Carbofilm total coating partial coating

312

The flow measurements were assessed every 20 minutes and the pharmacological

treatment did not include any anticoagulant or antiplatelet treatment.

The experimental plan is reported in tablel.

At the end of the test the Gott's rings were removed harvesting at list 1 cm on

both sides of the native vein. The silk ligature was cut away and the rings carefully

examined.

RESULTS

The blood flow measurements did not show any flow modifications that could be

considered significative for the tested alloy rings. This result could be due to the

high basal thromboresistance of these tested metals and more over to the strong

increasing of haemocompatibility induced by the high quality surface polishing. The

observed slight flow variations were attributed to systemic modifications of the

blood flow with compensation between the right and the left veins.

The data reported in literature are referred to polymers with a broad range of

thrombogenicity that in many ca ses induced a drastic reduction of the venous

blood flow [10,121. These results could also depend from the smoothing of the

polymer surfaces on whose were not possible to reach a finishing comparable with

those we obtained on the metals,

The recorded flow data were plotted comparing in the same graph the right and the

left jugular vein blood flows (figure 2 and figure 3).

The macroscopical evaluation of the implanted rings permitted to exclude the

presence of heavy thrombotic matrixes. The uncoated metal rings showed a thin

thrombotic layer on their internal surfaces without a significative reduction of the

blood flow.

The Ti 6AI4V totally Carbofilm coated rings did not show any thrombotic depot on

the surface exposed to the blood flow.

The Ti 6AI4V partially coated rings showed the same results and more over the

scanning electron microscopy investigations did not evidence any thrombotic

starting between the coated and the uncoated surfaces.

A thin layer of adherent plasmatic proteins was detected, but no red blood cells or

platelets were identified in this protein matrix.

-c ,-E ...... -E -): C .... LI.

CI C C .... ID

-C

313

1200

• Ti6AI4V 1000

000

600 -e-- STllLlTE 25

400

200

O~---r--~----r---,----r---.---,.---~---r---.--~

o 20 40 60 00 100 120 140 160 100 200 220

TIME (min.)

Figure 2. Blood flow curves of uncoated Stellite and Titanium rings.

1200~---------------------------------------------,

'e 1000 ......

---a-- Ti6AI4V + CARBOFILn (parU.Uy co.ted) -E 000 -): o .....I LI..

CI o o .....I ID

600

400

200

.~

• Ti6A14V + CARBOFILM ( tot.ny coated)

O~---r--~----r---,----r---.---,r---~---r---.--~

o 20 40 60 00 100 120 140 160 100 200 220 TIME (min.)

Figure 3. Blood flow curves of totally and partially Carbofilm coated Ti 6AI4V

rings.

314

TABLE 2

Experimental results of Gott's rings in vive test.

METALLIC ALLOYS

Ti6AI4V

STELLITE 25

Ti6AI4V + CARBOFILM PARTlALL Y COA TED

PATENCY

DO DO DO

FLOW MODIFICATIONS

NEGLIGIBLE

NEGLIGIBLE

NONE

Ti6AI4V + CARBOFILM D 0 TOTALL Y COA TED NONE

CONCLUSIONS

The thromboresistance test with the Gott' s rings is considered one of the most

reliable and effective for biomaterials to be applied in manufacturing of

cardiovascular devices. The results obtained by Gott and others. using polymers

rings, with this kind of test were very often negative.

The alloys employed in this experiment have demonstrated to be not thrombogenic;

this is especially due to the highly sophisticated surface polishing and smoothing.

The Carbofilm (thin film of high density carbon) potentiates the haemocompatibility

characteristics af Ti 6AI4 V giving to the coated surfaces an increased

thromboresistance.

The SEM investigations put in evidence marked differences in composition of the

deposited organic matrix on the inner side of the rings.

These organic debris have not thrombotic origin when the surface is Carbofilm

coated.

The purpose to realize metal rings partially coated with Carbofilm is to put in

evidence possible phenomena of galvanic potential between the coated and the

315

uncoated surfaces. This carbon-metal couple could induce in principle a localized

thrombotic starting.

In the galvanic serie the carbon is placed between the noble metals in a sequence

in which the titanium is present. Some authors [1.2.31 reported that never

observed any pitting phenomena at the interface between carbon and meta!.

Not so clear is the probability of a thrombotic starting localized in areas in whose a

negligible galvanic potential are present like as in this case.

The experimental results obtained with this test. considered extremely severe

Ismall diameter of Gott's rings. implantation into a venous vessel and without a

pharmacological support). substantially confirm the high characteristics of

haemocompatibility of the Carbofilm coating.

In the same experimental conditions. also when the metal ring (Ti 6A14V) is only

partially coated with Carbofilm. the thrombotic phenomena are totally absent.

These results indicate that the electrochemical phenomena at the interface between

Carbofilm and Ti 6AI4V are not able to induce thrombosis of the surfaces.

This aspect is particularly important in devices in whose an interface of this type

could be present after wearing of Carbofilm coated surfaces.

REFERENCES

1. Tombrei. F .• Preparation et structure des pyrocarbones. In !..!l Carbones. Masson et C.ie Editeurs. 1965. pp. 783-838. .

2. Kaae. J.L.. Gulden. T.D .• Structure and mechanical properties of co-deposited pyrolytic C-SiC alloys . .L. Amer. Ceram. Soc .• 1971. 54 1121. 605-09.

3. Bokros. J.C .• Variation in the cristallinity of carbons deposited in fluidized beds. ~. 1965. 3. 201-11.

4. Shim. H.S .• The behaviour of isotropie pyrolytic carbons under cyclic loading. Biomat. Med. Dev. Art . .Q.r.Q... 1974.2111. 55-65.

5. Bard. R.J .• Baxman. H.R .• Bertino. J.P .• O'Rourke. J.A .• Pyrolytic carbons deposited in fluidized beds at 1200 to 1400 xC from various hydrocarbons. ~. 1968,6,603-616.

6. Paccagnella, A .• Majni. G .• Ottaviani, G., Arru, P., Santi. M .• Vallana. F .• A new pyrolytic carbon film for biomedical applications. In Ceramies in clinical aDDlications: Satellite Symposium 2f 1h§ VI World Congress 2!!. High il.!<h Ceramies. Milan, 1986.

7. Thompson. N.G., Buchanan. R.A .• Lemons, J.E., In-vitro corrosion of Ti6AI4V and type 3161 stainless steel when galvanically coupled with the carbon . .L. Biomed. ~ Bn.... 1979. 13. 35-40.

316

8. Rostoker, W., Pretzel, C.W., Galante, J.O., Couple corrosion among alloys for skeletal prostheses. ,L, Biomed. Mater. futL, 1974, 8, 407.

9. Miller, B.A.Jr., The galvanic corrosion of graphite epoxy composite materials coupled with alloys. Report AD-A019322, Wright-Patterson Air Force Base, 1975.

10. Costello, M., Stanczewski, B., Vriesman, P., Lucas, T., Srinivasan, 5., Sawyer, P.N., Correlation between electrochemical and antithrombogenic characteristics of polyelectrolyte materials. Trans. Amer. Soc. Artif. lr!l!!!!.... Qm.., 1970, XVI, 1-6.

11. Whalen, R.L., Jeffrey, D.L., Norman, J.C., A new method of in vive screening of thromboresistant biomaterials utilizing flow measurements. I.r.i.r!L Amer. Soc. Artif. Intern. Qm.., 1973, XIX, 19- 23.

12. Marconi, W., Bartoli, F., Mantovani, E., Pittalis, F., Settembri, L., Cord ova, C., Musca, A., Alessandri, C., Development of new antithrombogenic surfaces by employing platelet antiaggregating agents: Preparation and characterization. Trans. Amer. ~ Artif. Intern. Qm.., 1979, XXI, 280-85.

13. Baier, R., Gott, V.L., Feruse, A., Surface chemical evaluation of thromboresistant materials before and after venous implantation. Trans. Amer. Soc. Artif. Intern. Qm.., 1970, XVI, 50-7.

14. Kusserow, B., Larrow, R., Nichols, J., Observations concerning prosthesis­induced thromboembolic phenomena made with an in vive embolus test system. Trans. Amer. ~ Artif. lr!l!!!!.... Qm.., 1970, XVI, 58-62.

15. Kusserow, B., Larrow, R., Nichols, J., The surface bonded covalently crosslinked urokinase synthetic surface. In vitro and chronic in vive studies. I.r.i.nL Amer. fuu;.. Artif. !.nnm... Qm.., 1973, XIX, 8-12.

16. Cumming, R.D., PhilIips, P.A., Singh, P.I., Surface chemistry and blood­material interactions (BMI). Trans. Amer. ~ Artif. Intern. Qm.., 1983, XXIX, 163-68.

17. Bruck, S.O., Biomaterials in medical devices. Trans. Amer. Soc. Artif. Intern. Qm.., 1972, XVIII, 1-9.

317

TCP- Impurities in HA- Granules and

Crystallinity Changes in Plasmaflamesprayed HA- Coatings

Detected by Spedrosc:opical Methods and their Consequences.

M. Weinländer .. , .... , H. PlenkJr ..... , F. Adar""" and R. Holmes ........

Orofacial Implant Center, UCLA, Los Angeles, USA *

Sone & Biomaterial Research Lab., Histol.· Embryol. Inst Univ. Vienna, A.**

Instrument S.A., Edison, N.J., USA***

Dept. Plastic Surgery, UCSD, San Diego, USA***"

Abstract

Two different commercially available granular hydroxyapatite (HA) ceramic materials were placed in experimental bone defects.The bone re action was evaluated histologically and histomorphometrically. The newly formed bone has tight contacts with the more rounded granules cf one implant material. Predominantly at the sharp edges of the ether more polygonal shaped granular material no bone appostion and degradation cf HA granules via foreign body giarit cells (FBGC) and macrophages could be demonstrated. X-ray powderdiffractometry (PDM) detected tricalciumphosphate (TCP) impurities in these HA granules. In animal experiments two series of plasmaflamesprayed HA coated dental implants have been evaluated histologically and histomorphometrically.Since the same biodegradation was again detected at plasmaflamesprayed HA coatings on metallic dental implants, spectroscopical evaluation cf the thin ceramic coatings was attempted. Ramanspectra of three diferent commercially available HA coated dental implants have been measured and compared to the spectra cf crystalline and amorphous HA as weil as to the spectrum of TCP. All implant coatings showed spectra more comparable with the amorphous phase of HA than with any other reference materials examined. Persisting FBGC and macrophages predominantly at the sharp edges and the presence of biodegradeable TCP impurities in the polygonal shaped HA granular material leads to areas cf missing bone formation .and biodegradation. In experimental studies HA coated dental implants performed superior in terms of earlier and more extensive bone formation compared with two titaniumsurfaced implants. The biological consequences cf biodegradeable amorphous calciumphosphates for long term periimplant bone formation around plasmaflame sprayed HA coated dental implants is not clear at the moment.

Introduction

Calciumphosphate - ceramic (CPC) materials are used extensively in contemporary bone reconstructive surgery. The most widely investigated CPC are hydroxyapatite (HA) and ß - tricalciumphosphate (TCP). The biodegradation cf these materials can take place either through the dissolution in physiological

318

fluids or cell mediated phagocytosis. It is generally known that dense hydroxyapatite materials show Iittle er no resorption and that porous TCP is resorbed at a higher rate than HA with a similar structure. It was the goal cf this study to evaluate the biological response cf two different HA - granular materials and two different HA- coated dental implants following a spectroscopical evaluation of their chemical composition and purity.

Materials and Melhods

x- Ray Powderdiffractom Prior to the implantation, both granular ceramic materials (Type A, Type B) were evaluated in vitro using the X- ray powderdiffraction technique (PDM) to determine their crystal structure. For this purpose both materials. were pulverized to a grain size of 10J,Jm and subsequently evaluated with a X- ray powder diffractometer (Phillips PW 1710) at the wavelength cf CU -Q. The resulting peaks were identified by comparison with the ASTM indices. [1]

Materials For implantation two commercially available granular hydroxyapatite (HA) ceramic materials Type A * ,Type B** and two HA - plasmaflamespray coated dental implants (Implant Type I ,cylindrical)', (Implant Type 11 , screw -shaped) n were selected for this study . • Periograt (COOk& Walle), .. C8lcitile (C8lcitek Inc), , Integral (C8lcitek Inc). " Steri- oss (Denar Corp.)

Raman Speclroscopy Macro (100J,Jm) and microspectra (5J,Jm) cf the two implants ( Implant Typei, Implant Type 11) and previously tested [2] reference samples cf tricalciumphosphate (TCP) § and hydroxyapatite (HA) Q were aquired on the U 1000 Raman microprobe. [3] § Synthograft (Johnson & Johnson, Q AlVeograt( COOk- Walle)

Animal Experiments Three series cf animal experiments were conducted.

Series 1: In eight experimental rabbits a monocortico - medullary defect was created at

both femoral diaphyses. Post implantation cf Implant material Type A and Type B into these defects, the surgical f1ap was sutured in two layers with resorbable suture material. The rabbits were sacrificed at six, eight , ten and twelve weeks postoperatively.

Series 2: Seven mongrel ( average weight 16.2 kg) dogs were edentulated bilaterally

in their premolar mandibular region. After six months healing time one HA­coated Implant Type I was implanted in each cf the dogs right mandible next to two titaniumsurfaced implants, which served as control. After three months healing time seven HA- coated implants were removed with b1ocksections.

319

Additionally fourteen titaniumsurfaced implants were harvested as contro!.

Series 3: Two mongrel dogs (average weight 17.0 kg) were used to evaluate the HA­

coated Implant Type ". Four implants each were placed in both ulnae of each dog. The animals were sacrificed at eight and twelve weeks. Eight implants were harvested after eight weeks, and the additiona"y eight implants were harvested twelve weeks post implantation. A" experimental animals received a three step intravital fluorochrome labelling consisting of tetracyclin at seven days post operative, Alizarinkomplexon at twentyone days post operative and Calcein green two days before sacrifice. [4]

Histological Technique The specimens were prepared according to the guidelines decscribed by Plenk. [4] After embedding in polymethylmethacrylate, undecalcified sections were prepared with a heavy duty microtome or with the sectioning - grinding technique. For further technical details see reference [4]

Five representative microradiographs (MRG) from each granular ceramic material (Type A, Type B) as indicated in Figure 2A and 2B were evaluated with a semiautomatic morphometric system (lBAS®,Zeiss). Fo"owing parameters were quantified. 1) Implant granule perimeter; 2) Length of bone - implant contact zones. The results were expressed as the percentage of the implant granule perimeter having bone contact zones. The me ans of this ratios were calculateed for the two groups (Type A, Type B) and the significance of the differences obtained , evaluated by the Students t - test.

One midline section of each Implant (Implant Type I, Implant Type ") was coated with Ag/Pd and placed into a scanning electron microscope (Cambridge StereoScan 360 SEM). Backscattered images were obtained with gun voltage and probe current settings that permitted digitization and computerized recognition of soft tissue, bone and implant based on their atomic number differences. Typica"y at 40x magnification, 6 images were aquired along the implant interface, each sampling 2mm of length. These images captured by PC microcomputer, were then analyzed for bone apposition.

Results

x- Ray Powderdiffractometry The PDM evaluation of Implant material Type B demonstrates a pure HA diffraction pattern at 1.943A, 2.262 A, 2.631 A, 2.720 A, 2.778A, 2.841A and 3.440A. (Fig. 1B) The diffraction pattern for Implant material Type A however indicates at 2.607 A, 2.880A and 3.210A characteristical peaks for TCP. ( Fig.1A) The peak intensities indicate the amount of TCP in the range of 5%.

Histological Evaluation (Series 1: Implant material Type A, Type B) In subperiosteal and endosseous implant sites of the two granular ceramic materials new bone formation could be demonstrated around a" implant granules which came in contact with endosseus or subperiosteal osteogenic cells. Characteristica"y the newly formed woven bone has tight contacts with the

320

flat or rounded implantsurfaces of Type A and Type B and shows no interposed connective tissue layer. (Figure 2 A, 2B) Predominantly at the sharp edges of implant Type A (arrows) bony recesses can be demonstrated. Within these bone recesses multinucleated foreign body giant eells (FBGC) with incorporated HA­particles could be identified. The implant Type B which has a more rounded spherieal character, in contrast to the implant Type A, is not showing these bone reeesses. The new bone is forrned according to normal bone healing and reparative patterns and shows transformation into mature lamellar bone.

The histomorphometrical comparison between the Type A and TypeB implant material demonstrated a statistically significant (p< 0.01) higher mean of implant -bone contact for Type B (70.04 ± 57.78) versus Type A (44.99 ± 23.73). Implant granules whieh eame not in contact with osteogenie eells, respectively were separated trom the periosteum during the bone healing process, did not elieit any new bone formation. COllagenous fibers are arranged in layers around these implant granules and FBGC with round cell infiltrates ean be notified with the granules. This ehronie inflammatory reaction is more visible around the polygonal sharp edges of the implant Type A than around the rounded surfaees of the implant Type B.

Rarnan SpecIroscopy Initially it was thought that variations in the performance of plasmaflamesprayed HA coated dental implants might be due to the deposition of small amounts of TCP in plaee of some of the HA. Therefore complete spectra of the reference materials were recorded and examined to determine the best region of the spectrum to use for determination of phase. It was apparant that the phosphate stretching region near 965 cm -1 presents the most sensitive region for diagnostics of phase. The spectrum of the Implant Typel and Type 11 (Figure 3, arrows) is overlaid with that of the reference materials and one additional implant. All implants investigated gave more or less the same spectrum. The strong phosphate band of HA dominates the spectrum of the implant. However there is another band at lower frequency (Figure 3, asterix), somewhat broader than bands of the referenee material. This band does in fact overlap with a band of TCP, but is not being assigned to TCP for the following reasoo. The TCP band shows another band, of even greater intensity, on the high trequency side of the HA band whieh was absent in all implant spectra. Therefore the extra low frequency band had to be accompanied by a higher frequency band if its origin were TCP. An alternate expression for the extra band is that it comes trom amorphous material codeposited during the plasma spray process whieh is known to melt the partieles partially. It is known that the Raman spectra of the amorphous phases are analogous to the spectra of the crystalline phases of the same material. The principal differences are that the bands seem to be broadened and shifted to lower frequency.

Histological Results, (Series 2: Implant Type I ) All seven cylindrieal Implants ( Type I) gave the same histologieal picture. The bone formation (B) on the HA - coating (HA) is of a developing lamellar eharacter in a paralell direction to the surfaee of the implant. This paralell layer is approximately 100jJm thick and in elose contact with the irnplant surfaee and the

321

secondary osteons extending from the original bone bed towards the implant. (Fig. 4, B = bone, HA = hydroxyapatite coating, Ti = titanium, arrow indicates paralelllayer of bone) Even in areas with no direct implant - bone contact (e.g. medullary cavity) a thin layer of bone covers the implant surface. The thickness of this layer is approximatelay 36IJm, as measured with spot measurment in the back scatter electron (BSE ) image .. Occasionaly FBGC and macrophages with incorporated calciumphosphate particles can be seen at the implant - bone interface. ( Fig. 5, BSE image, arrows indicate desintegrating HA coating, B = bone, Ti = titanium)

The histomorphometrical evaluation of the percentage of bone contact with the HA - coated Implant Type I indicated 71.35 ± 11.79. The titaniumsurfaced controlls demonstrated 45.66 ± 16.42, and 54. 96 ± 10.85 respectively. (p < 0.01)

Histological Results (Series 3: Implant Type 11) The bone formation at the 8 week specimens of the screw shaped Implant (Type 11 ) imitated the one discussed above. As demonstrated with the tetracyclinlabelling implantofugal bone formation takes place directly at the implant surface and simultanously implantopetal bone growth can be demonstrated at the original bone bed. Alizarin red demonstrated concentric bone formation in secondary osteons . Whereas Calcein green labelling indicated further concentric closing of secondary osteons, there was no indication for continuing bone tum over at the implant - bone interface. The twelve week specimens of the Implant Type 11 gave a more' moth eaten • image of the implant - bone interface than seen with the eight week specimens. On close viewing of the implant - bone interface pronounced biodegradation of the HA - coating with the presence of FBGC and macrophages could be demonstrated.

Whereas the eight week sampies of the HA - coated implant Type 11 demonstrated histomorphometrically a mean percentage of 74% implant - bone contact, the twelve week specimens showed a decrease of the mean values to 50%.

Discussion

The histological and morphometrical evaluation of two granular calcium -phosphate materials of different shape in an endosseous and subperiosteal experimental defects are consistent with previous studies conceming the soft tissue response to different shapes of implanted HA ceramic granules. [4] The bony recesses which could be demonstrated in this study could indicate that a irregular shaped implant configuration leads to a persistent foreign body reaction and therefore prevents tight bone apposition to the irregular shaped implant surface. An enhancing factor for this process could be the pesence of biodegradeable TCP detected in the irregular shaped implant material. This fact could be demonstrated through the presence of macrophages with incorporated calciumphosphate particles in this study. The presence of similar biodegradation processes at the interface of plasmaflame sprayed HA - coated implants favoured the idea of having similar

322

TCP impurites within the plasmaflame sprayed coatings.Since the results of the Raman spectroscopy did not proof this theory, but biodegradation still takes place, one has to conclucle that the presence of amorphous calciumphosphate phases (ACP) in this coatings is the reason for the biodegradative process. The initial bone formation at HA coated implants is superior in terms of amount and maturity compared with titaniumsurfaced implants of a similar macroscopic design. Nevertheless the impact of biodegradeable ACP in plasmaflame sprayed coatings on long term implant stability has to be evaluated. One previous report indicates a continuing biodegradative process between the 12 and the 32 week, documented with decreasing push out values for the evaluated HA coated cylindrical implants. [5] The histomorphometry of the screw shaped HA coated implants in our stucly indicated a decrease in the amount of the implant - bone contacts from the 8 to the 12 week from 74% to 50%.

Conclusion

The histological and X- ray powderdiffraction evaluation of two granular ceramic materials of different shape indicated that it is more likely to get bone apposition to flat or rounded surfaces of a pure HA material ( Type B) than at the irregular shaped SUrface or sharp edges of a HA material (Type A) with a significant amount of tricalciumphosphate ( TCP) impurities. Not only TCP , but also amorphous calciumphosphates ( ACP) lead to biodegradation of calciumphosphatebiomaterials. Since no longterm results of the interface between loaded HA coated dental implants and surrounding bone are available, the initial encouraging experimental results of unloaded implants should be viewed critically. The role of ACP on HA coated endosseous dental implants deposited during the plasmaflame spraying process requires further investigations in terms of bioresorption and impact on biointegration.

References

[1] ASTM, F4 Comittee, Philadelphia, PA [2] Weinländer M., Plenk H.Jr, Halwax E., und Annemarie Nikiforov: In vitro Untersuchungen verschiedener Hydroxylapatitmaterialien. Z. Stomatol 85/7:399, 1988 [3] Adar, F.: • Developments of the Raman micrprobe: Instrumentation and applications, Microchem J. 38: 50 -79, 1988 [4] Plenk, H. Jr. : The microscopic evaluation of hard tissue implants. In: Williams D.F. (ed) Techniques of biocompatibility testing. CRC Press, Boca Raton, FI, 35-81, 1986 [5] Misiek, D.J., Kent J.K., Carr R.F. : Soft tissue responses to Hydroxylapatite Particles of different shapes. J Oral Maxillofac Surg 42: 150-160,1984 6] Cook,S.D., Kay, J.F., Thomas, K.A. ,Jarcho, M.: Interface Mechanics and Histology of Titanium and Hydroxylapatite- Coated Tittanium for dental Implant Applications, Int J Oral Maxillofac Implants 2, 15,1987

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326

LONG1ERM STABllJ'IY OF 11N

MAX SCHALDACH and ARMIN BOLZ Zentralinstitut für Biomedizinische Technik der Friedrich-Alexander-Universität

Erlangen-Nürnberg, Turnstraße 5, D-8520 Erlangen

ABS1RACT

Metallic materials which are used in orthopaedics, heart valves and pacemaker electrodes have to be biocompatible; i.e. they have to avoid the reaction with the surrounding tissue. Especially for long-term stability it is essential to inhibit chemical reactions. Titanium and its compounds like oxides and nitrides are wen known to behave biocompatible. Tberefore, electrodes with a porous surface were produced by sintering titanium alloy powder or PVD of TiN. Tbe electron conductivity can be further increased by Boron doping. To differentiate between threshold changes resulting from surface oxidation or from fibrotic layer, second­ary ion mass spectroscopy (SIMS) was perfonned. Depth profIles of the oxygen concentra­tion allow the evaluation of the electrode interface characteristics with regard to time dependent on the formation of oxide layers. Tbe results indicate that the high long-term efficiency of titanium compound pacing electrodes is a result of the interface structure, the high dielectric constants of the thin surface oxide layer detennines the potential and charge distribution resulting in an improved biocompatibility. Tberefore the low stimulation threshold is a good indicator of an improved tissue compatibility inhibiting the protein activation.

IN1RODUcnON

Tbe success of titanium and its alloys as biomaterial is indisputed. A number of reports have proven the superiority compared to other groups of metallic materials in regard to macro­scopic, microscopic and molecular aspects of the interface [1, 2]. Tbe structure and the electrochemical properties of the surface atomic layers of the implant and the first molecu­lar layers of the adsorbates are of essential influence towards biocompatibility. Tberefore the tight oxide layer of Ti prevents an exchange of electrons and thus inhibits currents across

327

the interface protecting the ionic equilibrium of the adjacent biological phase. However this fUm under implant conditions may be locally removed by a mechanical process such as wear resulting into high corrosion rates until the oxide layer is self-repaired [3]. This situation is relevant to all implants exposed to mechanical interaction to the adjacent tissue. Studies were initiated to investigate the potential improvements, that could be provided by surface modifications. Thermal nitration, physical vapor deposition (PVD) of TiN, and ion implantation provide a variety of surface modification techniques [4, 5, 6]. The electro­chemical and physical performance of surface coating was deterrnined by standard poten­tiostatic impedance measurements. In addition TiN-coated, as weIl as B-doped titanium stimulation electrodes have been implanted in the right ventricle and their stimulation threshold and intracardiac electrogram (lEG) postoperatively monitored by bidirectional telemetry with the pacemaker. The sensing and pacing performance of the coated elec­trode-surface provides directly information of the material-tissue interaction. As previously reported [7, 8, 9], the electrochemical properties of sintered and surface-treated electrodes prove the predicted improvement of sensing performance if titanium-coated electrodes are used. TiN as weIl as B-doped Ti-Electrode coatings in comparison to Pt, Pt-Ir-Alloys show lower stable thresholds. The results of sintered and nitrided electrodes TiA15Fe2.5 demon­strate the superiority to all other materials presently known. The advantages are due to the micro-crystalline surface structure achieved by sputter-deposited coatings and the kinetics of the ionic exchange. Furthermore, the acute thresholds achieved with the TiN-system are significantly better than with those of a smooth-metallic surface. These results are also confmned for chronic implants and are attributed to the known biocompatibility of titanium and its alloys. The electrical performance describes the simplified equivalent circuit consisting of a parallel circuit of the Helmholtz capacity and the Faraday impedance. If the double-Iayer capacity is large, the pacing losses will be low, while the improved fUter characteristics will result in a higher sensitivity of the depolarization signals. This behaviour explains, in general, the advantages of porous- over smooth-surface electrodes, which were reported earlier. Consequently, it was expected that the efficacy of the artificial stimulation may be improved by electrode surface coatings based on better biological compatibility. As a result, metals such as titanium, tantalum, niobium and zirconium and their compounds (oxides, carbides and nitrides) should have promising possibilities. The conductivity may be widely adjusted by changes in the stoichiometry and, therefore, titanium oxide and tantalum (V) oxide may also serve as coating or sinter material for stimulation electrodes. The same applies to electrode materials such as glassy carbon - its biocompatibility and microporous structure is very weIl known [10, 11].

SURFACE COATING

The influence of surface coating on the electrical, mechanical and physicochemical be­haviour of titanium was studied under in vivo and in vitro conditions. Electrodes of hemisperical shape were made either by PVD on cp-Ti-substrates following the Thomton­model [5] or by sintering powder of TiAI5Fe2.5-Alloy. The powderization was done by HDH, where the bulk material was processed pressurized with pure hydrogen at 1.6 bar at

328

750·C [8, 12]. The electrodes were sintered under Ar-athmosphere at 950·C. With an average porosity of 70% (grain size of 40-60 ~) a pore size of 50 ~ was obtained. The subsequent gas nitriding at 9OO·C over 70 W results in a TIN-Iayer of approximately 13 ~ thickness. Similar results were obtained with cp-Ti-spheres. Sputter-deposited thin films are formed by a material flux approaching the substrate from a limited set of directions. Consequently, the metallurgical grain tends to be columnar. Furthermore, at low ho­mologous temperatures (substrate temperature T, relative to the coating material's melting point: Tm), a growth structure defined by voided boundaries (also columnar) is superim­posed on the intrinsic grain structure. Thin-film microstructure tends to show an anisotropic character which affects current transport properties as weIl as the nature of barriers. In particular, the grain or growth boundaries present preferred diffusion paths often extending the entire layer thickness. For most applications of sputter-deposited layers as in electronic circuits, for wear resistance or decorative purpose, the formation of a columnar growth structure defined by voided boundaries is undesirable. Many efforts have been undertaken to avoid those structures and to support the formation of dense structures with closed boundaries. As aprerequisite for the use of porous TiN-coating for pacing electrodes, the understanding of the "undesired" voided growth defects have been fully utilized in order 10

enlarge the open microstructure of the surface, gaining a maximum value of double-layer capacity. The main factors to be considered during deposition are:

limited deposition angles given by sputter process set-up;

low mobility of absorbed atoms (low substrate temperature relative to coating material melting point);

atomic shadowing - elevations on the growing surface are preferred receiving more coating flux than valleys - inducing open boundary;

oblique deposition angles favoring open microstructure; suppression of intense energetic ion bombardment, leading to a low resputtering rate, again unfavorable to dense microstructure.

Figure 1. SEM micrographs of the TiN coating.

In fact, a balanced set to the above-mentioned process factors have been perfectly tailored to the properties required for the described TiN coating. Figure 1 shows SEM photographs of the TiN coating . The highly porous structure is clearly visible.

329

CHARACIERIZATION OF 1HE ELECIRODES

In order to understand the electrochemical properties of the interface electrode/myo­cardium, the potential and charge distribution can be explained by a simplified Gouy-Chap­man model: the double-Iayer capacity depends on the charge distribution within the phase-boundary. The capacity of the double layer in equilibrium is given by:

Cg= E E() A/d

with A = surface area, d = thickness, and E() = relative dielectricity constant of the dou'ble-Iayer. As soon as adsorption occurs on the surface, the description of the double­layer becomes rather complicated by the addition of serial capacitors. To enable a compari­son between various electrodes in clinical use, adsorption was not considered. Potentiostatic in vitro measurements were used to determine the electrical properties of the electrodes

........ 10' c: ......... pF'/cm' - 1 Tim ive. uncOO1od 10

2Tim ive. thermally nitrided 50 CI) 3Ti m ive. therma.\ly nitridod 250 u c 4 Ti inlered. une Iod 1000 co

inlered. thermal1y nitride<! 3000 "'0 '0'

STi CI)

6 Ti sinlered. thermally nitride<! 6000 0. E 7Ti intered. TIN·PV]).(:o:ued 20000 -

10' 10' 10' 10'

Fr quency [Hz]

Figure 2. Impedance of different electrodes.

using the equivalent circuit and experimental set-up as published earlier [9]. The in-vitro experiments were done in a 0.9% NaCI solution where the pacemaker electrode acts as

330

working electrode, a kalomel electrode as reference electrode and a glassy carbon electrode as counter electrode. By potentiostatic polarization between -200 and +200 mV the impedance was measured as function of the frequency and the capacity is calculated. Figure 2 shows the results of the impedance measurement as function of the frequency and the calculated capacities of diffe~nt thennal nitrided and PVD-coated porous electrode tips compared with non porous and non nitrided electrodes. The non porous massive electrode tips show as one can expect a lower capacity than the sintered porous electrode tips. A surface nitriding is able 10 decrease the impedance because of the better electrical conduc­tivity compared with the non nitrided electrodes where the surface consists of an oxide layer with a low conductivity. Due 10 a larger surface even a non nitrided porous electrode had an increased capacity with 1 mF cm-2 compared with a massive, but surface nitrided electrode (250 ~ cm-~. Due 10 a stem-like columnar microstructure of the TiN layer the PVD-coated porous electrode shows the highest capacity of 20 mFcm-2.

CORROSION BEHAVIOUR OF TITANIUM NITRIDE

When TiN is exposed 10 physiologic NaCL-solution a thin passivating layer of titanium oxynitride is formed. Tbe chemical analysis of this layer was performed with secondary ion mass spectroscopy using 5 keV argon ions (ion source IQE 38/100 from Leybold) and a quadrupol mass analyzer (QAS 511, Leybold). Tbe mass spectra of positive as well as negative ions show titanium, nitrogen and oxygen; beside this A12+ , Na + ,K+ , F and cr ions were found in the topmost atomic layers. In spite of the high intensities the concentra­tion of these impurities is below 1 % due 10 their high ionisation coefficient Tbe oxidation kinetics was examined with SIMS 02--depth proftles after different exposure times of the sarnples [13]. Tbe results indicate an increase of the oxide thickness during the first days followed by a stable equilibrium. Tbis is in agreement with the corrosion kinetics of pure titanium and titanium nitride by measuring the corrosion current of these electrodes against a glassy carbon counter electrode, which is caused by the diffusion of hydroxide ions 10 the interface between the oxide layer and the bulk and the subsequent oxidation reaction. Tbe growing oxide layer acts as a diffusion barrier for the hydroxide ions leading to a decay of the corrosion current. Tbe time dependence of this reaction is described by

Cl C2 f I(D = Q(T) - Q(T)+ C3 with Q(T) = :(t)dt

where T is the experiment time, I(T) the corrosion current, Q(T) the cummulated charge and Cl, C2 and C3 are constants describing the geometry and the transport kinetics of the reactants. On the one hand the actual corrosion current represents a measure for the actual growth velocity of the oxide layer. On the other hand the cummulated charge is a measure for the overall oxidized ions and therefore for the oxide layer thickness. Figure 3 compares the results of cp-Ti electrodes with those of TiN showing a similar oxidation behaviour; on the right side (for low oxide thicknesses) the corrosion current is high, because the diffusion barrier is thin. This high current gives rise to an increase in charge (inverse to the x-axis)

331

cOITesponding 10 an increase in oxide layer thickness. As a result the corrosion current decreases indicating the passivating properties of the oxide layer. However, whereas the titanium electrode is showing remarkable inhomogeneties in the growth leinetics (denoted by the arrows this indicates spontaneous cracleing in the oxide layer) titanium nitride show

Corrosion current IIlAI 1.----------------------------------------------,

0,8

0,6

0,4

0,2

titaniuryl nitri<!e

titanium

0,1 0,2 0,3 0,4 0,5 0,6 0,7 0,8 0,9 lICharge 11IIlAsI

Figure 3. Corrosion current versus cummulated charge for TiN coated and uncoated electrodes.

a smooth decrease according 10 the equation. This proves that oxide layers on titanium nitride are more stable in comparison 10 titanium, especially do not tend 10 produce cracks or pinholes. As a result TiN demonstrates better passivating properties and therefore a higher biocompatibility.

CLINICAL RESULTS AND DISCUSSION

Clinical results were reported on 245 unipolar TiN PVD coated electrodes, 35 tennal nitrided sintered and 20 Ti-B-doped electrodes implanted into the ventricle via the cephalic, extern al jugular or subclavian vein. A quadrifIlar coil, made of cobalt-chromium­nickel alloy MP35N and insulated with silicone rubber, serves as conductor providing a lead with high flexibility. As part of the study, handling was foHowed a strict protocol and the electrode characteristics were evaluated and thresholds as weH as intracardiac potentials measured, using the pacing system analyzer, BIOTRONIK ERA 20. The mean R-wave

332

potential was 14.4 mV at a standard deviation (SD) of 4.23 mV (Figure 4a). The mean threshold voltage as 0.5 ms pulse duration was 0.24 V at a standard deviation of 0.08 mV. The mean value for 0.1 ms pulse duration was 0.66 V (SD 0.21 V) and for 0.3 ms pulse duration 0.36 V (SD 0.11 V). The Chronaxie-Rheobase relation behaved normal. Nonin­vasive postoperative measurements so fare were performed only on TiN-coated electrodes; they confmn the long-term stability and the excellent sensing and threshold performance (Figure 4b).

QWlnUty Ouantlty

% %

50 50

.. 0 40

30 30

20 20

10 10

0 0 5 10 15 20 mV 0,1 0,2 0,3 0," 0.5 V

R-potentlal Pul .. amplitude

F i i I 2 3 • 5 23. 5

wooks post implant monlhs post mplant

Figure 4. Distribution of measured intracardiac signal amplitudes and thresholds ( (A) intraoperative, (B) postoperative measurements).

The balanced double layer at the electrode/tissue interface determines the performance characteristics of the pacing electrodes. Longterm observations confmne that the pacing and sensing functions reflecting the biocompatibility and the large capacity of the double

333

layer. As exemplified by TiN, selection of the appropriate materials as weIl as the applied deposition processes, the electrode-tip configuration results in a significantly improved signal-to-noise ratio and reduced threshold, thereby permitting extended service life. Clinical results obtained with sputter-deposited electrode coatings prove the importance of the microstructure not only for the elctrochemical efficiency of the paced-sensing system but also give evidence of a better understanding of the biocompatibility of Ti and its compounds. In contrast to other applications, columnar growth structure defined by voided boundaries is of particular significance. The results also confmn the experimental inves­tigations, which served as a basis for the expectation, that further advancements in the electrical therapy of rhythmic disturbances can be achieved. The goal is, in particular, to provide a substantial prerequisite for the automatic control of output voltage and sensitivity setting based on the detection of the evoked potentials.

REFERENCES

[1] Albrektsson, T., Bdnemark, P.J., Hansson, H.A., The interface zone of inorganic implants in vivo: Titanium implants in bone. Ann of Biomed En~ 1983,11,1.

[2] Breme, J., Titanium and Titanium Alloys, biomaterials of preference. Proc Sjxth World Conf on Titanjum, France, 1988

[3] Zwicker, U., Etzold, U., Moser, Th., Abrasive properties of oxide layers on TiAl5Fe2.5 in contact with high density polyethelene. In T'84 Science and Technology , eds. G. Lütjering, U. Zwicker, W. Bunk, Proc of the 5 Int Conf on Tjtanjum, Munich, 1985.

[4] Brophy, J.H., Rose, R.M., Wulff, J., The structure and properties of materials. In Thennodynamjes of Strueture, Vol. 2, Wiley, New York, 1964,82.

[5] Thomton, J.A., The microstructure of sputter-deposited coatings. J. Yac. SeL Tecb­IWl.., 1986, A4(6), 3059-3065.

[6] Münz, W.D., Hofmann, D., Herstellung harter dekorativer goldfarbener Titannitrid­schichten mittels Hochleistungskatodenzerstäubung. MetaUoberfJäehe ,1983,7,279-285.

[7] Sehaldaeh, M., New Pacemaker Electrodes. Trans Amer Soe ArtjfjcjaJ Int Or~ans 1971,17,29.

[8] Breme, J. and Schaldach, M., Porous Heart Pacemaker Electrodes of TiA15Fe2.5 Alloy. In Proe of Int Conf on Iitanjum Produets and AppUeations, Orlando, 1990, 604-611

[9] Schaldach, M., Hubmann, M., Weikl, A. and Hardt, R., SputtepfC~osited TiN Electrode Coatings for Superior Sensing and Pacing Performance. , 1990,13, 1891-1895.

[10] Shigemitsu, T., Matsumoto, G., Tsukahara, S., Electrical properties of glassy-carbon electrodes. Med mol En~ Camp, 1979,17,465-470.

[11] Mund, K., Electrochemical_ Pn>~es of Platinum Glassy-Carbon and Phyrographite as Stimulating Electrodes. fMJi, 1986,9,1225.

[12] Schaldach, M., Bolz, A. and Breme, J., Porous Heart Pacemaker Electrodes of TiAl5Fe2.5 and TiN. Proc. 1st Eum.pean Con! on Bjomed En~., Nice, 1990,312-313.

[13] Bolz, A., Matlok, H., Still, M., Schaldach, M., Hubmann, M. and Hardt, R., Langzeits­tabilität des Detekions- und Reizschwellenverhaltens von Titannitrid-Herzschrittma­cher-Elektroden. Bjomedjzjnjsche Technjk,35, (Ergänzungsband), 1990, 131-133.

334

IN VITRO TOXICITY OF FINE PARTICLES OF HYDROXYAPATITE.

E.I. EVANS and E.M.H. CLARKE-SMITH Department of Anatomy

University of Wales College of Cardiff PO Box 900, Cardiff CFl 3YF, UK.

ABSTRACT

Although experimental and clinica1 studies have indicated that hydroxyapatite is a well tolerated material, in vitro studies using cultured fibroblastic cells and fme powders of hydroxyapatite have shown toxicity with both ceIl death and an inhibition of mitosis. The cell damage occurs with fine particles of different makes of hydroxyapatite and requires direct contact between cells and particles. Damage does not occur If the particles are separated from the ce1ls by a microporous membrane. The damaged cells release LDH suggesting some form of membrane damage and microscopy shows particles adhering to the cell membrane. There is no evidence to suggest toxicity of larger granules of hydroxyapatite intended for clinica1 use, nor is anything yet known of the importance of this mechanism of cell damage in vivo.

INTRODUCTION

Hydroxyapatite is considered, both experimentally and clinica1ly to be a well tolerated

material [l ,2]. Any ions released by breakdown or degradation of hydroxyapatite are

expected to be normal constituents of the cell environment and therefore harmless.

However, we have recently shown that fine particles of hydroxyapatite damage fibroblastic

cells in vitro, while larger particles do not [3]. These observations support the suggestion,

raised previously for metallic particles [4,5], that cell damage may be caused by direct

contact with fine particulate materials.

The work described here attempts to investigate, in a variety of different experiments, the

mechanism by which this cell damage occurs. Cells have been grown in direct contact with

hydroxyapatite powders or separated from them by a microporous membrane. Release of

LDH from cells exposed to powders has been measured and microscopic observations of

the ceIls have been made.

335

MATERIALS AND METHOnS

Fibroblastic cells from rats were grown using standard tissue culture techniques [3] in

Eagle's MEM medium (Flow Labs) supplemented with 10% calf serum and penicillin and

streptomycin. An atmosphere of 5 % C02 in air was used.

Three types of hydroxyapatite were used. One (Biomatrix D.G.) was a type

supplied for clinical use as a bone graft substitute: its manufacture involved sintering at

13OO°C. It was ground by hand in a glass pestle and mortar to produce fine particles with

a mean size of approximately 3.5", with 90% of particles between 0.5", and 8.5",. The

second type was prepared for experimental work in particle sizes of 15", and 100",. The

third was a commercial product (Biogel) supplied as a fine powder. All types were

sterilised by autoclaving before use.

Tests were carried out with cells grown in Petri dishes or Millcell-HA inserts

(Millipore Ud) which are small plastic cylinders with a microporous membrane (0.45",

pore size) sealing the lower end. The Millicells were placed in multiweIl plates containing

additional culture medium outside the membrane. For each experiment cells were

inoculated into the culture vessel and 24 or 48 hours later the hydroxyapatite was added.

The number of cells was chosen to give the same density of cells per unit area in all

experiments. Concentrations of hydroxyapatite are expressed in milligrams per 104 cells

originally inoculated.

Cell counts were made by trypsinising the cells and counting them in a

haemocytometer. LDH content of the medium in contact with cells was estimated from the

rate of conversion of pyruvate to lactate in the presence of NADH whose conversion to

NAD was measured spectrophotometrically using a commercial assay kit (Sigma

Diagnostics). Cells for SEM were prepared by standard techniques, leaving the cells

attached to the surface on which they were growing.

RESULTS

The growth of cells in Millicels either in contact with ("inside") hydroxyapatite powder, or

separated from it by a microporous membrane ("outside") is shown in figures 1, 2, 3 and

4. In each case the cells in contact with powder ("inside") show reduced growth compared

to those separated from the powder by a microporous membrane ("outside"). The reduction

of cell growth occurs in all cases, but is greatest with the finely ground powder and marked

with 15", powder.

336

CEUS .1000 80

60

40

20

O L---~-=~~--~---+--------~--~ o 2 l 5 6 7

DAYS

COIHROL -+- INSIDE OUTSIDE

Figure 1. Cell numbers (mean and standard error) at different times after exposure to 0.5mg fine1y ground Biomatrix hydroxyapatite.

CELLS x1000

40 \ _ CONTROL ~OUTSIOE U INSIOE

32.6

30

20

3 6

DAYS

Figure 2. Mean cell counts at 3 and 6 days after exposure to O.5mg hydroxyapatite of 15J.' diameter.

* = p < 0.001 compared to control.

337

CELLS x1000 14

12 10.7 10.9 9 .6

10

8

6

:f 3 .5 *

0 6 4

DAYS

CONTROL B OUTSIDE o INSIDE

Figure 3. Mean cell counts at 3 and 6 days after exposure to 0.5mg hydroxyapatite of 1001' diameter.

CELLS xl000 40

* = p < 0.001 compared to control.

_ CONTROL r.:::Il OUTSIDE D INSIDE 33 .5

20

10 7.4 6 .8

o 3 6

DAYS

Figure 4. Mean cell counts at 3 and 6 days after exposure to 0.5mg Biogel fine hydroxyapatite.

* = p < 0.01 compared to control.

For comparison, the effect of fine powders of two glasses, known to leach toxic

components into the medium, is shown in figure 5. Here the growth of cells not in contact

338

with powder ("outside") is also reduced, hut the reduction is not so marked as for cells in

contact with the glasses.

CELLS x1000 12

10

8

8 6

2

O L-._-

_ CONTROL

_ G2 OUTSIDE

DAYS

GI OUTSIDE

c:J G2 INSIDE

9

6

o GI INSIDE

Figure 5. Mean ceH counts of ceHs exposed to 0.5mg/104 ceHs of two glasses, partiele size 3-40JL.

G1 = G338; G2 = Mp4.

LDH in the medium at various times for ceHs grown in Petri dishes in contact with hydroxyapatite is shown in figure 6.

LOH X100. unita 500

400 CONTROL

300

200

100

TEST

/

OL---__ ~ __ ~ __ ~~~~~ ______ L_ __ ~~~~~~

1 10 100 HOURS (log acalel

Figure 6. LDH levels in medium for test and control cells at times after addition of O.2mg finely ground Biomatrix hydroxyapatite per 104 ceHs

Scanning electron micrographs show partieles of hydroxyapatite in elose contact

with the ceHs, and the fact that these partieles remain after preparation of the specimen

339

suggests that they are adherent to the cell membrane. Phase contrast rnicrographs of living

cens were taken at various times after addition of fine hydroxyapatite partieles. After half

an hour the powder is still scattered fairly evenly over the eulture surface. It apparently

sticks to the cens so that at two hours most of the powder is concentrated around the

elumps of cens and by 6 hours all the particles are on or around the cens. Tbe particles do

not spread freely over the eulture surface again until most of the cells are dead after 4 or 6

days.

DISCUSSION

Tbese results show that cell damage by fine particles of hydroxyapatite requires direct

contact between cells and partieles. Tbe particles appear, by mieroscopy, to adhere to the

cens, and release of LDH suggests damage to the cell membrane.

Tbe test with bioglasses shows cen damage when the ceHs and powder are separated

by the rnieroporous membrane but inereased toxieity with direct contaet. Tbus the method

is ahle to distinguish two mechanisms of ceH damage, one requiring direct contaet with

cells and the other caused, presumably, by release of toxie ions.

Tbere is no evidence to suggest toxieity of the larger granules of hydroxyapatite

intended for elinical use, nor is anything yet known of the importance of this physical

mechanism of ceH damage in vivo.

ACKNOWLEDGEMENTS

We are grateful to Mr I A Brook, University of Sheffield and to Orthodesign Ud who supplied the sampies of hydroxyapatite.

Tbe work was supported by a grant from the Wellcome Trust.

REFERENCES

1. Holmes, R.E., Bueholz, R.W. & Mooney, V., Porous hydroxyapatite as a bone-graft substitute in metaphyseal defects - a histometrie study. I. Bone Int. Surg., 1986, 68, 904-911.

2. Chao, S-Y. & Poon, C-K., Histologie study oftissue response to implanted hydroxyapatite in two patients. I. Oral Maxillofacial Surg., 1987,45,359-362.

3. Evans, E.I., Toxieity of hydroxyapatite in vitro: the effect of particle size. Biomaterials, 1991, (in press).

4. Evans, E.I. & Benjamin, M., Tbe effect of grinding conditions on the toxieity of cobalt-ehrome-molybdenum particles in vitro. Biomaterials, 1987,8, 377-384.

5. Rae, T., Tbe haemolytie action ofpartieulate metals (Cd, Cr, Co, Fe, Mo, Ni, Ta, Ti, Zn, Co-Cr alloy). I. Pathol., 1978, 125,81-89.

340

DIFFERENCES IN BEHA VIOUR OF CULTURED FETAL RAT OSTEOBLASTS UPON BIOGLASS & NONREACTIVE GLASSES.

W.C.A. Vrouwenvelder; C.G. Groot; K. de Groot. Biomaterials Research Group, School of Medicine, Rijnsburgerweg 10, 2333 AA LEIDEN, The Netherlands.

ABSTRACf

Fetal rat osteoblasts (OB) were cultured upon bioglass (BG.45S5) & nonreactive glasses (NRG) for several periods and their behaviour was examined morphologically as weil as biochemically. Thefollowing parameters were used to demonstrate dijferences in OB-behaviour: cell morphology; presence and distribution of collagen type I & 11 and an OB-specijic membrane antigen; DNA-content. total alkaline phosphatase activity (APA) and the distribution of APA. Bioglass cultures demonstrated a better OB-like morphology, a higher proliferation rate and generally a better osteoblast expression in comparison with nonreactive glass cultures.

INTRODUCI10N

Many in vivo studies [1] have been performed demonstrating the bonding-to-bone capability

of bioglasses due to the formation of the bioactive layer [2] at the interface.

It was shown that the best bone-bonding properties were obtained with the bioglass encoded

45S5 containing 45 % Sial, 24.5 % N~O, 24.5 % CaO and 6 % PlO, (in weight %). The

interface reactions between bone tissue and bioactive implant; however, are still not fully

understood. As osteoblasts (OB) are responsible for bone formation at the interface we

investigated the behaviour of OB in elose contact with the bioactive substrate in an in vitro

model, with the elimination of unwanted in vivo complexities.

The purpose of this study was to examine the behaviour of fetal rat osteoblasts [3, 4] cultured

upon bioglass (45S5) compared with three nonreactive glasses (NRG).

341

MATERIALS & METHODS

Preparation of glasses: The bioglass 45S5 was prepared from high grade chemicals. After

mixing, the basic compounds were preheated. The mixture was allowed to react for 2-3 h at

1350 °C and then cast in preheated graphite moulds. The rods were annealed at 500 °C for

6 h then cooled down slowly for 10 h. The rods were sawn transversely and the resulting slides

were grinded and polished. Three nonreactive glass types were chosen: (1) a soda-silica-lime

glass (NG) with 72 % Si02, 16 % N~O and 12 % CaO, (in weight %); (2) commercially

available quartz glass (QG) with purity, > 99.9 % ; (3) microscopical coverslips (DG). Except

for the commercially manufactured coverslips (with a larger diameter of 10 mm), we obtained

thin, translucent slides with a diameter of 8 mm and thickness of 0.4 mm. The corrosion

behaviour of the glass types was investigated with routine scanning electron microscopy (SEM)

combined with a X-ray micro analysis.

Cell Isolation & Culture procedure: Osteoblasts were isolated from calvaria of 20-day-old fetal

rats under aseptic conditions by collagenase treatment [(1 mg/mI), 1 x 10': pre-incubation,

2 x 30': harvesting]. The obtained OB-enriched cell population was suspended in a-MEM,

containing 5 % fetal calf serum, 1 mg/mI glucose and 90 Jlg/ml gentamycin. Glass sampies

were placed in 24 weIl culture dishes. We applied 50 JlI of cell suspension to the glass sampies.

The cells were allowed to attach for 2 h, then 800 JlI a-MEM was added. The cultures were

kept in a humidified incubator with 5 % CO2 and maintained for different periods of time.

Preparations: Some cultures were fixed with 2.5 % glutaraldehyde in 0.1 M Na~cacodylate

buffer after 2, 6 and 12 days and prepared for examination with SEM. After 12 days part of

the cultures were prepared for immunocytochemistry. They were examined for the presence

of collagen type I & 11 (acetone fixation) and an OB-specific membrane antigen (fixation with

1 % paraformaldehyde in cacodylate buffer). The cultures were stored at -20°C after 3,6 and

8 days and collected to determine the DNA-content and Total alkaline phosphatase activity

(APA) biochemically. To demonstrate the distribution of the APA some cultures were fixed

with 10 % neutral formalin after 3, 6 and 8 days and stained histochemically.

RESULTS

Osteoblast-momhology: OB cultured upon BG for 2 days are compact with many dorsal ruffles

and filapodia (Figure 1). The toplayer is corroded. [We confirmed by X-ray micro analysis that

an increasing calciumphosphate enriched silica layer appeared on top of the surface of the BG

during exposure to the culture medium for periods up to 12 days.]

OB cultured upon NRG for 2 days are flattened with almost no dorsal ruffles or filapodia

(Figure 2). No surface corrosion is observed for any of the NRG.

342

Figures 1 and 2. Scanning electron micrographs showing osteoblasts (OB) cultured for 2 days upon bioglass (BG) and quam. glass (QG). (Bar = 0.1 mm). Figure 1: OB on BG have a compact structure with dorsal ruffles and jilapodia (arrows). Note the corroded toplayer. Figure 2: OB on QG are jIaltened and lack dorsal rujJles and jilapodia. No corrosion of toplayer is observed.

After 12 days confluent BG-cultures show large clusters of cells on top of the existing

monolayer. At higher magnification it can be observed that the clustered cells have many dorsal

ruffles at their surface and collagenous fibers in the extracellular matrix (ECM). On intact

NRG-cultures smaller clusters on top of the monolayer are observed. The clustered cells have,

besides dorsal ruffles, numerous blebs on their membranes, indicating a poor cell condition.

Collagenous fibers are also observed in the ECM. In general a lot of NRG-cultures detach

spontaneously from their substrate after reaching confluency, which is not found for the

confluent BG-cultures. In our model collagen type I is used as a marker for osteoblast

expression and is found in all cultures after 12 days. In BG-cultures the GY-collagen type I

antibodies are located in the cytoplasm and the ECM of the celis showing a dense fluorescence

over the monolayer with the clusters accentuated. In NRG-cultures the GY-collagen type I

antibodies are clearly observed in the cytoplasm and the ECM because of the more flattened

morphology of the osteoblasts upon NRG. Collagen type 11, a marker for chondrocytes, is

hardly produced in either of the cultures. The monoclonal antibody directed against a

membrane associated antigen of rat osteoblasts is used as an exclusive marker for osteoblast

expression. In BG-cultures the GY-OB antibodies are generally distributed in the monolayer but

stronger in the clusters showing polygonally shaped cells. In intact NRG-cultures the GY-OB

antibodies are mainly located in the clusters showing polygonally as well as spherically shaped

cells. The DNA-content and Total APA increase for all cultures after 3, 6 and 8 days (Figures

3 and 4). The APA is a marker for osteoblast-like differentiation. The DNA-content and the

APA are significantly higher for BG-cultures in comparison with the NRG-cultures after 6 and

8 days (p < 0.001, n = 6).

343

8 DNA-content (ug) 3 2 1 APA (nmolImin) 4

6 14

4

1 2

o o 3 6 8 3 6 8

Time (days) Time (days)

BG D NG

Figures 3 and 4. Histograms ofthe DNA-content and APA after 3,6 and 8 days of culture. DG* is the calculated value (because of the larger area for DG).

The ratio ofthe APA and DNA-content (the mean APA per cell) between BG-cultures and

NRG-cultures is not significantly different after 6 and 8 days of culture. Histochemical staining,

however, shows a more evenly distributed APA in the BG-cultures after 3, 6 and 8 days in

comparison with the NRG-cultures. In confluent BG-cultures ( ~ 6 days ) the APA is more

evenly distributed over the monolayer with higher activity in the large clusters, whereas in

NRG-cultures the APA is mainly concentrated in the smaller clusters with a relative high

activity .

DISCUSSION

We investigated the behaviour of fetal rat osteoblasts [3,4) cultured upon BG. Nonreactive

glasses were used as a control. We showed differences in behaviour and expression of

osteoblasts cultured upon BG and NRG by means of the above mentioned parameters.

The spreading of OB is influenced by the nature of the underlying substrate. On highly

negatively charged surfaces (such as the calciumphosphate enriched silica gel layer on top of

the BG) OB adopt a 'stand-off morphology [5) with many dorsal ruffles and filapodia. More

compact cells pro area are seen on BG after 2 days resulting in a denser cell layer. On neutral

344

surfaces OB are flattened and show aImost no dorsal ruffles. In the denser cell layer upon BG collagen type I is more concentrated in the cytoplasm and

the ECM of the cells. On NRG the cells are more flattened showing collagen type I as

collagenous fibers in the ECM. In BG-cultures the OB-associated antigen is more evenly

distributed over the monolayer. This antigen is mainly concentrated in the clusters of the NRG­

cultures, where possibly a favourable micro-environment is created for the osteoblast

expression. The spherical shaped cells also indicate a poor cell condition in the clusters of the

NRG-cultures.

A significantly higher DNA-content is observed for BG in comparison with NRG after 6 and

8 days of culture indicating a generally higher proliferation rate. The mean APA per ceIl,

however, is not significantly different. Visualization of the AP A establishes our hypothesis that

the OB expression on BG is more evenly distributed due to the stimulation of the appearing

bioactive layer.

CONCLUSION

Osteoblasts show a better osteoblast behaviour when cultured upon bioglass due to the

contribution of the appearing bioactive layer. The higher proliferation rate leads to a higher cell

density during culture and a generally better osteoblast expression in comparison with

nonreactive glasses.

References

1. Hench L.L.; Splinter R.J.; Allen W.C.; Greenlee T.K.; 'Bonding mechanisms at the

interface of ceramic prosthetic materials'.

J. Biom. Mat. Res. Symp., 1971,2(1), pp. 117-141.

2. Sanders D.M.; Hench L.L.; 'Mechanisms of glass corrosion and durability'.

J. Am. Cer Soc., 1973,56(7), pp. 373-377.

3. Matsuda T.; Davies J.E.; 'The in vitro response of osteoblasts to bioactive glass'.

J. Biomat. , 1987, 8, pp. 275-284.

4. Cohn D.V.; Wong G.L.; 'Isolated bone cells'. From: Skeletal Research

(Simmons D.J. & Kunin A.S., eds.), Acad. Press NY, 1979, pp. 3-20.

5. Maroudas N.G.; 'Adhesion and spreading of eells on charged surfaces'.

J. Theor. Biolo&y, 1975,49, pp. 417-424.

345

COMPLEMENT ACTIV ATION BY CERAMICS

ISABELLE DION*&****, ANIE BAQUEY**, CHARLES BAQUEY*, TIllERRY MESANA***, MONIQUE POURTEIN **, BERNARD CANDELON****,

JEAN-RAOUL MONTIES***

* INSERM U.306- Universite de Bordeaux 11, 146, rue Uo Saignat, F33076 Bordeaux ** Service d'Immunologie, Höpital Pellegrin, Place Amelie Raba Lean, F33076 Bordeaux *** Laboratoire de Recherches Chirurgicales, Universite Aix-Marseille 11, 27 Boulevard

Jean Moulin, F13385 Marseille-Cedex 5 **** TERACOR, ZI Toulon-Est, BP 238, F83089 Toulon Cedex.

ABSTRACT

The left ventricular assist device is based on the principle of the MaiIlard-Wenkel rotative pump. The materials which make up the pump must present particular mechanical, tribolo­gical, thermal and chemical properties. Graphite has been selected as a substrate, ceramic as the coating.The purpose of this study is to evaluate the in vitro complement activation eventually initiated by ceramic powders (AI2O:3, 'hO.fY20:3, AlN, B4C, BN, SiC, Si3N4, TiB2, TiC, TiN), graphite and diamond. The morphology of powders has been studied by Scanning Electron Microscopy and X-Ray Diffraction investigations have been carried out. The evaluation of complement activation was undertaken by measuring the total hemolytic complement (CHSO). No activation has been detected.

INTRODUCTION

The left ventricular assist device we are developping is based on the principle of the MaiIlard-Wenkel rotative compressor [1].

Wehave defined stringent requirements for construction materials. Coated materials seem to be the most suited to these requirements. The choice of the substrate depends on bulk characteristics and that of the coating on surface properties such as chemical, biolo­gical and mechanical behavior as weIl as potential interactions with the organism [2].

Parallel to this particular application, ceramic coatings as weIl as bulk ceramics are expected to be more and more used in the field of biomedical devices whether the latter are in contact with blood or not.

346

Tbe purpose of this stody is to evaluate the in vitro complement activation eventoally initiated by ceramic powders which have been stodied by Scanning Electron Microscopy and X-Ray Diffraction.

MATERIALS AND METHODS

Ceramic powders Alumina (A~0:3), Zirconium oxideIYttrium oxide ('bOfY20:3) and Silicon Carbide (SiC) were supplied by LONZA France - MARTINSWERK GmbH (10-12 rue des Trois-Fontanot F9200 Nanterre). Aluminium Nitride (AlN), Boron Carbide (B4C)' Boron Nitride (BN) and Titanium Diboride (TiB:z> were provided by COMAIP-ESK (68, Avenue G. Bizot F75012 Paris), Silicium Nitride (Si3N4) by METABAP-HCST (17, rue Eugene Delacroix, F75116 Paris). Titanium Carbide (TiC) and Titanium Nitride (TiN) were purchased from CEREX (24A, rue de la Resistance, F74108 Annemasse). Diamond powder was elaborated by DE BEERS INDUSTRlAL DIAMOND DNISION-ESKENAZI SA (24, rue Joseph Girard, CH-1227 Geneve) while graphite powderwas supplied by SUPERIOR GRAPHITE Co (120 South Riverside Plaza, Chicago, n. 60606).

Scanning Electron Microscopy Tbe SEM used here was an HIT ACHI S 2500. All powders have been metallized by gold. Tbe magnification varies from 50 to 50 000, at 15 or 20 kV, depending on the nature of the powders.

X-Ray DitTraction Analyses Tbe X-Ray Diffractometer was a PHll..IPS PW 1050 Spectrogoniometer (Cu KaI' A. = 1,5405). Each powder was deposited on an Aluminium plate.

Complement Activation Tbe evaluation of complement activation was undertaken by measuring the total hemolytic complement (CH 50) which is a parameter representing the hemolysis ability. Each powder was placed in a glass test tobe, the weight of powder was calculated with respect of the sur­face area/volume of serum ratio standardized to 500 cm2/ml [3] which allows a correct detection. Frozen sera sampies were thawed at room temperature and mixed prior to use. One ml of the sera pool was placed in each test tobe and incubation went for 60 minutes at 37OC. A positive control was obtained with a mixture of serum and Zymosan A, a known potent activator (S.Cerevisiae, Sigma, St-Louis, Mo, USA), with the respect of 500 cm2/ml. A negative control was obtained with serum incubated without any powder. After centrifu­gation at 1800 g for 10 minutes, the CH50 test was assayed on each serum (4). Tbe hemo­lysis ability is expressed by each serum towards sheep red cells sensitized by antisheep red cells rabbit antiserum, and quantified by the determination of the fraction of a constant globular pool which is hemolyzed in the presence of decreasing dilutions of each serum.

347

RESUL TS AND DISCUSSION

Scanning Electron Microscopy ~D.3 (Fig. 1, x 50 000) and BN (Fig. 5, x 10 000) powders show a sticked plate-like morphology. ZrOfY2D.3 (Fig. 2, x 40 000) and TiC (Fig. 9, x 40 000) are submicronic ultrafine spheroidal particles. AlN (Fig. 3, x 10 000) is composed of parti.cles of various sizes and shapes. SiC (Fig. 6, x 10000), B4C (Fig. 4, x 10000) and Diamond (Fig. 12, x 10000) particles are angular while Si3N4 (Fig. 7, x 50 000), TiB2 (Fig. 8, x 1 000) and TiN (Fig. 10, x 10000) present irregular shapes. Graphite parti.cles (Fig. 11, x 50) are nodular.

X-Ray Diffraction ~<l.J (Fig. 13) is pure a-alumina with a trigonallattice. Zr02 (Fig. 14) is present both in monoclinic and tetragonal systems. AlN (Fig. 15), B4C (Fig. 16), BN (Fig. 17), S~N4

(Fig.19), TiB2 (Fig. 20) and graphite (Fig. 23) show an hexagonal structure, they are an pure and (A1N)4H, (BN) 4H, (Si3N4) 28H and (TiB2) 3H are present respectively instead of AlN, BN, Si3N4 and TiB2• TiC (Fig. 21), TiN (Fig. 22) and diamond (Fig. 24) present a cubic lattiee, TiC and TiN are pure while diamond is mixed with Zr02 (tetragonal structure). The SiC phase (Fig. 18) is an a-phase with a rhombohedral structure.

Complement Activation Table 1 shows the date from the CH50 tests.

TABLEI CH50 Unit - Total complement : consumed percentage.

Sampie Powder CH50Unit Total complement Number assays

Consumed percentage I sampie powder

Sera Pool 95 0 3

ZymosanA <2 98 3

~0.3 95 0 3

~/Y20.3 87 8 3

AlN 91 4 2

B4C 80 16 2

BN 84 12 2

SiC 87 8 3 Si3N4 87 8 2 TiB2 95 0 2

TiC 91 4 3

TiN 91 4 3

Diamond 84 12 3

Graphite 95 0 2

348

Figure 1. A120.3 (x 50 (00) Figure 2. ZrOzlY203 (x 40 (00)

Figure 3. AlN (x 10 (00) Figure 4. B4C (x 10 (00)

Figure 5. BN (x 10 (00) Figure 6. SiC (x 10 (00)

349

Figure 8. TiB2 (x 1 000)

Figure 9. TiC (x 40 000) Figure 10. TiN (x 10 000)

Figure 11. Diamond (x 10000) Figure 12. Graphite (x 50)

- ...... p ........... __ .. _ .- ... _- -

fl Figure 13. A12O:3

._ .. .. . _. - .. _-~ _ .. ---

Figure 15. (AIN) 4H

Al I

Al , Al

•• .. •• • {. '-::.,=". -:": • • ~ • • '-':1.0,-. -!-. • • '""'.~.-= • .,... ~ •• ,...,.~. -=.-t-'. ':-: • • ~.:-=-. -: .. Figure 17. (BN) 4H

350

.. __ .. ----­_._--

-----­_ .. _._--

..... ~ ... ~-~~. Figure 16. B4C

Figure 18. SiC

---_ ....... . _.- ..-_-

--- . .. _._--­_ .- --

.. =00;:::--:"=_-

I .~.l! ~i • • • " J.'Y~L.--.-...Jt~L.

Figure 21. TiC

Al

._-- --. -_ ..... _- ... _ .. ---

~~ ~A1 t Al

.~~~-.~.~.~.~'"""""~!~ Figure 23. Diamond

351

....... UI"

L 1 IILU - --.. -- -- -. .. ~ - .. .. .. ~~ ~ . . ~ .. . -

I

- - --~... . _0- _. ___

I .

..----,..-~.'y.~ ..... . dJ.L.

.' I

Figure 22. TiN _111 ' .. _. ___ _

-'. ---

Al Al Al t

Al t

__ ... ) '--"""""~ ."-._ • ./1._. _._. .... . . _ .. _._."'._._._. _ • •

Figllre 24. Graphite

352

The positive control shows a total complement activation while the pool of sera shows none. Despite the variation of percentage of consumed complement (from 0 to 12) due to the sensitivity of the method itself, we can say that no ceramic powder initiates a significant in vitro complement activation.

CONCLUSIONS

The results of the tests described here are very encouraging and on the basis of them, we concluded that the tested ceramic powders and therefore the corresponding bulk ceramies or ceramic coatings could be candidate materials for our device as weH as for many other biomedical devices.

REFERENCES

1. Monties, J.R. and Havlik, P. , Mechanical pump destined for left ventricular assist device and for total permanent cardiac replacement. Acta Cardiolol:ica, 1982, Suppl. xxvm, pp. 136-138.

2. Dion, 1., Baquey, Ch., Candelon, B. and Monties, J.R., Hemocompatibility of Titanium Nitride. To be published in the International Journal of Artificial Orl:ans.

3. Payne, S.M. and Horbett, T.A., Complement activation by hydroxyethylmethacrylate­ethylmethacrylate copolymers. Journal of Biomedical Materials Research. 1987,21, pp. 843-859.

4. Kazatchkine, M., Hauptmann, G. and Nydegger, U., Dosages hemolytiques du complement. in Techniques du Complement. Ed. INSERM, Paris, 1985, pp. 15-48.

ACKNOWLEDGEMENTS

We would like to thank, Pr. R. Salmon, Mr. Trut and Mr. Cazorla from the Laboratoire de Chimie du Solide - CNRS (Univ. Bordeaux I). We also thank, F. Rouais for his technical assistance, M.O. Rubio for her photographic work and M. Rouais for typing the manuscript.

353

summ CHROMATID EXCHANGBS (seEs) AHn PROLIFERATION RATE INDEX (PR!): THE APPLICATION OF CYTOGENETIC KK'.ftIODS IM BIOCOKPATIBILITY FIELD.

*MARIO CANNAS, *SANDRA BIASIOL, *ALESSANDRO MASSE', AALESSANDRO RUGGERI, ARITA STROCCHI *Universitä di Torino, Dipartimento di Anatomia e Fisiologia

Umana, C.so M. D'Azeglio 52, 10126 Torino, Italy. AUniversitä di Bologna, Istituto di Anatomia Umana Normale.

Via Irnerio 48. 40126 Bologna, Italy.

ABSTRACT

Cytologically visible damage in human chromosome detected as Sister Chromatid Exchanges (SCEs) or other structural chromosomal aberrations is one of the "In Vitro" steps to assess the biocompatibility of soma materials to be used in orthopaedy or dentistry. The utility of these observations is that they point out that similar alterations may have occurred in other tissues, either somatic or germinal cells, evenience with possible clinical implications. The absence of effects on the DNA of replicating human lymphocytes cultivated in presence of three different types of Alumina (105 NS, 105 SFP and 130 SF, Metco Ind.) and an Hydroxyapatite, at different concentrations, let these materials to be classified as "negative", (from the point of view of being a possible genotoxic substance), as they failed to induce a significant increase in SeEs. They result also not inducing effects on the proliferation rate index (PRI) on the same cell type.

IHTROOOCTION

Different biological effects may be observed following cellular contact with materials used for implantation orthopaedy and dentistry, so cytogenetic techniques may useful in preliminary steps of biocompatibility testing.

the in be

The induction of sister chromatid exchanges (SeEs) by different groups of bioceramics, Alumina and Hydroxyapatite. respectively not bioactive and bioactive. was investigated in an in vitro culture of lymphocytes, as SeEs have been

354

suggested to be a valid method for predicting and evaluating the mutagenic/carcinogenic potential of a chemical in a dose dependent induction. SCEs represent symmetrical exchanges between sister chromatids; they do not result in alteration of the chromosome optical morphology or genetic information; their presence is restricted to cases where the exposing agents are strong alkylating compounds (e.g. cytostatic) or to some multi-exposure conditions (e.g. cigarette smoking). For example, Triethylene melamine (TEM) , a known animal carcinogen, was demonstrated to be an effective inducer of persistent SCEs in murine lymphocytes so demonstrating the ability of the method to identify genotoxins [1]; cyclophosphamide, a mutagen carcinogen, caused dose-related SCEs (five-fold increase over controls) both in vive and in vivo/ vitro [2].

Other authors used Short-term tests for the control of genotoxic potential of chemicals [see, for example, 3].

In the present study we describe a cytogenetic analysis specifically the SCEs value of lymphocytes from normal human donors exposed to culture medium containing different types of Alumina and an Hydroxyapatite.

The employ of Alumina in the orthopedic field is due to its high resistance to mechanical stresses and the weIl clinical documented biological compatibility of the "in toto" prosthesis. Hydroxyapatite is employed due to its osteoinductive effect as powder blocks to fill cavities or bony defects in order to induce bone ingrowth, or like plasma­sprayed on metallic prosthesis to create a good connection between metal surface and bone tissue on the implanted site [4]. Some observations about a mitogenic effect of synthetic Hydroxyapatite have been reported [5].

The investigations on the biological compatibility of these two bioceramies are numerous and especially based on the evaluation of tissue reaction to their implantation as weIl as of their effects when tested in cultures with macrophages [6].

MATERIALS AHn METHODS

Alumina of different types, 105 NS (5-25 micron, A12 =3 99%), alumina 105 SFP (5-35 micron, A12 03 98%) and Alumina 130 SF (5-35 micron, A12 03 80% and Ti 02 18%) obtained directly by the producers, METCO Industries, were sterilized with gamma rays (1.106 rad) and used as powders at different concentrations (1 mg/10 ml, 1/100 ml and 1/1000 ml); Hydroxyapatite (5-300 micron), some concentrations.

Human peripheral lymphocytes were collected from heparinized whole blood from a cytogenetically normal male, set up in TC RPMI 1640 (Difco) supplemented with Fetal Bovine Serum (FBS, Gibco, 33%), Phytohemoagglutinine (PHA, Gibco 2.5%), Penicillin and Streptomycin (1%) and the tested sampies powders; after 24 hours Bromodeoxyuridine (BrdU, Sigma) was added to the cultures at a final concentration of 5 ug/ml, as it is incorporated as a detectable marker instead of thymidine

355

in the chromosomal material. Cultures were incubated in the dark and Colcemid was added 1 hour before the harvesting at a final concentration of 0.25 u~/ml. After 72 houre the lymphocytes have completed three cycles of DNA replication.

To minimize laboratory variability that might result in uncertain results, all epecimens were treated in the same standardized method, ueing a simple batch of all materiale.

The culturee were fixed accordin~ to standard methode, employing an hypotonie treatment in 75 mM KCl for 10 minutee (37 degrees centigrade) and a fixation in methanol/glacia1 acetic acid 3:1; 2 drops of the fixed cells suepeneion were dropped on a clear wet elide and allowed to dry overnight. Slides, air dried, were then etained with Hoechst 33258 and exposed to U.V. light. Giemea dye followed (1% in buffer solution, pH 6.9 for 25 minutes) according with etandard method, Fluorescence-plus-Giemsa (FPG).

All experiments were done in quadruplicatee. Controle were repreeented by lymphocytes in the same

experimental culture conditione but without the adding of the powdere.

Twenty-five Giemea-etained well-spread second divieion metaphaees from each culture experiment were analyzed; slidee were randomly coded and scored blind to avoid ecore bias; all SCEs scored were done by the same individual.

To better define the cellular kinetics, the proliferation rate index (PRI) of the cultures was calculated; from each culture the number of cells in first (M!), second (M2) and third (M3) divieion was quantified (100 metaphasee were considered in all). Cell kinetic as PRI, is a factor involved in SCEs frequencies, significantly affected by mutagens and carcinogene [7].

All reeults were stored in a data-base and then processed with a commercial statistical package, in order to per form two-ways and one-way ANOVA teet and linear re~reseion tests.

RESULTS

Not statietically significant differences in SCEs values were observed for all the tested sampies respect to the contro1s (p>0.1) with no clear dose-response relationship, as the values were similar at the different concentrations (Figure 1 ), ( p>O . 5) .

the (not

value

Similar results were obtained from the controls and experimental eamples with reepect to the PRI index statietically eignificant differences, p>0.5, in PRI respect to the contro1s), Figure 2.

From literature, the mean frequency of SCEs in normal subjecte variee from 8.29/ce11 (+/- 0.08, SD), [8] to 5.43/ce11 (+/-2.14, SD) [9], statietica11y higher in fema1es that in males.

356

SCE. e 5

4

3

2

1

0

1:10 4.n5 4 .64 4.7 4.925 1:100 5.463 4.6715 4.15 5.575 1:1000 15.425 15.075 4 .6152 4.675 CONTROL. 4 .55 4.155 4.55 4.55

.1:10 0 1:100 1:1000 - CONTROL.

Flgure 1. For the explanation see text

PRI 2. 15 ~--------~----------,---------~r---------,

2

1.15

1

0 .5

o

1:10 1:100 1:1000 CONTROL.

2.2515 2.338 2.26

2.355

2.23 2.313 2.408 2.355

2.333 2.223 2.213 2.355

2.413 2.275 2.353 2.355

• 1:10 E2J 1:100 _ 1:1000 - CONTROL.

Figure 2. For explanation see text

357

DISCUSSION

In preventive medicine, biological monitoring methods are used for early detection of harmful exposure; these methods consist of measurements among a group of exposed people, using validated measurements to estimate the interna 1 dose (10]. Exposure to agents that may result to be mutagenic happens regularly by using substances that substitutes specific parts of the normal body constituents: as in every implantological actions; from this basis principle derives the importance to declare a substance to be implanted absolutely safe from the point of view of its mutagenic effects.

At present, the FPG staining of BrdU-labeled metaphases is the most frequently used technique for sister-chromatid differentiation, due to its excellent microscopic resolution (11], as the induction of sister chromatid exchanges is an exquisitely sensitive detective method to detect the exposure to hazardous mutagenic chemicals, as cigarette smoking and coffee consumption (12]; for example smoking is associated with an increase of approximately 2 SeEs per cell and a decrease in cell proliferation [12, cited].

An exaustive study of factors that may be responsible for SeEs response variability is impractical, and is by-passes in the present study by using the same plasma pool for the controls as for the experimental phase.

From our observations, both Alumina (in the different formulations) and Hydroxyapatite, didn-t demonstrate to produce chromosomal damage in human cultures of lymphocytes. From the literature [13] all agents which induce verifiable statistically significant increases in SeEs (p<O.OOl) should be classified as "positive"; agents which produce variable results (e.g. 0.05>p>0.001)should be classified as "indeterminate" or "provisionally positive"; "negative" should identify agents which fail to induce significant increases in SeEs.

The results obtained from our experiments 1st the sampies fail in the third situation.

The parameters connected with the proliferation rates of the cultures demonstrate that the tested materials do not interfere with the ce11 kinetics expressed by the PRI formula:

as in (7], cited.

Ml + 2 M2 + 3 M3 100

OONCWSIONS

in do

the not

The reeulte obtained with the samples of alumina different concentrations and physical configurations differ from what obtained by two of us in observations using a different experimental model both the previous and the present investigations no

previous [14]; in

indication

358

of chromosomal damage may be reported; for the used experimental conditions the results appear to be reproducible.

The observations reported in this paper will be further assessed by using the micronucleus technique, a method for the measurement of chromosomal damage in mutagens-stimulated lymphocytes, able to determine the proportion of cells that have responded to amitogen (PHA) , the proportion of the responding cells that have divided and the fate of micronuclei (an acentric chromosome fragment or whole chromosome not incorporated in the main nucleus at cell division) in the cells [15].

Acknowledgments The authors are grateful to Mrs. Elena Rizzo for her technical assistance. This research has been supported by M.U.R.S.T. grant ("quota 60%") .

REFERENCES

1. Neft, R.E., Schol, H.M., Casciano, D.A., Triethylene melamine-induced sister-chromatid exchange in murine lymphocytes exposed in vivo., Mut Res, 1989, 222, 323-328.

2. Krishna, G., Nath J., Ong T., Comparative In Vivo and In Vitro Sister Chromatid Exchange Studies in Chinese Hamster Bone Marrow and Spleen Cells., Teratogenesis. Carcinogenesis. and MutaQenesis, 1986, 6, 321-330.

3. Sobels, F.H., International symposium on short-term tests for genotoxicity., Mut. Res., 1986, 164, 389-394.

4. Cook, S.D., Thomas, K.A., Kay, Hydroxyapatite-Coated Titanium for Applications., Clin. Orthop Rel Res,

J. F. , J archo , M. , Orthopedic Implant

1988, 232, 225-243.

5. Cheung, H.S., Story M.T., McCarty D.J., Mitogenic Effects of Hydroxyapatite and Calcium Pyrophosphate Dihydrate Crystals on Cultured Mammalian Cells, Arthritis and Rbeuroatism, 1984, 27 (6), 668-674.

6. Guizzardi, S., Di Silvestre, M., Tarabusi, C., De Pasquale, V., Ruggeri, A., Pizzoferrato, A., ATEM investigation on macrophages, exposed to alumin and related materials., Ahs IIIrd World Biomaterials Congress, April 1988, Kioto.

7. Lamberti, L., Bigatti Ponzetto, P., Ardito, Kinetics and sister-chromatid exchanges frequency lymphocytes., Mut Res., 1983, 120, 193-199.

G., Cell in human

8. Bender, M.A., Preston, R.J., Leonard, R.C., Pyatt, B.E.,

359

Gooch, P.C., Shelby, M.D., Chromosomal aberration and sister­chromatid exchange frequencies in peripheral blood lymphocytes of a large human population sample., Hut Res, 1988, 204, 421-433. .

9. Cannas, M., Bigatti, P., Rossi E., Rossi P., Ricerche in vitro sulla possibilita di danno cromosomico da polimetilmetacrilato in Ortopedia., Giornale ltaliano di Ortopedia e Traumatologia, 132 (3), 1987, 399-403.

10. Sorsa, M., Monitoring of sister chromatid exchange and micronuclei as biological endpoints., in "Monitorinlj! human exposure to carcinogenic and QlUtagenic agents", Berlin A., Draper M., Hemminki K. and Vainio H edts, lARC, Lyon, N.59, 1984.

11. Haaf, T., Ott, G., Schmid, M., Differential inhibition of sister chromatid condensation induced by 5-azadeoxycytidine in human chromosomes., Chromosoma, 1986, 94, 389-394.

12. Reidy, J.A., Annest, J.L., Chen, A.T.L., Welty, T.K., Increased Sister Chromatid Exchange Associated With Smoking abd Coffee Consumption., Enyironmental and Molecular Hutalj!enesis, 1988, 12, 311-318.

13. Best, R.G., McKenzie, W.H., Variable sister-chromatid exchange response in human lymphocytes exposed in vitro to gossypol acetic acid., Hut Res.,1988, 206, 227-233.

14. Guizzardi, S., Oi Silvestre, M., Govoni, P., Ruggeri, A., Biocompatibility of implants of alumina-powder in rat., ~ & Applied Histocbem , 1988, 33 suppl., 148, 9.

15. Fenech, M., Morley, A.A., Measurement of micronuclei in lymphocytes, Mut Rea, 1985, 147, 29-36.

360

SURFACE COATING OF PECVD a-SiC:H TO IMPROVE BIOCOMPATIBILI1Y

ARMIN BOlZ and MAX SCHALDACH Zentralinstitut für Biomedizinische Technik der Friedrich-Alexander-Universität

Erlangen-Nümberg, Turnstr. 5, D-8520 Erlangen

ABSTRACT

Hemocompatibility of a material is essentially determined by the electronic properties of its surface, whereas functionality is provided by the bulk properties. As most materials are not satisfying both requirements, a hybrid design is the only way to solve the material related problems in replacement surgery. In the following a microscopic model of thrombogenesis induced by solid surfaces is shortly reviewed and the electronic requirements for high hemocompatibility are deduced. The physical properties of amorphous silicon carbide (a-SiC:H) concerning the band gap, the density of states and the conductivity show, that a-SiC:H coatings meet these requirements. In vitro measurements (TIRIF and TEG) prove the hemocompatibility of a-SiC:H coated parts. Additionally corrosion tests confirm a good long time behaviour, so that a-SiC:H is well suited as antithrombogenic coating material for implants.

INTRonUCTION -TIIE SURFACE INnUCEn CLOTTING

The decomposition of fibrinogen into fibrin - the most important and irreversible step in thrombogenesis - is not only induced by thrombin but also by an electron transfer from fibrinogen to a solid's surface. Band structure calculations as weIl as electrochemical measurements at semiconductor-fibrinogen interfaces demonstrated that the transfer level must be between -0.9 and -1.3 eV below Fermi's energy (see also fig. 1) [1,2]. The loose of an electron causes arelaxation of the atomic structure in the fibrinogen molecule. Due to the Franck-Condon-principle the charge transfer is much faster than the relaxation resul­ting in a high relaxation energy of about 2 e V per amino acid, large enough to break the peptide bond of the fibrinopeptides.

361

In order to improve the hemocompatibility of an implant the electron transfer to its surface must be inhibited. Figure I compares the density of oceupied states in the protein DFb-(E) with that of unoccupied states in the solid D:lJI + (E). The density of states (DOS) in the valence band offibrinogen NFb(E) is about 10 cm-3 eVl and at body temperature almost any state is occupied by electrons [3]. According to

Ev Ev

jv=C e J DJiI(E) DFiJ(E) dE = Ce J NHI(E) (l-f(E» NFb(E)f(E) dE (1) -00 -00

(where C is constant, e the charge of an e1ectron and feE) Fermi's function) the number of unoccupied states in the valence band of the solid DHl + (E) determines the exact value of the exchange eurrent jv and with that the thrombogenicity of the material. However, DHI + (E) depends on the distance E between the energy level under consideration and Ferrni's energy as weH as on the density of states NHI(E) in the material.

Q) Cl) E E

complete denslty of denslty unoccupled of states s ates

0 -------- - ---- .... - ....

0 [

-{)9

- 3

Surface of the Solid

E

A

denslty of occupled states

Charge rans fer Inter val

Eg : 1 8 eV

Fibrinogen

E

Figure 1. Schematic distribution of the densities of states at the interface between the solid's surface and an aqueous fibrinogen solution.

The second picture of figure 1 compares four different distributions of unoccupied states in the solid (A, B, C and D), their correspondin~ densities of states NHI(E) are shown in picture 1. Case A with a constant DOS of about 102 cm -3 e V l is equivalent to a common noble metal; B, C and D correspond with semiconductors with increasing band gap. Thus in order to decrease the exchange current significantly, NHl(E) has to be lowered. As a resuIt, the most important requirement for good hemocompatibility of a material is a low DOS in the critical energy interval between 0.9 and 1.3 e V below Fermi's energy. In reality this low density can only be achieved by a semiconducting material with a band gap of more

362

than 2 eV. Additionally the conductivity should be more than 10-4 (Qcmr1 in order to stabilize the electrochemical equlibrium at the interface. On the one hand this is an explanation for the better hemocompatibility of passivated metals in comparison to noble metals due to the semiconducting properties of most metal oxides. On the other hand special semiconducting coatings with tailored electronic properties should show superior hemocompatibility.

MATERIALS AND METHODS

In order to prove this theory semiconducting coatings of amorphous hydrogenated silicon carbide (a-SiC:H) were produced using the plasma enhanced chemical vapor deposition process [4]. The filrns were deposited from mixtures of silane (10% diluted in hydrogen) and methane (pure) in a 13.56 MHz capacitively coupled plasma at 0.1 mbar. In order to achieve a low DOS in the band gap the substrate temperature was hold at 250 oe. The n-doping was done with phosphine (0.1% in hydrogen) resulting in conductivities of up to 10-3 (Qcmr1. The exact details of sampie preparation has been discussed previously [5]. A schematic of the charge and potential distribution at the interface of these coatings is shown in figure 2. Due to the hydroxide ions in the solution, a-SiC:H get a thin passivating layer of silicon dioxide on its surface with an equilibrium thickness of about 2 nm.

Energy of electrons

EHl V

OXIde layer

Olstance X

Figure 2. Charge and potential distribution at the interface between pure amorphous silicon and an aqueous fibrinogen solution (pH 7.4). Ee is the lower edge of the conduction band, Ev the upper edge of the valence band and Er is the abbreviation of Fermi's energy. The superscripts Fand HI denote the fibrinogen molecule (F) and the semiconductor (Hl).

363

The variation of the gas composition gives an opportunity to modify the difference between Fermi's energy Ef and the upper edge of the valence band Ev• The exact values for a-SiC:H were derived from band gap and activation energy measurements; the former were performed with UV -VIS-NIR transmission experiments using the Tauc-formalism [6], the latter by measuring the temperature dependence ofthe electronicconductivity. Amorphous silicon has an Ef-Ev value of 1.4 e V, steadily increasing with increasing carbon content leading to Ef-Ev of 1.6 e V at 70% methane concentration. The effect of different values of Ef-Evs (with Evs as the valence band position at the surface) on the exchange current is demonstrated in figure 3 for crystalline and amorphous semiconductors with 0.6 e V broad tail states. For low values of Ef-Evs, the charge transfer is possible between the two valence bands resulting in a maximum exchange current. However, as soon as Ef-Evs exceeds 0.9 e V the current decreases by several orders of magnitude. However, depending on the band gap DOS there is a saturation effect for high Ef-Evs. The field effect derived band gap DOS of a-SiC:H is some 1016 e Vi cm'3. Therefore, taking a little band bending of about 0.1 e V into consideration, a-SiC:H should expose increasing hemocompatibility with increasing carbon content with a saturation effect at high carbon concentrations.

Exchange Current [Rel. UnitsJ

(EI'Eva) • 0 .9 .V IEI'Eva' ' 0 .9 .V 1000000~ __________________ ~

100000

10000

1000 NIE)

I 100 10'22 amorPf '

10 10'16

10'18

10'17

I E EI Ev

10'18

0,1 '---'---o 0,2 0 ,. 0 ,8 0 ,8 1,2 1,. 1,6

Ef-Evs [eVI

Figure 3. Effect of Ef-Evs on the exchange current respectively the activation of fibrinogen.

RESULTS

This prediction was checked with pure bovine fibrinogen solutions using the TIRIF- (Total Internal Reflection Intrinsic Fluorescence) as weIl as in vitro with human blood using the TEG- (thrombelastography) technique. TIRIF utilizes the intrinsic fluorescence of proteins containing aromatic amino acids like tryptophan or tyrosin enabling a concentration measurement at interfaces [7]. For this purpose ultraviolett light (285 nm) is totally reflected

364

at the interface, whereas the evanescent wave excites adsorbed molecules. A photomulti­plier measures the emitted fluorescence light. After flushing the interface with saline at pH 7.4 and 37°C for one hour, 0.5 % fibrinogen solution was added. After different contact times the interfaces were flushed again with pure saline and the remaining protein was measured in relation to the maximum adsorbed protein concentration. Tbe results are demonstrated in figure 4.

Remaining Protein 1%1 80 ----------------------~~

a-SiC:H

3

20 o 300 1100 900 1200 1500 1800 2100 2400 2700 3000 3300

Contact Time Isl

Figure 4. Comparison of the remaining protein concentration after different contact times for silica and a-SiC:H (50% methane) coated silica. Tbe a-SiC:H was deposited from a mixture of 50% silane and 50% methane. 100% corresponds to the maximum concentration

achieved during the adsorption phase.

Time [mini 18

16 Reaction Time ~ Clotting Time

14

12

10

8

6

4

2

O ~~~~~~~~ __ ~~~~~UU~-L~~L-~ Titanium 0'" 21% 50% 75% Methane

Figure 5. Effect of different a-SiC:H coatings on the clotting and reaction time measured with the thrombelastography technique.

365

Silica activates fibrinogen due to its low conductivity and leads to an increasing polyme­rization and therefore an increasing concentration of remaining protein. Sputtered carbon films revealed a similar behaviour due to their high DOS. On the contrary a-SiC:H coatings show no time dependent increase in the remaining protein concentration confirming that no fibrinogen activation and polymerization takes place. Additionally the hemocompatibi­lity of a-SiC:H was checked with the TEG-technique in order to prove the predictions for whole blood [8]. The test specimens were milled from pure titanium as a reference material and coated with a-SiC:H of different compositions. All experiments were performed with blood of the same donor, the results are summarized in figure 5.

DISCUSSION

The results verify the electrochemical model for the activation of fibrinogen at solid surfaces. Using a-SiC:H deposited from the correct gas mixture (corresponding to a correct value of Ef -Ev), the clotting time is enhanced by more than 200% in comparison to titanium. Furthermore corrosion experiments proved an excellent corrosion behaviour due to the thin passivating layer on the surface; corrosion rates at physiologie conditions are below 30 nm per year [9]. Additionally a-SiC:H has a microhardness of about 20 GPa resulting in high wear resistance. All results together conflrm that a-SiC:H coatings are improving the blood compatibility of implants.

REFERENCES

[1] Bolz, A and Schaldach, M., Improved materials for heart valve replacement. Proc. of the First Euro.p. Cont on Biomedical Engineering, 1991, 303-06.

[2] Baurschmidt, P. and Schaldach, M., The electrochemical aspects of thrombogenicity ofa material. J Bioeng., 1977,1,261-69.

[3] Bakshi, A K, Ladik, J., Seelz M. and Otto, P., On the electronic structure and conduction properties of apenodie DNA and proteins. IV. Electronic structure of aperiodic proteIns. Chemical Physjcs, 1986, 108, 233-41. .

[4] Kushner, M. J., A model for the discharge kineties and plasma chemistry during plasma enhanced chemical vapor deposition of amorphous silicon. J Appl Pbysics, 1988, 63, 2532-51

[5] Bolz, A and Schaldach, M., Artificial Heart Valves: Improved blood compatibility by PECVD a-SiC:H Coating. Artjficial Organs, 1990, 14, 260-69

[6] Tauc, J., Amorphous and liquid semiconductors. Plenum Press, New York, 1976. [7] Van Wagenen, R. A, Rockhold, S., Andrade, 1. D., Probing protein adsorption by total

internal reflection intrinsic fluorescence in biomaterials: lnterfacial phenomena and applications. Ady Chem. Series, 1981,199,351-70

[8] Hartert, H., Blutgerinnungsstudie mit der Thrombelastographie, einem neuen Unter­suchungsverfahren. Klin. Wocbenschr., 1948,37.

[9] Bolz, A, Brem, B. and Schaldach, M.t Electrochemical corrosion behaviour of anti­thrombogenic amorphous silicon carbIde coatings. In Proc of tbe twelftb Int. Conf of tbe lEEEI EMBS, Philadelphia,1990, 2087-89

366

BONE TISSUE RESPONSE TO HYDROXYAPATITE - COATED AND UNCOATED TITANIUII WIRE - IlESHS IN AN INPECTED SITE

RESULTS OP AN ANIIIAL EXPERDIENT

AXEL WILKE, JOACHIM ORTH, MARKUS KRAFT, PETER GRISS Orthopaedie Department of Philipps-University,

Baidingerstraße,3550 Marburg, Germany

ABSTRACT

The authors studied the ingrowth dyn ami es of bone tissue into the pores of Hydroxyapatite-eoated (plasmaspraying teehnique) and uneoated wire-meshs of pure titanium in an infeeted implantation site. Sampies of the tested materials were implanted into the femora of 16 adult Göttingen minipigs. Just before implantation they were eontaminated with Staphyloeoeeus aureus. The animals were saerifieed after 4, 8, 12 and 24 weeks. Undeealclfied ground­seetions of bone tissue were prepared (app. 100 mierons) and stained with toluidin blue for eomparative histologieal evaluation.

The HA-eoated implants demonstrated advaneed new bone formation already after 4 weeks and nearly eomplete osseointegration after 12 weeks although a11 sampies showed gross and histologieal signs of persisting infee­tion. Comparable reaetions of the uneoated Implants eould only be observed after 24 weeks. Signs of degradation of the hydroxyapatite-eoating eould be seen in eontaet to soft tissue and were more extensive eompared to the behaviour in an uninfeeted site.

INTRODUCTION

Durlng the last deeade HA-eoated implant materials were introdueed into elinieal praetiee (1, 2, 3). Experimental results proved, that HA ean be regarded as a bioaetive material, whieh demonstrates no toxieity, antigenity and cancerogenity (1). Furthermore a direet osteogenesis was reported on HA-surfaces (4). These promising experimental results where eonfirmed by the clinical appllcation of HA-coated implants (2). Nevertheless desinte­grative processes 01 HA-eoatings and loosening of some of these implants are not entirely understood.

In the course of an earlier experimental study in our institution we observed, that in spite of ineidental infeetion 01 HA-eoated implants osseointegration was only slightly delayed (6, 6).

IlATERIALS AND IlETHODS

Three different preparations 01 a 4-layer sintered pure titanium (ISO 6832-2) wire-mesh were tested.

367

Uncoated titanium (Tl) wire mesh (7) Ti-wlre-mesh coated with an HA-ceramic (8) by plasmaspraying wlth an average thickness of 200 microns . Ti-wire-mesh (5-layers) wlth an unilateral coating of Ultra High Molecular Weight Polyethylen (UHMWPE) RCHR-I000 (7) (Figure 1)

Figure 1. Implantmaterlals. I-HA. 2-Tl. 3-PE.

The wire-meshs were press-fitted into a slot of the femur created by a sagittal osteotomy of left and rlght femur of 15 adult Göttingen Minipigs . The osteotomies were located at the proximal (HA and Tl) and distal (UHMWPE) metaphysis.

The implants were contaminated just before implantation wlth 50 microllter of a blood-brain medium with added niac1n. that contained 103 Staphylococcus aureus. This bacterium was donated by The Faculty of Veterinary Medicine of The University of Gießen. Germany (Strain derived from the pig).

The animals were sacriticed after 4. 8. 12 and 24 weeks. Before explantation of each femur two bacteriological swabs of the perifemoral abscesses were taken under sterile conditions and evaluated by The Microbiological Institute of The Phllipps - University Marburg.

The pH-value of the perilmplant tissue was examined with a surface microelectrode U-402-M8 (Fa. Ingold Meßtechnik GmbH. 6374 Steinbach/Ts .• Germany).

After explantation and preparation of the femura contact x-rays of the implantation site were taken. Hard tissue secUons (app. 100 microns) were prepared by grlnding and stained with toluidine blue for comparatlve histological studies using standard fixation and embedding techniques (for technical details see Donath. K. (9»

RESULTS

lIacroscoplcal observations All animals. that were sacrificed after 4 and 8 weeks showed an extensive perifemoral abscess . This dramatic reaction at the infected implantation site decreased wlthin the following weeks. 24 weeks after operation none of the implantation sites demonstrated macroscopical signs of infection.

Fractures occured in 6 cases and were always located at the implanta-

368

tion site. Accordingly 13 out of 33 implants were displaced into the surrounding soft tissue. It is note worthy. that 8 of the 13 displaced im­plants were those coated with polyethylene.

The extend of bony integration was roughly esti~ated by manual force applied to each implant in the implantation bed. It turned out. that the stability of osseointegration seen around the HA-implant 4 weeks after implantation was already comparable to the stability of Ti-implants after 24 weeks. All implants with a unilateral PE-coating were displaced into surroundlng soft tissue after 8. 12 and 24 weeks. Those after 4 weeks showed no osseointegration at a11 and were surrounded by soft granulation tissue.

Bacterlologlcal re8ults and pH-value8: Staphylococcus aureus could be cultivated in 70 % of a11 bacteriological sampies. that were taken trom each implantation slte prior to explantation of each femur. 30 " of a11 sampies showed bacteriologlcal results. differing from the strain lnocculated at operation. There were no negative bacteriological tests throughout the experiment. PH-values at the lmplatation sites. that were evaluated Just be fore explantation had a range from 6.44-7.60 and no differences could be found after 4. 8. 12 and 24 weeks (see also dlcussion) .

1I1croscoplcal observations: 4 weeks HA: The osteotomy slte around the lmplant and the entire medullary cavity showed a vast lnflammatory reaction with numerous granulocytes and multiple fresh abscesses of variing slze. Endostal. periostal and medullary new bone formation were seen and appeared frequently in association wlth bone resorption in preexisting cortical bone structures. Numerous active osteoblasts were involved in this reaction. A discontinous contact of newly formed woven bone was evident between cortical bone of osteotomy and implant surface. No interposed layer of connective tissue was seen at the interface and primary bone formation occured directly on HA -coatings. Thls bone formation was associated with a Une of active osteoblasts. Evidence for desintegrative processes at HA-Iayers could not be demonstrated (Fig.2) .

. 11 [ 11111111111111111111.1111111111111111

Figure 2. HA-coated implant after 4 weeks.

4 weeks Tl: Reactions due to induced infection were similar to those described around HA-coated specimens. The implants were however completely encapsulated by an abscess-membrane. that was filled wlth numerous

369

granulocytes and tissue debris . Newly formed woven bone was rarely seen and termlnated at the abscess membrane, separating cortical bone and Implant surface (Fig.3) .

•

Figure 3. Ti-implant after 4 weeks.

4 weeks PE-Ti: In PE-Ti implants baslcally the same reactions could be demonstrated as in pure Ti implants.

8 and 12 weeks HA: The discontinous woven bone at the interface, which was seen after 4 weeks now formed a continuous layer between implant and cortical bone. Pores of the HA-coated Ti-meshwork were entirely filled with bone tissue. Parts of the HA-layer, that were in soft tissue contact demonstrated obvious signs of desintegration. HA-particles were phagocytosed by macrophages. This loss of coating amounted partly up to 50'" of the original thickness of the HA-coating. Compared to the 4 weeks results the inflammatory tissue reactions appeared less cellular and more organized. The number of granulocytes decreased and fibrotic granulation tissue was more prominent.

8 and 12 weeks Ti: The implants evaluated were still embedded in an active granulation tissue presenting neither significant new bone formation nor bone-implant contact.

8 and 12 weeks PE-Ti: All implants at this time of sacrifice were displaced and fully encapsulated in an abszeding granulation tissue.

24 weeks HA: The entire osteotomy slot and the voids within the implant were filled with bone, which was almost completely of lamellar structure. The number of osteocytes per area had however decreased. Haversian canals were seen as a sign of good vascularisation of bone. In those areas where HA-coating was in direct contact to bone no desintegrative processes of HA occured. In contrast almost complete desintegration of HA was seen in areas of HA-soft tissue (granulation tissue) contact. Mild signs of infections could be demonstrated.

24 weeks Ti: Titanium implants were completely integrated into bone similar to those coated with HA. In approximately 60 '" of the implant surface direct contact to bone was evident. In the remaining areas a thin connective tissue layer at the interface was seen. Similar to the results after 12 weeks weak signs of infection were evident (Fig. 4)

24 weeks PE-Ti: All implants at this time of sacrifice were displaced and fully encapsulated in infected granulation tissue.

370

Figure 4. Ti -implant after 24 weeks.

DISCUSSION

The infection model introduced has proved to be reproducible and standar­dized enough for studies in an infected implantation site in mini pigs. The desired infection occured in a11 animals and persisted in most up to 24 weeks. During the experiment bacterial shift could be seen in 30% of animals. Local pH-values of the infected tissue have proved to be relatively constant within a range of physiological pH-values.

It was clearly seen, that in spite of local infection HA-coated implants demonstrated good osseointegration already after 4 weeks . Signs of degradation of the HA-coating were evident after 8 weeks and increased during the entire experimental time. This resulted in a complete loss of HA­coating after 24 weeks in those areas where no bone contact was present.

Compared to earlier experiments with HA-coated implants without induced infection, completed recently in our institution (6, 6), osseointegration was only slightly delayed in case of infection. The accelerated HA-degradation seen in this experiment was exclusively restricted to these areas of the infected implantation si te where no bone contact was seen.

Infection is combined with a granulocyte and macrophage activation. Thus it can be concluded, that HA-degradation is due to an active granulocyte and macrophage activity (10, 11).

Experiments performed by Van Blitterswljks where macroporous HA­ceramics were implanted into middle ear of Wistar rats with local infection induced by Staphylococcus aureus demonstrated slmilar results (12). His studles showed a 76 % increase of macroporosity of HA-ceramic. It was postulated, that this was a result of increased activity of macrophages, where Incorporated HA-particles could be seen (13) .

Our experiments can not completely answer the question whether HA­degradation Is pH-dependent. The pH-values measured ranged trom 6.44-7.60 in the infected implantation sites. Ungethüm and Fink pointed out, that pH­values of 3-4 oblgate for an Increased degradation of bioactive materials (14). The Ti-implants showed a delay of osseointegration for approximately 20 weeks durlng infection. In contrast to this, osseointegration of both, HA-

371

coated and uncoated Ti-implants, in a non infected implantation slte have both been proved to be well advanced after 4 weeks (5, 6). Regarding cUnical aspects of the results described here the following conclusions can be drawn:

After infection of HA-coated implants an explantation can be expected to be more difficult, compared to an uncoated material for bone-bonding of the HA -coating is not disturbed.

It needs to be discussed, whether an infected HA-coated arthroplastic has to be removed at all or whether it demands only conditions of conservative (antibiotics) and operative (debridement) treatment without prosthesis exchange, for integration occurs despite local infection.

More research is needed to investigate the possible benefit of implants coated with HA and an additional antibiotic loading.

1. Osborn, J,F., Implantatwerkstoff Hydroxylapatlt. Grundlagen und kUnische Anwendung, Quintessenz Verlag, BerUn 1985.

2. Osborn, J.F., Die biologische Leistung der Hydroxylapatit Keramik­Beschichtung auf dem Femurschaft einer Titanendoprothese-erste histologische Auswertung eines Humanexplantates, Biomedizinische Technik, 1987, 32 7-8, 177-183.

3. Osborn, J.F., Hydroxylapatitgranulate und ihre Systematik, Zahnärztliche !4itteilungen, 1987, Heft 8/87, 77. Jahrgang.

4. Herren, Th., Remagen, W., and Schenk, R., Histologie der Implantat Knochengrenze bei zementierten und nichtzementierten Endoprothesen, Der Orthopäde, 1987, 16, 239-251.

5. Orth, J., Kautzmann, J. and Griss, P., Bone tissue response to porous hydroxyapatite and wire meshs of stainless steel wlth and without coatings of hydroxyapatlte and titaniumnitrlte, in CUnical Implant Materials, Advances in Biomaterials, Volume 9, ed. G.Heimke, U.Soltesz and A.J.C.Lee, Elsevier Sc1ence Publ1shers B.V.,Amsterdam,1990, 283-287.

6. Orth, J., Griss, P. and Wllke, A., Tierexperimentelle Beobachtungen zur Frage der Resorbierbarkeit von Hydroxylapatitbeschichtungen auf Dauerimplantaten, Vortragsmanuskript, XII. Münchener Symposium für experimentelle Orthopädie, 1990

7. Oehy, J., letter of 9.5.90 Fa. Sulzer, CH-8401 Winterthur. 8. Dörre, E., Hydroxylapatitkeramikbeschichtungen für Verankerungsteile

von Hüftgelenksprothesen. Biomed. Technik, 1989, 34, 46-52. 9. Donath, K., Die Trenn-Dünnschliff-Technik, Exakt-/Kulzer-Druckschrift

Norderstedt, 1987 10. Carlson, S., Josefsson, G. and Lindberg, L., Revision with Gentamicin­

Impregnated Cement for Deep Infections in Total Hip Arthroplasties, The Journal of Bone and Joint Surgery, 1978, Vol. 60-A, No. 8.

11. Lang, F., Pathophysiologie, Pathobiochemie, 1987, Enke-Verlag, Stuttgart. 12. Blitterswijk, van C.A., Grote, J.J., De Groote, K., Daems, W. Th. and

Kuijpers, W., The Biologieal Performance of calcium-phosphateceramics in an infected implantation site. I. Biological Performance of Hydroxyapatite during Staphylococcus aureus infection, J. of Biomed. Mat. Res., 1986, Vol. 20, 989-1002.

13. Blitterwijk, van C.A., Bakker, D., Grote, J.J. and Daems, W.Th., The biologieal performance of calcium-phosphate ceramies in an infected implantation site.II. Biological evaluation of Hydroxyapatite du ring short-term infection, J. of Biomed. Mat. Res., 1986, Vol, 20., 1003-1015.

14. Ungetüm, M. und Fink.~Bioaktive Werkstoffe, Eine kritische Obersieht, Z. Orthop., 1988, 126, 697-708.

15. Winter,M.,Griss,P.,Comparative histocompatibUity testing of seven calcium-phosphate ceramies, Biomat., 1981, 2, 159-161.

372

BIOCOIIPATIBILITY OP SILICONCARBIDE AND SILICONNITRIDE CERAIIICS. RESUL TS OP AN ANIIIAL EXPERIIIENT

JOACHIM ORTH, MARTIN LUDWIG, WOLFGANG PIENING, AXEL WILKE PETER GRISS

Orthopaedlc Department of Philipps-University, Baldingerstraße,3550 Marburg, Germany

ABSTRACT

To test the biocompatibility of siliconcarbide and siliconnitride ceramics little tiles of both materials were implanted into the femora of 30 rats. The bone tissue response was evaluated by qualitative and quantitative light microscopy. In a second experiment ceramic powders of SiC and Sb N 4 were injected into both knee joints and the left foot of rats. The tissue reponse at locus of application, in the organs of the reticulohistiocytary system and the kidneys was examined by lightmicroscopy. In both experiments the silicon containing ceramics demonstrated less biocompatibility in comparison to alumina, we used as reference material.

INTRODUCTION

For many years alumina ceramics are widely used in orthopaedic and dental surgery (2,3,8). The most important properties are good biocompatibility in contact to bone and other tissues and a high wear resistance. On the other hand there is not much information about the biocompatibility of alternative ceramic materials like silicon-carbide and silicon-nitride, which because of their mechanical properties could be of comparable importance. According to the guidlines of the Working Group Biomaterials of the German Society for Orthopaedics and Traumatology (DGOT) (7) new biomaterials should not only be tested in the final definite implant form.

Since wear debris partlcles of implants can induce fatal outcomes(6), the tissue re action to possible arising particles of new implant materials should be tested too. Further a material of reference with well known biological behaviour should be used.

OB.JECTIVES

The aim of the two presented experiments therefore was to test the biocom­patibility of SiC and Sb N4 in form of highly sintered bulk material in a non weight bearing animal model and in a second experiment in form of ceramic powder by injection into both knee joints and into the soft tissue of the left

373

foot of rats. In both experiments the material of reference was alumina ceramic.

MATERIALS AND METHODS

The implants used in the non weight bearing animal experiment were smooth surfaced little ceramic tiles with outside dimensions of 6x3xO.8 mm. They consisted out of Atz Da, Sie and Sb N4. Impurities are described in table 1.

TABLE 1 Impurities and me an grain sizes of used ceramics

mean gr.size MgO CaO SiOz Fez 03 ZrOz

Al. 0. 2.40 <0.01 0.008 0.01 0.008 0.01

Sie 3.05 <0.01 <0.005 <0.005 0.03

SiaH .. 2.94 <0.01 0.073 0.006 0.03

11m % % % % %

The animals used for the first experiment were 30 mature Wistar rats with a mean weight of 260gr (200-300gr) and a me an age of 56 days (47-62). The animals were anaesthesized by Pentobarbital Na. Monocortical slot osteotomies were prepared through a lateral approach to both femora using a dental drill. Each animal received one type of implant at both femora, resulting in 4 implants (2 animals) per time period. After 2, 4, 8, 12 and 24 weeks the animals were sacrificed.

After harvesting both femora they were contact radiographed and pre­pared for histological evaluation by formalin fixation and ethyl alcohol dehydration. The undecalcified femura were embedded in methylmethacrylat. Transversal sections of app. 140 llm were cut with an annular diamond saw. The slices were grinded reaching app. 50 llm thicknes and polished by use of alumina grains. Toluidine blue and Giemsa stained sections were evaluated qualitatively and quantitatively. For quantitative histology we used a semiautomated analysis system (ASM 68 KR, LEITZ). We exclusively measured in the cortical area by two different ways (Fig.I). First we measured the me an distance between implant surface and neighbouring old or new formed bone at both sides of the implant (Fig.1. left side). Secondly we measured the area of direct contacts of new formed bone to the implant surface and related it to the total implant surface in the cortical slot (Fig.l right side).

I M P L A N T

Figure 1. Methods of quantitative histology

I M

A1 P L A A2 N

A3 T

374

For the second experiment we used ce ramie powders. Grain sizcs are demonstrated above (Table 1.). The animals were 36 female Sprague rats with a me an weight of 336 gr (300-350gr). We prepared 20 % suspensions of the ceramic powders in 0.9% sodium solution. Under ether anaesthesia 0.05 ml of the suspension were injected in both knee joints and 0.1 ml in the soft tissue of the left foot. 2, 4, 8, 12, 24 and 52 weeks after treatment the animals were sacrificed. This resulted in 2 animals for each material group and each time period.

We harvested both knec joints and the left foot, inguinal and paraaortic lymphatic nodes, the spleen, the liver, both lungs and both kidneys for histological evaluation. After formalin fixation and ethyl alcohol dehydration the knee joints and the foot were embedded in methylmethacrylate, sectioned using a dense-section microtome and stained with toluidine-blue and Giemsa solution. The other organs were embedded in paraffin, sectioned with a standard microtome, and HE and van Gieson stained. Qualitative tissue reaction was evaluated by light-microscopy.

RESULTS

First experbnent: 2 weeks after implantation new formation of woven, unmineralized bone was seen only intramedullary in case of alumina and with a delay of further two weeks in case ofSb N.. In the cortical area new bone formation could be observed after 4 weeks only around alumina implants. After 8 weeks rarely direct bone contacts to the alumina-surface could be seen. SiC and Sb N4 implants demonstrated new bone formation in the cortical slot , seperated from the implant surface by a thick layer of fibrous tissue. The rate of direct bone contacts to the alumina implant increased time dependend up to 83 % until the end of the experiment (Fig.2, Fig.3a).

'ocr o. <o'''ca' .... 80

80

40

20

o ta 24

weeke after Implantation

_ ALUMINA ~ SILICIUMNITRIDE

Figure 2. Rate of direct bone/implant-surface contacts in the cortical area

After 12 weeks the maximum of bone contacts was reached in case of Sb N. (32%) (Fig.2), whereas no direct bone contacts appeared at the surface of SiC implants (Fig.3b).

During the whole experiment a maturation of new formed bone was ob­served, accomponied by a decreasing thickness of fibrous tissue, interposed between implant surface and bone (Fig 4.).

Figure 3a. Ah 03 -implant 24 weeks after implantation (cortical area)

(Magn. app. 50x)

micron

375

Figure 3b. Si 3 N4 -implant 24 weeks after implantation (cortical area)

(Magn. app. 50x)

::j 40 +_ ~ __ ----~~~

30

~: 1+-...,....--,.---,-------.---:-.----------, -----,..-----~--~~~ 2 4 8 8 10 12 14 18 18 20 22 24

weeks after implantation

ALUMINA SILICONCAR810E SILICONNITRIDE

Figure 4. Mean-thickness of interconnective tissue between implant surface and surrounding bone.

Second exper1m.ent: The second experiment demonstrated increasing fibrotic changes of subcuta­neous and synovial tissue at the site of injection after appl1cation of SiC and Sb N4 powder. when compared to alumina. This was not l1mited to the area particle deposition. but involved the neighbouring tissue (Fig.5a). A high number of histiocytes. fibroblasts and round cells characterized the tissue reaction.

Independ of material a high number of ceramic grains were found in paraaortic lymphatic nodes. After initial inflammatory reaction SiC powder particles induced interstitial fibrosis faster than Sb N4 particles. This reaction was not seen with alumina powder particles.

A varying amount of ceramic particles was found in the spleen indepen­dent of the material used. After 52 weeks a lot of grains without patho­logical tissue reactions could be observed.

376

The appearanee ot eeramie particles in the llver was irregularly and minor to the upper mentloned organs, never Induclng pathologieal tissue reaetlon.

In the lung eeramle part1cles ot small amount eould be seen interstltlal in lymph traets, phagoeytosed in alveolar maerophages and in some eases inducing eap1llary embolism (Flg.5b) . Atter 24 and 52 weeks the lungs were tree ot eeramie toreign bodies.

Figure 5a. Sie powder particles in synovial tissue and Hotta's fat pad inducing fibrotie tissue ehanges, 52

weeks post surg., magn. app. 40x

Figure 5b. Sie powder particles phagoeytosed in alveolar macro­phages, 8 weeks post surg.,

magn. app.320x

At no time ceramie part1cles could be found in the kidneys . No malignant tlssue changes were observed in any of the analyzed

organs.

DISCUSSION

Both experiments demonstrated striking differenees between tested materials with inferior blological behaviour of Sie and Sb N4 ceramics when compared to alumina ceramics.

Our results concerning osseointegration of Si3N4 are similar to those of Rether (5), who also compared this material to alumina ceramic . Howlett et al.(4) described good tissue ingrowth of a porous Sb N4 ceramic in an in vitro and in vivo study. His experiment on the other hand has the lack of not testing a material of well known biological properties for reference under the same conditlons. The tissue reaction to Sb N4 powder was unexplored until now.

In respect to bone ingrowth Sb N4 ceramic showed discret better results compared to Sie ceramic in our experiment. Further information about bone ingrowth of this material is not available. Sie powder until now only was tested in comparison to graphite (1). Striking dlfferences of tissue response to both tested silicon ceramics were not visible .

377

The results of our study clearely demonstrate the blologlcal superlorlty of alumina ceramic. Even If sillconceramics will proove to have better mechanlcal properties, our data suggest, that in respect to thelr mlnor blocompatibllly the appl1cation ,of the tested slllconceramics as biomaterials cannot be recommended.

REPERENCES

1. Alexnat, H., Gewebereaktion auf Kohlenstoff- und Siliziumpartikel in Pulverform, Inaugural-Dissertation, Frankfurt/Main, 1980.

2. Boutin,P. and Blanquaert,D., Le frottement alumine-alumine en chirurgie de la hange. 1205 arthroplasties totales: avril 1970 - juln 1980, Rev Chir Orthop, 1981, 67, 279-287.

3. Griss,P. and Heimke,G., Blocompatiblllty of high density alumina and 1ts application in orthopedic surgery, Blocompatibillty of clinical implant materials, ed. D.F.Williams, CRC Press, Inc., Boca Raton, Florida, 1981, pp 155-198.

4. Howlett,C.R., McCarthney,E. and Ching,W., The effect of silicon nitride ceram1c on rabbit seeletal cells and tissue, Clin Orthop, 1989, 244, 293-304.

5. Rether,U., Vergleichende tierexperimentelle Untersuchung über die Gewe­beverträglichkeit von Aluminiumoxidkeramik, AISI 316 L Stahl und Siliziumnitrid, Inaugural-Disserteation, Heidelberg, 1979.

6. Willert,H.-G., Semlitsch,M., Buchhorn,G. and Krlete,U., Materialverschleiß und Gewebereaktion bei künstlichen Gelenken, Orthopäde, 1978, 7, 62-83.

7. Willert,H.-G., Buchhorn,G. and Ungethüm,M., Proposed guldeline for the biological testlng of orthopaedic implant materials and Implants, Biomaterials, 1980, I, 197-182.

8. Ungethüm, M. and Blömer,W., Technologie der zementlosen Hüftendoprothetik, Orthopäde, 1987, 1.&, 170-184.

378

EFFECT OF Ti02 CERAMIC PRECURSORS ON HUMAN LYMPHOCYTE MITOGENESIS.

S. PIANTELLI., G. MACCAURO. P. FASSINA*. A. BUKAT** Istituto di Ortopedia, Univ. Cattolica S.Cuore,Roma.Italy; *TEMAV .• Medicina.

Bologna. ltaly; ** ENEA, Casaccia, Roma. ltaly.

ABSTRACT

In this paper we re port the effects of 5 different batches of Ti02 ceramic precursors on the phytohemoaggtutinin (PHA) induced proliferation of human mononuetear eells from peripherat blood. Our resutts suggest that the size of the preeursor partieles is inversely corretated with the tolicity. In addition, among the various impurities present in the dirrerent batehes. Pb eontamlnation is important in determining the tOlieity of a given titania bateh.

INTRODUCTION

Immunocompetent eells constitute an important part of the general framework of response to prosthetic implants. In particular. Iymphocytes are obviously involved in the cell-mediated hypersensilivity reaelion which develops in response to metal implants. On the other hand the impairment of the immuno-response at the implant-bone interface could be responsible for the implant failure due to infections. The lectin induced proliferation of human lymphocytes. an in vitro model of the immune response, mayaiso be employed to quanlitatively analyze the effect on cell activation of metal ions or powdered biomaterial precursors ( 1-4).

379

MATERIALS AHD METHODS

Human mononuclear cells from peripheral blood were stimulated with phytohemagglutinin as previously reported (5). Briefly, cells were cultured in Microtiter plates ( Microtest 11 Falcon) at a concentratlon of I x I OS cells I weil. Each weil contained 0.2 ml RPMI 1640 medium (GIBCO) supplemented with 2 mM glutamine, 60IJg/ml gentamicin and 10 I heat-inactivated fetal calf serum. After 72 hours of incubation in the presence of optimal phytohemagglutinin (PHA, Welcome) concentrations (0.6 IJg/ml) the cells were pulsed with 0.5 IJCi/well of (3H]-methyl thymidine (Amersham; specific activity: 5Ci/mmol) for 6 hours. At the end of the incubation period, the cells were harvested with a semiautomated cell harvester (Titertek, Flow) and the radioactivity determined by scintillation spectrometry. Powder sampies of TiOz ceramic precursors were prepared from 5 different batches (TEMAV Spa, Medicina, Bologna). These powders were gamma-ray sterilized, suspended in RPMI culture medium and added to the culture weHs at final concentrations ranging from 0.6 to 6 mg/mI.

RESUL TS AHD DISCUSSION

As reported in Table I, TiOz ceramic precursors exerted a dose-dependent inhibition of the PHA-induced Iymphocyte proliferation.

TABLE I Inhibitory effect of various doses ofTi02 powders of different grain sizes on

PHA induced human Iymphocyte proliferation

Ti02 doses (g/l 00 mD Grain size(IJm) 0.06 0.15 0.3' 0.6

TI <38 81 ± 9(a 69± 10 40±4 3±2 TI 38-106 91±11 83± 8 48±3 6±4 TI 106-150 79±12 67± 7 5l±5 8±2 T3106-150 95±15 81± 9 45±6 4±3 T4 < 38 65 t 7 45:t 5 14 ± 3 3 ± I T4106-150 80tl0 75tl0 38±6 4±2 T5 38-106 82 ± 9 53 ± 7 10 ± 2 4 ± 2 T5106-150 79± 8 56± 6 17±5 5±3 T6 <38 85 ± 12 76 ± 11 24 ± 3 3 ± 1 T6 106-150 93 t 13 71 t 9 33 t 6 5 t 4

a) I of control; mean t S.O .. ; N-6

380

An almost complete inhibition of the proliferative responses of lymphocytes was obtained by the addition to the cultures of aJJ ceramic precursors tested, at a concentration of 0.6 mg/mt. Low inhibitory activity was exhibited by the same precursors at 0.06 mg/mt. From these dose-response experiments it was possible to calculate the inhibitory dose 501 UD 50) for alJ the powdered precursors tested (Figure 1).

Tl02:

TI6 106-150 TI6 <38

TI 5 106-150 TI 5 38-106

TI4 106-150 TI 4 <38

T13106-150 Tll 106-150 TII 38-106

TI 1 < 38

0,0

mean +/- S.o. N = 6

... ,.. ... ,.. ~ ,.. ... ,..

~ ,.. ~ ,..

~ ,..-.... ,..

~ ,.. ~ r

0,1 0,2 0,3 0,4 Inhlbltory dose (% w Iv) SO~

Figure I. Inhibitory effects of Ti02 powders on PHA induced human Iymphocyte proliferation. Categories indicate the size range (J-Im) of the

powder grains.

The size of the partieles was inversely correlated with the to!icity, the <38 microns partieles being more inhibitory than the >38 ones. Precursor particles of the same dimensions belonging to the 5 different batches exhibited different inhibitory activities on Iymphocyte mltogenesls (Figure 2). This finding depended on the amount of Pb impurity as suggested by the positive correlations between the relative Pb concentrations in the precursor partieles and the entity of the inhibitory effect (figure 3).

381

Ti02: mean +/- S.o. N = 6

TI6 106-150

TI5 106- 150

Tl 4 106-150

T13106-150

Tl 1 106- 150 ~~~~~~~~

0,0 0,1 0,2 0,3 0,4

Inhtbttory dose (~w Iv) 50% Figure 2. Inhibitory effects of five different Ti02 powders with the same

grain size on PHA induced human Iymphocyte proliferation.

220 r = 0.92 ; p< 0.05

200

-00 180 ..... ~ 160 .......

Co 140

000

120 0,1 0,2 0,3 0,4

Inhlbltory dose (~w Iv) 50~

Figure 3. Correlation between Ti02 powder inhibitory capacities on PHA induced human Iymphocyte proliferation and their Pb impurity concentratlons. The size of the grains for all the five Ti02 powders analysed

were 106-150 J,Jm.

382

The 5 different Titania batches also differed in their content of other impurities such as AI. Ca. Fe, Hg. Mg. Mn. Sn, Zn, y, Si. However, no significant correlation was found between these differencies and the correspondiog inhibitory actlvity on mitogenesis (data not shown). The different specific surfaces of the 5 ceramic precursors (from 102 m2/g to 117 m2/g) did not influence the inhibitory actlvity on mitogenesis.

REFERENCES

1. Piantelli, S., Tranquilli-Leali, P., Lorini, G. and Piantelli, M., Effetti inibitori di alcuni metalli di interesse ortopedico sulla trasformazione blastica dei linfociti umani. Min. Ort.. 1983, 34f, 1-3.

2. Panush, R.S., Petty R.W., Inhibition oe human 1ymphocyte responses by methy1metacrylate. eUn. OrthQP. ReJ. Res .. 1987, 134f, 356-63.

3. Pizzoferrato, A., Vespucci, A., Ciappetti, G. and Stea, 5 .• Biocompatibility testing oe prosthetic imp1ant materials by cell cultures; Biomaterials. 1985. 6,346-51.

4. Itoh,S., Ishibashi. K .• Inada. K. and Yoshida •. M .• The effects oe bioactive glass on the function of 1ymphocytes. In: Handbook of Bioactive Ceramics. I ,ed Yamamuro. T .• Hench. LL. and WUson. J., CRC Press Boca Raton. USA. 1990. pp227-33.

5. Piantelli, M., Lauriola. L .• Maggiano. N .• Ranelletti, F.O. and Musiani, P., R01e oe interleukins 1 and 2 on human thymocyte mitogen activation. CdI.. ImmunoL 1981, 64f. 337-42.

383

EXPERIMENT AL STUDY OF TWO CORALS USED AS BONE IMPLANT IN THE SHEEP.

P. JAMMET, F. BONNEL, P. BALDET, F. SOUYRIS, M. HUGUET Faculty of Medicine - Montpellier - FRANCE

ABSTRACT

Goniopora Lobata and Polyphyllia Talpina were tested on ten sheep with a follow-up of 12 months. The material of reference is a vitroceramic CAP 42. Different implantation sites were used : craniofacial and orthopedic. A macro and microscopic postoperative study was performed at 3, 6, 9 and 12 months. The resul ts show :

- biocompatibility with bone, - a disappearence, without colonisation, of Goniopora Lobata associated

with a complete bone regeneration at 12 months. - a later resorption of Polyphyllia Talpina with a fibrous colonisation,

without osseous transformation. The possibility of using these materials in Cranio-Maxillo-Facial surgery are discussed.

INTRODUCTION

In the department of maxillofacial surgery of Montpellier we have used the madreporaria skeletons for 10 years as bone implants. In the same way, GUILLEMIN (3), PATAT (4) and CHIROFF (2) studied the bone ingrowth into a porous coral skeleton. They used the Porites variety. This experimental study proposed to analyse the involvement of two new varieties of coral from the Great Barrier Reef and the Coral sea.

MATERIAL AND METHODS

Goniopora Lobata and Polyphyllia Talpina were tested on ten sheep. The general structure of Goniopora Lobata is quite different from Polyphyllia Talpina because of the ingrowth of the coral's colonies under the water. Goniopora Lobata has an open porosity in contrary to Polyphyllia Talpina which has a closed porosity. The structure of Goniopora Lobata is similar to Porites (3). The difference is a greater porosity (70%/50%). All of them consist of calcium carbonate. We proposed to test them as bone substitute in the sheep with an aluminophosphate glass-ceramic of reference (porosity ~ 60 %) (I) (5).

384

Implantation

The natural structure of Goniopora Lobata permitted the manufacture of cylindrical implants. Polyphyllia Talpina permitted only rectangular implants. The implantation sides used were :

- the cranial vault, the nasal area, the mandibule for Polyphyllia Talpina - the femur and the posterior iliac crest for Goniopora Lobata.

120 corals were implanted. 70 of them analysed. 20 for Polyphyllia Talpina and 50 for Goniopora Lobata. A macroscopic and a microscopic study were performed. Qualitative and quantitative ingrowth measuring bolt.

RESULTS

Macroscopic study

During the follow up, we described fractures in 8 of the 10 femurs implanted. Every Goniopora Lobata implant tested disappeared from 3 months postoperative. Every time cortical restoration associated with no periphereal tissue reaction was found. In the ca se of Polyphyllia's implantation the apparent volume was conserved for 12 months. No periphereal tissue reaction was observed.

Microscopic study

The qualitative microscopic parameters studied were the implant resorption, the interface bone/implant and the bone and fibrous connective tissue ingrowth at 3, 6, 9 and 12 months postoperative.

The rating system assigned : o in ca se of no bone or fibrous ingrowth + in ca se of ingrowth of immature bone trabeculae +++ in ca se of differentiation of lamellar bone ++ inter mediate between + and +++

TABLE 1 Qualitative microscopic study-results (Goniopora) (0 to 3)

Implant Bone Fibrous Months resorption Interface ingrowth ingrowth

3 +++ 0 0 0 6 +++ 0 0 0 9 +++ 0 ++ 0

12 +++ 0 +++ 0

For Goniopora Lobata (Table 1) we found a total resorption after 3 months associated with a progressive bone ingrowth. At 12 months the bone is completely regenerated. Goniopora Lobata disappeared without real colonisation.

385

TABLE 2 Qualitative microscopic study-results (Polyphyllia) (0 to 3)

Months Resorption

12 +

Bone Interface ingrowth

o 0

Fibrous ingrowth

+++

For Polyphyllia Talpina (Table 2) we found later implant resorption associated with a fibrous ingrowth without osseous transformation (Fig.l - Fig.2).

TABLE 3 Qualitative microscopic study-results (Ceramics) (0 to 3)

Months Resorption

12 +

Bone Interface ingrowth

+++ ++

Fibrous ingrowth

+

In the case of ceramic implantation (Table 3) the resorption at 12 months is low (l %). We found the same results as Pernot (5). They showed at 12 months fibrous ingrowth with osseous transformation.

The quantitative microscopic parameters were the implant's matrix, the tissue component and the depth of the cortical bone (Table 4).

12 months

Polyphyllia Goniopora Ceramic

TABLE 4 Quantitative microscopic study-results (0 to 3)

Matrix (%)

95 o

41

Tissue component (%)

5 100 59

Cortical bone (mm)

o 1,78

o

386

HISTOLOGICAL ASPECTS AFTER POLYPHYLLlA T ALPINA IMPLANT A nON

(I2 MONTHS POST OPERA nVE)

FIGURE 1

MANDIßLE IMPLANT A nON

1. Coral skeletton 2. Masseter 3. Mandible bone ~. Fibrous tissue

FIGURE 2

NASAL IMPLANTATION

1. Coral skeletton 2. Nasal bone 3. Fibrous tissue

FIGURE 3

NASAL IMPLANT A nON MACROSCOPIC VIEW

1. Coral

387

At 12 months we observed : - very little matrix disappearenee, no eortical bone formation for

Polyphyllia Talpina. - eomplete matrix disappearenee, bone regeneration (cortical and spongious)

for Goniopora Lobata. - little resorption, fibrous and osseous eolonisation for the eeramic.

IN CONCLUSION

We think that Goniopora Lobata has a too high a porosity. That explains perhaps the quicker resorption than with Porites (3). Polyphyllia Talpina seems interesting beeause of its very slow resorption. However, no osseous eolonisation was no ted beeause of its closed porosity. We are testing sueeessfully this last sort of eoral on the human body in eraniomaxillofacial surgery (6).

REFERENCES

1. BALDET, P., PERNOT, F., ZARZYCKI, J., BONNEL, F., RABISCHONG, P., Study of Bone ingrowth in porous calcium alumino-phosphate Glass eeramics. Biomaterials, 1980, pp. 73-85.

2. CHIROFF, R., Tissue Ingrowth of Replamineform Implants. J. Biomed. Mater. Res. Symposium, 1975, 6, pp. 29-~5.

3. GUILLEMIN, G., FOURNIE, J., PATAT, J.L., CHETAIL, M., Contribution a l'etude du devenir d'un fragment de squelette de eorail madreporaire im plante dans la diaphyse des os longs ehez le ehien. C.R. Aead. Sc. Paris, 1981, 293, pp. 371-5.

~. PATAT, J.L., GUILLEMIN G., Le eorail naturel utilise eomme biomateriau de substitution a la greffe osseuse. Ann. Chir. Plast. Esthet., 1989, 34, pp. 221-5.

5. PERNOT, F., BALDET, P., BONNEL, F., ZARZYCKI, J., RABISCHONG, P., Development of Phosphate Glass-Ceramics for Bone Implants. Ceramics International, 1983, 9, pp. 127-131.

6. SOUYRIS, F., CHEVALIER, J.P., PAYROT, Cl., PELLEQUER, Cl., GARY BOBO, A., MERLIER, Ch., Bilan apres quatre ans d'experimentation du eorail a titre d'implant osseux. Ann. Chir. Plast. Esthet., 198~, 29, pp. 256-60.

7. SOUYRIS F., Utilisation des Madreporaires a titre d'implant osseux en Chirurgie Maxillo-Faeiale et Cranio-Faciale. Rev. Chir. Orthop., 1983, 69, pp. 577.

388

INTERFACIAl STUDY OF SOME INERT AND ACTIVE CERAMICS IMPlANTED IN BONE.

D. lAFFE, *S. GIANNINI, *A. MORONI, °A. KRAJEWSKI and °A. RAVAGLIOLI

Institute of Human Anatomy, Via del Pozzo 71, 41100 MODENA, Italy *Orthopaedic School, Univ. of Bologna, Rizzoli Inst., Italy

°IRTEC-CNR, Faenza (RA), Italy

ABSTRACT

Cylinders of biologically inert (Alumina and lirconia) and active (Hydroxyapatite, Bioglass and AKRA 15 glass) ceramics, were inserted as a press-fit for 6 and 12 months in the femurs of adult Merinos sheep. It resulted that all the cylinders were surrounded by an envelope of new cortical bone; however those made up of inert ceramic were separated from the bone by a capsule of loose connective tissue, whereas a direct contact occurred between the new formed bone and the active ceramics cylinders, except for AKRA 15 rods whose surfaces appeared to be almost completely covered by fibrous tissue. A bidirectional flow of ions occurred in the outer reactive layer of the two glasses examined. The results are discussed in relation with the biological behaviour of the tested materials in bone implants.

I NTRODUCTI ON

The materials used in skeletal implantology have various behaviours and

induce different biological reactions when implanted in vivo. They can be

roughly classified in three categories: a) biotolerant materials (PMMA,

Polyethylene, Steels and most part of alloys), b) bioinert materials

(Alumina, lirconia, Carbon, Titanium) and c) bioactive materials (Ca-P

based ceramics, Ca-P containing glasses, glass-ceramics) (1). In the first

389

group, the metals are the most interesting type of materials, though they

undergo some alterations, as for instance the corrosion involving ions

release that affect normal cell differentiation and proliferation (l). To

avoid this, ceramic coating of metals have been used in order to impart to

the metal surface the characteristics of the ceramics used. Since the

components of ceramics, such as Alumina or Zirconia, are shown to be not

solubles in organic fluids they should not influence the organic tissues

(l). However such materials do not bind to bone {2}. On the contrary,

bioactive ceramics, with respect to inert one, give a lesser degree of

protection to the metals, but show a greater degree of biocompatibility

(3), though the mechanism of their binding to bone is not yet completely

understood (l).

The aim of the present investigation is to bring additional

morphological and chemico-physical data on the behaviour of some inert or

bioactive ceramics implanted in bone sheep.

MATERIALS AND METHODS

On the whole, 40 cylinders (5x12 mm) of the following ceramics were

implanted in the femurs of four adult Merino sheep: Zirconia (Zr02),

Alumina (A1 203), Hydroxyapatite (HA) (Ca50H{P04)3), Bioglass (same 45S5

formulation (4}), AKRA 15 (a new bioactive glass (5». Additionally, 6

Alumina cylinder, having the same dimensions as those mentioned above but

whose convex lateral surface was furrowed by circular grooves, were

implanted in two sheep. The ceramic prostheses were prepared by IRTEC-CNR

(Faenza, Italy).

In each femur of the right and left side, the cylinders were

implanted as a press-fit inside five transcortical holes drilled at

diaphyseal level, 2 cm apart from each other. The grooved Alumina

cylinders were inserted at the metaphyseal level. The cylinders were

pressed inside the hole until their external flat surface was level with

390

the periosteal surface of the shaft. The two sheep, with implants of flat

cylinders only, were sacrificed after 6 months from the implant, and the

remaining two after 12 months. The bone samples containing the implants

were fixed, dehydrated and embedded in methylmethacrylate resin. Cross

sections (500 micron thick), taken from the levels of the implants by

means of a diamond saw microtome (mod. 1600 - Leitz, Germany), were

perfectly polished with emery paper and Alumina. SEM observations (SEM 500

- Philips, The Netherlands) and X-rays Energy Dispersion System

microanalyses (EDAX - Philips International) were performed on such

polished thick sections after sputtering their surface with a thin

conductive layer of carbon.

RESUlTS

No significant differences were observed between Zirconia and

Alumina cylinders. After 6 months (Fig . s la and 2a), the bone surrounds

both their convex lateral surface and also the flat surface, at their

external end . A layer of loose connective tissue, whose thickness ranges

between 50 and 150 microns, intervenes between the bone and the inert

Figure 1. SEM micrographs of Zirconia cylinders (Zr) for 6 (a) and 12 months (b and c) in sheep's femur. (Bars: a band c = O. 1 RIß) . --

implanted = 1 nlll;

391

Figure 2. SEM micrographs of A1umina cy1inders (Al) imp1anted for 6 (a) and 12 months (~and~) in sheep's femur. (Bars : a and c = 1 mm; b 0 . 1 mm) .

Figure 3. SEM micrographs of Hydroxyapatite cy1inders (HA) imp1anted for 12 months in sheep's femur. (Bars: a = 1 mm; band c = 0.1 mm).

ceramic . After 12 months, the fibrous 1ayer appears to be invaded by new1y

formed trabecular bone (Fig. 1b and 2b) that comes into contact with the

surface of the cylinders. The newformed bone also grows inside the grooves

on the surface of the A1umina cy1inders (Fig. 2c). It is to note, however,

that at the externa1 edges of the cy1inders, fibrous tissue usua11y

intervenes between the bone and the imp1ant (Fig. 1c).

Severa1 large vascu1ar cavities (0 . 5xO.1 mm or more) may be observed

392

Figure 4. SEM micrographs of Bioglass cylinders (8G) implanted for 12 months in sheep's femur. (Bars: a = 1 mm; b = 0.1 mm; c = 0.01 mm).

Figure 5. SEM micrographs of AKRA 15 cylinders (AK) implanted for 12 months in sheep's femur . (Bars: a = 1 mm ; b = 0.1 10m; c = 0.01 mm).

in the bone adjacent to the surface of Hydroxyapatite cylinders. After 12

months (Fig. 3a), the cavities increase in number but decrease in size,

and the surface of the cylinders shows many indentation completely filled

by new bone (Fig. 3b), particularly opposite of bone vascular cavities

(Fig. 3c).

No differences with time were observed between Bioglass and AKRA 15

cylinders. The two glasses show at their periphery a degraded layer, whose

393

Figure 6. Digital scanning electron micrograph (se) and digital X-ray dot maps for Sodium (Na), Silicon (Si), ?hosphorus (P), Potassium (!), Calcium (Ca) and Iron (Fe) recorded in the intramedullary segment of an AKRA 15 cylinder. Sheep's femur 12 months after the implant. NDG = Ilot-degraded glass; Dl = degraded layer (Bar = 0.1 mm).

thickness ranges between 100 and 200 microns

(Fig.s 4a and 5a). The bone comes into contact

with the outer intracortical segment of

Bioglass cylinders (Fig. 4a) and, after 12 months, also envelops the

intramedullary segment close to the endosteal surface (Fig. 4c). Contrary

to HA cylinders, large vascular cavities are lacking in the bone adjacent

394

to the Biog1ass; this bone contains globous osteocyte 1acunae, typica1 of

woven bone (Fig. 4b). The AKRA 15 cy1inders are most1y separated from the

bone by a thin 1ayer (about 30 micron thick) of fibrous tissue (Fig. 5a).

Sometimes the thickness of this 1ayer great1y reduces in size (Fig. 5b) or

it may even disappears 1eaving the glass in contact with the bone

(Fig. 5c).

X-ray analysis shows that in the degraded 1ayer (OL) of both

Biog1ass and AKRA 15, the so-ca11ed "Silicon barrier" either has an

irregu1ar thickness or is 1acking (Fig. 6). Furthermore, the periphera1

Ca-rich 1ayer may not be present, particu1ar1y in the intramedu11ary

segment of cy1inders. Topographic X-ray analysis carried out on the

intramedu11ary AKRA 15 cy1inder (Fig. 6) shows that: 1) Sodium is present

in the not degraded glass on1y (NDG); 2) Silicon forms an irregu1ar

barrier, far from the NDG; 3) Phosphorus gradua11y increases in the DL, on

going towards the bio10gica1 fluids; 4) Potassium, an element absent in

the composition of the glass, is present in all parts of DL, particu1ar1y

c10se to the NDG; 5) Calcium is almost absent in the DL near the NDG, but

is present at the periphery, though in a 10wer content; 6) Iron is present

both in the NDG and in the DL, but in the 1atter its content is higher.

Quantitative X-ray analysis, using not imp1anted glass as standard, shows

that from the NDG to DL: P increases from 2.3% to 11%, Ca/P ratio (wt/wt)

decreases from 6.9 to 1.3 and Fe increases from 5% to 8-9%.

DISCUSSION AND CONCLUSIONS

On the basis of the resu1ts here reported, it appears that "contact

osteogenesis" does not occurs throughout the surface of the inert ceramic

imp1ants, as suggested by Osborn (6). In fact, it was not observed in our

experiment a10ng the externa1 edges of the Zirconia and A1umina cy1inders.

Since in these regions mechanica1 strains are presumab1y higher, it may be

inferred that an inverse corre1ation might exists between osteointegration

395

and mechanica1 forces. Thus tests performed on un10aded imp1ants of inert

ceramics do not seem to give tenab1e informations. The observed

degradation of Hydroxyapatite casts some doubts on the stabi1ity of this

material in bio10gica1 fluids. Ana1yses of active glasses show that a

diffusion of the main glass components occurs toward the bio10gica1

environment; not on1y P and Ca enter the periphera1 1ayer of DG , but also

a K and H enter the glass probab1y to counterbalance its 10ss of elements.

The increment of not-diffusible ions (Fe, Al) probab1y depends on the

vo1umetric shrinkage of the degraded layer. One can speculate that the

latter phenomenon, together with the lower diffusion rate of Si with

respect to Ca and Na, may be the cause of the Si-rich layer (i.e. the

so-called Silicon barrier) in the inner part of the DG. However, further

investigations are needed to a better understanding of the biological

reaction to bioceramic implants.

REFERENCES

1. Heimke, G., Biomaterials highlights 111, Bone replacements: 1mplant materials and the mode of implant fixation. Adv. Mater., 1989, 10, 345-348.

2. Ono, K., Yamamuro, T., Nakamura, T. and KOkubo, study on osteoconduction of apatite-wollastonite ceramic granules, hydroxyapatite granules, and Biomaterials, 1990, 11, 265-271.

T., Quantitative containing glass

alumina granules.

3. Gross, U., Kinne, R., Schmitz, H.J. and Strunz, V., The response of bone to surface-active glasses/glass-ceramics. CRC Crit. Rev. Biocompatib., 1988, 4, 155-179.

4. Hench, L.L. and Ethridge, E.C., Biomaterials, an interfacial approach, Academic Press, New York, 1982.

5. Ravaglioli, A., Krajewski, A., Zini, Q. and Venturi, long range silicate release from doped bioglass into Technical note. Biomaterials, 1986, 7, 76-78.

R., Short and 199 medium.

6. Osborn, J.F., 1mplant material hydroxylapatite ceramic. Basic considerations and clinical applications, Quintessenz-Verlag. Berlin, 1985.

396

SURFACE CHARGE OF THE BIOGLASS TREATED BY A PHYSIOLOGICAL SOLUTION

STANISLAWA SZARSKA Institute of Physics, Technical University of Wroclaw,

SO-370 Wroclaw, Wybrzeze Wyspianskiego 27

ABSTRACT In result of interaction of the simulated physiological soluti­on with bioglass, at the surface the layer is produced, enri­ched with Si02 and calcium phosphate. The sUbject of investi­gation using the OSEE method was the surface charge appearing on the glass surface in an aqueous environment. The results of measurements have been compared with the surface charge on the porous silicate glasses.

INTRODUCTION When the bioglass is immersed in the simulated physiological

solution, the sodium ions are leached from the surface and

are replaced by H+ (H 30+) ions from the solution through an

ion-exchange reaction producing a silica-rich surface layer.

(1) On this layer, a caO-P 20 S-rich film is formed. This film

in contact with the bone surface makes it possible to the

osteogenesis process. This ion-exchange process gives in re­

sult a negatively charged surface (2). Li and Zhang reported

that a charged surface is developed, leading to the formation

of an electric double layer in the glass-solution system.

The ion distribution inside the electric double layer is dif­

ferent from that in the bulk of sOlution and a protective

layer is formed. They had stated that the charged surface is

either positive or negative depending on the glass composition

and of the pH solution. As the charged surface plays an impor­

tant role in the reaction of the glass with the solution, the

optical stimulated exoelectron emission (OSEE) was applied to

determine the degree of charging of the bioglass surface.

397

Application of the OSEE method for investigations of the

glass surface is reported in references (3-6). Exoemission

method for biomaterials research was applied by J.E.Davies(7).

Exoelectron emission as a phy.sical effect,is a non-stationary

phenomenon occuring on the surface which is in a noneqilibra­

ted state.A surface layer of nonmetallic solids was examined

using the exoelectron emission method. Low energy electrons

(exoelectrons) are emitted from the surface layer during

optical or thermal stimulation caused by excitation mechanical,

chemical or ionizing radiation of the surface.(8)

EXPERIMENTAL ARRANGEMENT AND PREPARATION OF THE SAMPLES

The measurements were performed in vacuum of 10-4 Pa,at about

300 K.The emitted electrons were focussed by the electrostatic

field and registered by the electron mUltiplier and recorder

(5).Because the sterilization of the sample was performed by

exposure to the ultraviolet light (9),the OSEE measurements

were stimulated using a deuter lamp with an interference fil­

ter with the wavelenght 325 nm.

The subject of investigations were:the 45S5 bioglass(10)

and the porous silicate glas ses obtained by K.Marczuk with

90% Si02 and 10% B203' The samples were exposed to a simulated

physiological sOlution with pH equal to 7.4 from 1 h to 2

weeks and for 2 hours in the sOlution with pH equal to 6~9-7.4

RESULT OF MEASOREMENTS AND DISCUSSION

The purpose of this paper was to discuss the possibility of

application of the OSEE method to characterize the dynamic

surface charge activity of the bioglass.Two types of dependen­

ce of the OSEE intensity on time can be observed. The first

one is the decay curve which gives information on the energe­

tic spectrum of the electron traps in the material(4,5).

Whereas for the bioglass and the porous glasses, the shapes

of the curves presented in fig. 1 are observed. The ealier

investigation (11) pointed out that this shape of OSEE curve

398

is connected with the surface charge of the sampIe.

1 so

! ~o ~ .... 30 ~ ~ 20 .~

~ 10 ~

2 6 B 10 12 U

Figure 1. The dependence of the OSEE intensity on time after

exposed to a physiological solution with pH equal 7.4 for 2 h.

The area under the OSEE curve marked on fig.1 is proportional

to the degree of charging Sc of the bioglass surface. In fig.2

is presented the dependence of the surfar.e charge Sc on the

time of exposure of the sampIe in the physiological solution

of the bioglass and porous glass.The shape of the curve points

out for the violent charging process of the surface when the

bioglass is exposed for the simulated physiological solution

till 2 hours. After 5 days the surface charge reaches its

maximal value. Kim at all. (12) stated that the concentration

of the phosphorus ions in the reacted sOlution drops after

2 hours. This phenomenon can be explained as crystallization

of the amorphous calcium-phosphate layer at the surface.

This process is connected with a change of the bond orienta­

tion and therefore with releasing of electrons and increase

in electron exoemission. The increase in the negative poten­

tial is connected with enrichment of the Ca 2+ ions in the

electric double layer with the time of exposure in the physio-

399

logical SOlution.

1 8 e e e • •

6 .-------.( I c"

CI) 2

S 7 9 11 13

hours tloys

Figure 2. The dependence of the surface charge Sc on time of

exposure in the simulated physiological sOlution with pH equal

7.4;0 bioglass;eporous glass.

In fig.3 is presented the dependence of the surface charge Sc versus pH of the solution for the 2 hours exposure in the

simulated physiological sOlution.

" -0

f

0 0

3 - 0

~ 2 -

0

1 0

• I I 1 I • • 6,9 7,0 7,1 7,2 1,3 1,.( 1,S

pH •

Figure 3. The surface charge Sc versus pH of the physiological

sOlution (2 hexposure).

400

The shape of the curve indicates that the value of the surface

charge increases when the solution is more alkaline. This

result agrees with the theoretical predictions of Li and Zhang.

They stated that the surface is charged negatively in the case

when the concentration of (=SiO-) silox group is larger than

that of (=SiOHR+) cationic group. The concentration of this

cationic group decreases with the rise of the pH value of the

sOlution. When the bioglass is immersed in the solution with

the increased pH value, Na+ ions exchange with H+. It leads

to a negative charging of the surface (2). This results in en­

richment of Na+ and Ca 2+ in the interface between the bioglass

and the sOlution. The concentration of Na+,Ca 2+ in the inter­

face is alway§ latger than thatrin the bulk.of the soiution

although they both increase gradually with the further ionic

exchange between the bioglass and the sOlution. This phenome~

non corresponds with the shap,e of the curve in fig.3.

CONCLUSION

The exoemission measurements pointed out that the surface is

emitting electrons during dynamic relaxation of excitations

and are capable to demonstrate the subtle changes of the bio­

gl ass surface, caused by the chemical pre-treatments.

REFERENCES

1. Hench,L.L., Ethridge,E.C., Biomaterials. An Interfacial Approach., Academic Press, New York, 1982, 77-78.

2. Li,p.,Zhang,F., The electrochemistry of a glass surface and its application to bioactive glass in sOlution. ~.Non­Cryst.Sol.,1990,119, 112-118.

3. Sobolew,J.W., Kortow,W.S., Kazjawina,I.W., Zatsepin,A.F., Szczeglowa,O.W., Primienienie metoda fotostimulirovannoj exoelectronnoj emisji dIa ocenki defektivnosti povierhnosti listovogo stiekla. Fizika i Chimia Stekla, 1985, 11, 4, 490-493.

4. Szarska, S., Magierski, W., Investigation of the EE from the gl ass surface obtained by ion exchange. Rad. Prot. Dosim., 1983, 4, 3/4, 201-204.

5. Szarska, S., Magierski, W., OSEE as the method of determi­ning of defect degree for a fibre preforms. SPIE Optical Fibres and Their Applications, 1986, 670, 109-111.

401

6. Szarska, S., Magierski, W., Dependence of the Parameters of OSEE from silicate glasses on the wavelenght of stimula­ting light. Acta Univer. Wrat., 1988, 1085, 54, 57-64.

7. Davies, J.E.~oemission for BiomaterialS Research.Japan. ~.~.Phys., 1985, 24, 24-30.

8. Glaefeke, H., Exoemission. In Thermally Stimulated Relaxa­tion in Solids. ed. P. Bräunlich, Springer-Verlag, Berlin, 1979,225-274.

9. Gross, U.M., Strunz, V., The anchoring of glass ceramics of different solubility in the femur of the rat. J. Biomed. Mater. Res., 1980, 14, 607-618. --

10.Hench, L.L., Paschali, H.A., Direct Chemical Bond of bioac­tive glass-ceramic materials to bone and muscle., ~.Biomed. Mater. Res. Sym., 1973, 4, 25-42.

11. Kamada, M., Asai, F., Tsutsumi, K., Optically Stimulated Exoelectron Emission from X-Irradiated KCI. Japan. ~. ~. Phys. , 1985, 24, 92-95.

12.Kim, Ch.Y., Clark, A.E., Hench, L.L., Early stages of cal­cium-phosphate layer formation in bioglasses. ~.Non-Cryst. Sol., 1989, 113, 195-202.

402

IN VIVO STUDY OF A NEW ACTIVE GLASS FOR BONE REPAIR: SHORT TERM RESULTS

A.M. GATTI,+D. ZAFFE, *0. ANDERSON School of Dentistry, Via del Pozzo 71, 41100 MODENA, Italy

+Institute of Human Anatomy, Univ. of Modena, Italy *Dpt. of Chemical Eng., Abo Akademi Univ. Turku, Finland

ABSTRACT

The "in vivo" behaviour of a new active glass, as granules, was checked for 3 months in mandibular sheep's bone. After explantation, SEM observations and X-ray microanalysis on the mandibular sections were carried out. New formed bone grew directly on the granules that show a biodegradation.

I NTRODUCT ION

Active glasses are promising materials for the repair of bone defects in

Dentistry and Orthopaedics (1,2). The present study reports the "in vivo"

behaviour of a new glass used as granules which develops a direct contact

with bone.

MATERIALS AND METHODS

A new type of bioactive glass called SS3P4 (3), composed of S3% Si02, 23%

Na20, 20% CaO, 4% P20S' was used as 300-S00 micron diameter granules in

mandibular sheep's bone. Two 4 mm diameter holes were drilled bilaterally

in two sheep"s jaws; two were filled with granules of SS3P4 and two were

403

left free in order to check the bone

growth without any "induction".

Three months after

implantation, the bone segments were

explanted , fixed in 4%

Paraformaldehyde, embedded in

IIlethylmethacrylate resin, thin-

-sectioned with a sawing rnicrotome

(Lei tz, Germany) and surface-

-polished with alumina . The

non-decalcified sections were coated

with a carbon fil rn and exami ned

under a scanning electron microscope

(SEM 500, Philips, the Netherlands)

with an X-ray microprobe (EDS) for

the mi croana lysi s . Non-illlp 1 anted

granules were elllbedded and prepared

in the sallie way to analyze the

morphology and the elernental

composition of the ori ginal glass.

RESULTS AND DISCUSSION

The SEM st udy of the morphology of

rlgure 1. Digital scanning electron Illicrograph (se) and X-ray analyses for Sodiurn --(Na) Silicon (Si), Phosphorus (~) and Calcium (Ca)--of non-implanted glass granules (Fieldwidth = 1.4 mm).

404

the non-implanted glass (Fig. 1)

proves that the material is

homogenous and the topographie X-ray

microanalyses for Sodium, Silicon,

Phosphorus and Calcium (Fig. 1),

show that these elements are

homogeneously distributed inside the

granules.

After 3 months from

implantation, the newly-grown bone

closes the surgically drilled holes

almost completely; the amount of

this new bone is greater than that

of the reference holes.Some granules

at the cent re of the pocket appear

to be surrounded by new formed bony

trabeculae and others by connective

tissue (Fig. 2). All granules show a

degradation in the outer part; this

degradation consists of the

appearance of white lines of

fracture crossing the outer part of

the granule and of the setting up of

concentration gradients . In granules

completely surrounded by bone, no

Figure 2. Digital scanning electron micrograph (se) and X-ray analyses for Sodium ~(Na) Silicon (Si), Phosphorus (P)--and Calcium (Ca) of implanted glass granules. surrounded by connective tissue (Fieldwidth = 1.4 mm).

405

fibrous tissue is observed in

between. The fracture lines,

crossing the external layer

sometimes propagate into the

surrounding bone. Probably, this

means that the brittleness of the

glass is increased and it is like

bone's. From a morphological point

of view, this part is very similar

to the bone and only very-high

magnification SEM observations allow

to distinguish the two materials.

The topographie X-ray

analyses (Fig. 3) show that : a) the

core of the granule is not degraded

and its composition (Fig.s 3 and 4a)

corresponds to the original one

(i.e. 17.1 wt% of Sodium,

24 .7 wt% of Silicon, 1.7 wt% of

Phosphorus and 14.3 wt% of Calcium);

b) the granule interface can be

subdivided into two layers,

different for elemental composition.

The inner 1 ayer i s mai n"iy composed

of Silicon (Fig.s 3, 4b and 4c) and

the outer layer mainly of Calcium

Figure 3. Digital scanning electron micrograph (se) and X-ray analyses for Sodium -(Na) Silicon (Si), Phosphorus (P) and Calcium (Ca) of a glass granule completely surrounded by new bone (Fieldwidth = 0.35 mm).

406

8.4 2.1 ....

a b .. . ... ... . ' . .,.

1.3 1.6 '1-.

c ....

c.,. e d ....

1.7 1.9

....

e f" f ... ... ..

Figure 4. Spot X-ray microanalyses carried out in the points indicated in Fig.3-se. The Ca/P ratio of each analysis is reported at the top-right side.

and Phosphorus. Sodium is absent in both layers. In these two layers

there are different behaviours of Calcium and Phosphorus concentrations;

the Calcium content in the inner layer appears to be decreased, whereas

the Phosphorus content is sirnilar to that of the core. These contents

gradually increase in the outer layer, reaching their highest values near

bone. So the Ca/P ratio reaches a minimum value in the inner layer close

to the outer one (Fig.s 3 and 4c). No Silicon is detected in the bone in

contact with the glass (Fig.s 3 and 4f).

407

CONClUSIONS

This new material seems promising for clinical applications since it

develops a very close contact with bone. This is due to the biodegradation

of the glass and the formation of a Ca-P rich layer at the periphery of

the granules. Nevertheless, the results show that there is a diffusion of

Calcium, Silicon and Sodium toward the biological environment and an

inverse diffusion of Calcium and Phosphorus toward the glass. Probably,

this is the cause of the leaking of Silicon in the granules surrounded by

fibrous tissue; on the contrary, in the granules surrounded by bone, the

degradation is stopped by the formation of bone itself (4).

REFERENCES

1. Hench, L.L. and Ethridge, E.C., Biomaterials, an interfacial approach, Academic press, New york, 1982.

2. Gross, U., Kinne, R., Schmitz, H.J. and Strunz, V., The response of bone to surface-active glasses/glass-ceramics. CRC Crit. Rev. Biocompatibil., 1988, 4, 155-179.

3. Andersson, 0., The bioactivity of Silicate glass. Ph.D. Thesis, Dpt. of Chem.Eng., Abo Akademi Univ., Turku, Finland, 1990.

4. Gatti, A.M. and Zaffe, D., Long term behaviour of active glasses in sheep's mandibular bone. Biomaterials, 1991, in press.

408

A STUDY OF HYDROXYAPATITE CERAMICS AND ITS OSTEOGENESIS

ZHANG XINGDONG, ZHOU PIN, ZHANG JIANGUO, CHEN WEIQUN, WU CHUONG

Institute of Materials Science and Technology, Sichuan University

Chengdu 610064, P.R. China(ll

ABSTRACT

A porous hydroxyapati te ceramic block wi th some tricalcium phosphate was prepared. The in vive experiments showed that the compression strength of the ceramic block can be reinforced to the level of nature bone by new bone ingrowth into the macroporosi ties in the ceramics; the deposition of new bone occurred and went into active period about 7 and 14 days after implantation into bone separately; for the specimen 4mm in diameter with suitable pore dimension and porosity, the mineralization of new bone in the ceramics was completed in 30 days after implantation approximately; the osteogenesis was observed 2 month after the ceramic blocks were implanted into the muscles of dogs. The porous hydroxyapatite ceramic blocks have been used as the substitutes for bone in clinic.

INTRODUCTION

In recent decade, porous hydroxyapatite ceramics (HAC) is paid close attention due to that new bone can grow into the interconnecting pores in the ceramics if the implant is in contact wi th bone and a bone-bonding can be formed at the interface between new bone and exposed surface of HA grains. As a results,

(llThis project was supported by the National Nature Science Foundation of China and the State Education Commission of P.R. China.

409

artificial bone may be constituted from porous HAC. This porous HAC is reinforced greatly, after pores are filled with new bone /1-7/. This paper gives a evaluation for the effect of reinforcing porous HAC by new bone ingrowth into pores and investigates the progress of osteogenesis in the porous MAC implants in order to seek the possibility of extending applications of the ceramics in clinic.

MATERIALS

PREPARATION OF POROUS HAC BLOCKS. The HAC block were prepared by traditional ceramic sintering technique /8-10/. A slurry of HA powder and an aqueous solution of hydrogenperoxide was made. Drying the slurry at about 100 °C results in formation of pores in the so-called green body. The pore dimensions and porosity in the green body were controlled by the mixed content of hydrogenperoxide and the rate of rising temperature . The green body, then, was sintered at 1250°C for 2 hours. Experimental specimens were obtained by cutting the ceramic block into the required shape.

COMPOSITION AND STRUCTURE. Analysis with x-ray diffraction, showed the phase composition of porous HAC to be essentially pure HA with some B-tricalcium phosphate « 5% ) • The analysis wi th ICP spectrometer showed that the total content of impurities was less than 5 ppm.

Fig. 1 A 8

Scanning electron micrography of the microstructure of a speciment. Original magnifications; a) x 12; b) x 3000

410

The microstructure of porous HAC block is shown by Fig. 1. There are two kinds of pores in the ceramies : macroporosi ties and microporosities.

EXPERIMENTAL METHODS

In the studies of reinforcing porous MAC by new bone ingrowth into pores, three types of speciemens wi th various pore dimensions and porosi ties were prepared. The specimens were shaped in cylinder (cI>4x8). Fi ve adult male dogs underwent the experiments, 6 holes 4mm in diameter were transversely drilled in the both left and right femur separetly. Specimens, 12 pieces (4 pcs x 3 types) were inserted into the holes of each dog 1.2.3.4 and 8 months after implantation, the dogs were sacrified and the implants were collected for the testing of compression strength and histomorphology observation. Same specimens of three above were also used for the implantation in muscles in order to evaluate the effect of soft tissue ingrowth into pores. In this esperiment, 18 pieces specimens, 9 pieces per side (3 pcs x 3 tYPes), were buried into left and right waist-back muscles of each dog separately. Five dogs underwent the experiments.

TABLE 1 The pore dimensions and porosities of three types of porous

HAC.

Sampie Pore dimension Porosity Appearent density (IJ. m) (%) (g/cm3 )

I 50-200, aVe 115 43.60 1.79

II 60-320, aVe 245 47.61 1.66

III 75-550, aVe 380 60.04 1.27

For studying the early stage of host and materials response after implantation, the porous HAC block was sawed in rectangular pieces (1 x 1 x 6 mm). Adult male Sprague-Dawly rats underwent the experiment, a hole cl> 1.5 mm was drilled in both left and right femur midshalfs and the rectangular implants were inserted. The rats were sacrified at 3.7.4.21 and 28 days postoperatively

411

and the implants were collected for examination of scanning electron microscope ( SEM) , transmission electron microscope

(LM) •

RESULTS AND DISCUSSION

REINFORCING EFFECT OF TISSUE INGROWTH INTO THE PORES IN POROUS HAC. The compression strength of implants collected from animal bodies in various periods after implantation was tested in order to evaluate reinforcing effect of tissue ingrowth into the pores in the porous HAC (Table 2). According to Table 2, the effect of reinforcement depends on the pore dimension and porosity in the specimens, Le. on the volume occupied by pores, and is mainly contributed by the ingrowth of new bone. The final compression strength of sample 111 is greater than that of sample 11, and sample 11 is greater than sample 1. The greatest compression strength is reached in sample 111 at 4 months after implantation, and is about 1300 Kg/cm2 which is the level of compression strength of nature bone. By prolonging the time of implantation, the further improvement of compression strength has not been observed obviously. This shows that the mineralization of new bone in porous HAC block in this experimental condi tion is completed 4 months after implantation approximately.

TABLE 2 The compression strengthes of specimens implanted in femur and

waist-bach muscles of dogs in various periods.

In bone In muscles Per iod (Kg. / cm2 ) (Kg. / cm2 )

month Type of sample Type of sample

I II III I II III

0 201-243 193-215 145-159 201-243 193-215 145-159

1 636-662 594-606 316-328

2 av.263 av.277.5 av.278

4 835-938 1003-1150 1170-1300 av.260.3 av.262 av.265

8 1070-1200 1210-1280 av.303 av.331 av.328

412

PROGRESS OF OSTEOGENESIS IN POROUS HAC. The following is abrief description of the results coming from collaborators cited in the aknowledgement. The histomorphology investigation was carried out with Leiz LM after implantation into the femur of rats. Results are given in Table 3.

TABLE 3 Percentage of bone, osteoid, chondroid and soft tissue at the

interface up to 28 days after implantation.

Days Bone Osteoid Chondroid Soft tissue

7 2.14 0 0 97.86

14 27.28 0 0 72.72

21 49.83 0.40 0 49.77

28 68.00 0.41 0 31.59

Fig. 2 ( a ) and Fig. 2 (b ) are TEM and SEM images 7 days after implantation. Fig. 2(a) shows that the osteogenisis has been in progress and many mineralized areas can be observed in both of the interface and non-interface. In Fig. 2(b) the fibrillar network has grown into the macroporosities in the specimen.

A B

Fig. 2 TEM image (a) and SEM image (b) 7 days after implantation. Original magnifications: (a) x 1600, (b) x 2800.

413

14 days after implantation the osteogenesis went into active period (see table 3) and the mineralized areas were increasing rapidly. At 28 days after implantation, most of the interface was covered by new bone.

Fig. 3 Decalfied light micrograph of sampIe 111 30 days after implantation into the femur of a dog.

Fig. 3 is a decalcified light micrographs of sampIe 111 30 days after implantation into the femur af a dog. The new bone fully filled the pores in the implant peripherally as weIl as centrally • Only the macroporosities were filled by new bone, whereas the microporosities with diameter less than 1 Nm were not. It is a interesting finding that the osteogenesis was found in muscles after the porous HAC were implanted in the waist-back muscles of dogs about 2 months (Fig. 4). This shows that the porous HAC has probably the ability to induce osteogenesis, which is contrary to all results published by other investigators. Further studies to confirm this surprising resul t are being carried out.

Fig. 4 Decalfied micrograph implantation into the magnification: x 300.

of implant muscles of

2

a months after

dog. Original

414

CLINICAL APPLICATION

The porous HAC blocks have been used not only in filling bone defect cavities but also in replacing injure bone, for example, mandible and so on Fig. 5 and Fig. 6 are the photographes for a mandible replacement and a spine repairment.

Fig. 5 Photograph for a mandible replacement.

Fig. 6 Photograph for a spine repairment.

A interesting phenomenon is that the spli ts in mandible substitute caused by a injury can be healed by itself one year after the porous HAC substitute was implanted. This shows that the porous HAC may from a reconstructed bone with some vitality after it was implanted in bone for enough time.

415

SUMMARY

1. The porous HAC imp1ants can conduce new bone to grow into the macroporosities in implant, and a bone-bonding can be formed at the interface between new bone and exposed HA grains after implantation. As a resu1 t, the porous HAC wi th enough pore dimension and porosi ty may be reinforced to· the level of strength

of nature bone by bone ingrowth about 4 months after implantation.

2. The active period of osteogenesis is reached about 14 days after implantation and the mineralization is approximetely

completed about 4 months after implantation in the experiment.

3. It is possible for the porous HAC implant to have the function

to induce osteogenesis after implantation into body. 4. The porous HAC can be extended to the application as the substitutes for some bone bearing load in clinic.

ACKNOWKEDGEMENT

The authors thank Miss Ou Ya for carrying out the experiment of

implantation in dogs; furthermore they thank U.M. Gross and Ch Muller for their collaboration.

REFERENCE

/1/ M. Jarcho, Calcium phosphate ceramics as hard tissue prosthetics, Clinical Orthopaedics and Related Research, 1981, 157, 259-278.

/2/ Heimke G. and Griss P., Tissue interaction to bone replacement materials, Ed. Oe Groot K., Bioceramics of Calcium Phosphate, Boca Raton, Florida, CRC Press Inc., 1983, 79-97.

/3/ Schoenaers J.H. , Holmes R.E. , Finn R.A. , Healing in interconnected porous hydroxyapatite blocks, Long-term histology and histomorphometry, Trans. Soc. Biomater., 1985, 110.

/4/ Pieeuch J.F., Coldberh A.J., Shastry C.V., Compressive

Strength of implanted porous replamineform hydroxyapatite. J. Biomed. Mater. Res., 1984, 18, 39-45.

/5/ Holmes, R. E. substi tute in Surgery, 1986,

416

Paraus hydroxyapatite metaphy seal defects, 68A, 904-911.

as a bane graft J. Bane and Joint

/6/ Tracy B.M. And Doremus R.H., Direct electron microscopy studies of the bone-hydroxyapatite interface, J. Biomed. Mater. Bes. 1984, 18, 719.

/7/ Rout P.G.J., Tarrant S.F., Frane J.W. and Davies J.E., Interaction between primary bane cell cu1tures and biomaterials, 111. A comparison of dehse and macroporous hydroxyapatite in Biomaterials and Clinical Application, Pizzoferrat, A., Eds., Elsevier, Amsterdam, 1987, 591.

/8/ Klawitter J.J., A basic investigation of bone growth into a porous material. Ph. D. thesis, Clemson University, South Carolina, 1970.

/9/ De Groot K., Ceramics of calcium phosphates: preparation and properties, Ed. De Groot K., Bioceramics of Calcium Phosphate, Boca Raton, Florida, CRC Press Inc., 1983, 99-113.

/10/ Posner A.S., Betls F. And BlumenthaI N.C., Formation and structure of synthetic and bone hydroxyapati te, Prog. Crystal Growth Charact., 1980, 3, 49-64.

417

BIOIDGICAL APATITE AB A MATERIAL FOR ARl'IFICIAL 10m A PRELIMINARY ~GATI{JII {JII I'l'S RmIBILITlC

Kazushi lUKJI'A National Institute for Research in Inorganic Materials,

1-1 Namiki, Tsukuba,Ibaraki, 305 JAPAN

ABSTRACT Stoichianetric hydroxyapatite, irnplanted in soft tissue of rats were exanined, and it was found that the hydroxyapatite has changed to non-stoichanetric one. The results suggest that the induction period necessaIY for bonding fonnation beb.1een sintered hydroxyap9tite and hard tissue can be shortened by using non-stoichanetric hydroxycpatite as artifical bane materials.

IN'l'RXlX:TI{JII

It has been generally considered that stoichianetric hydroxyapatite is a thennodynamically stable phase at neutral pli region of the system CaO­P20S-H20. This means that the solubility of the stoichianetric hydroxy­apatite is lower than the other calcium phosphate phases in neutral aqeous solution. Due to this character, sintered stoichianetric hydroxyapatite can keep its shape for long time in living bodies, being an excellent material for implantation.

On the other hand , apatite phase found in living bodies contains minor inorganic canponents. It is also known that, within living bodies, the activities of those, for exernple, H,Na,K, Mg, Ca are kept constant, playing an essential role. Since the apatite phase in the living body is assumed to have sane kind of interaction through metabolism, i t is considered that there exist a certain relation beb.1een the activities. and the canposition of the apatite phase.

As is well known, the apatite which is used for irnplant material is stoichianetric hydroxyapatite, of which the canposition is different fran that of the apatite in living bodies. IXles this difference give any effect on its biological affinity? Or, is it possible to design and prepare the apatite for irnplant material with higher biological affinity with long life ?

As a preliminaIY approach fran this point of view, the possibility of the change of stoichianetric hydroxyapatite within living bodies was investigated.

418

MMERIAI.S AN[) HE'I'R:U)

Sintered dense hydroxyapatite grains of c.a. 100 .,m in diameter were bonded with each other, by means of sintering, to fo:rm parous pellets of 3 nm in diarreter and 2 nm in thiclmess. These pellets were implanted in the muscle of rats for 4 or 12 weeks. Then, these pellets were analysed by means of EIMA. The results are shown in Table 1.

Fig. 1. After 4 weeks, the pellets with muscle was mounted in plastics, cutted and palished for the analysis by means of EIMA.

Fig. 2. For canparison, a pellet made of apatite cement was implanted. After 4 weeks, the pellet dissolved partly into the tissue.

Fig. 3. nor size.

419

4 weeks after, the hydroxyapatite grains do not change in shape

Fig. 4. 12 weeks after, no change in the shape nor size of the grains can 'oe found. Tissue did not enter into the pellet.

420

TABLE 1 Chernical canposition of hydraxyapatite grains detennined by means of EPMA

Ca P

wt% atans wt% atans before use 55.8 (±0.2) 10.00 42.4 (±0.2) 6.00

4 weeks 55.0 9.85 42.4 6.00 12 weeks 54.8 9.82 42.5 6.00

CalQ( P04)6(OH)2 55.82 10.00 42.39 6.00

RESJLTS ( 1) Expressing the canposi tion of hydroxyapati te by the fonnula,

Ca10-x(M,H)x(P04)6(OH)2-x nH20 , where M is metal cations, the values of x deduced fran Ca and P concentration for the grains of sintered dense hydroxyapatite shown in Table 1 are as follows ;

x = 0.00 ± 0.03 (as sintered, before use) x = 0.15 + 0.03 (4 weeks) x = 0.18 + 0.03 (12 weeks)

( 2) Minor metallic element detectable by means of EIMA was only Na. Est.imated concentration was 0.01 wt%Na or less for both 4 weeks and 12 weeks spec:imen. By means of SIMS, using CAMECA 4F, Na was detected but the quantity was unable to detennine.

DISCUSSI(lIl AN[) CXH:LUS1(lIl It has been generally considered that, in the system CaO-P~5-H20, the non-stoichametric hydroxyapatites are not stable.

For canparison, stoichianetric hyroxyapatite powder was mixed with aqueous solution of 0.1 F NaH2P04 and kept at 1400c and 3.6 atme for 600 hrs. the hydroxyapatite powder thus obtained was examined by means of gravimetrie and volunetric analysis. The results are shown in Table 2. SEM.image of starting and product powder is shown in Fig. 5.

TABLE 2 Chemical canposition of hydroxyapatite powder treated in 0.1 F NaH2P04

aqueous solution at 140<{:: and 3.6 atme

Ca P Na Ca/P wt% atans wt% atans wt% atans

before use 39.71(±0.03)9.954 18.42(±0.02)5.975 0.000 0.000 1.666 300 hrs. 39.61 9.929 18.42 5.975 0.005 0.002 1.662 600 hrs. 39.60 9.926 18.43 5.978 0.004 0.002 1.660

In the starting powder, inpurity of max. concentration was carbonate ion (20 ppm as 002). Adsorption of moisture is not negligible.

421

Fig. 5. SEM image of the hydroxyapatite powder at starting A, and the product B.

In Table 2, a little but detectable decrease in Ca/P ratio was fo~n~ Thus, it is possible to consider that the non-stoichiametric hydroxyapatite is thermodynamically stable phase, at least in living bokies. Ion exchange equilibria may takes place between living body and hydroxyapatite. A precise catching in canposition of hydroxyapatite is desirable for ion activity matching as an implant material.

422

THE DESIGN AND MANUFACTURE OF JOINT PROSTHESES AND STRESS DISTRIBUTION

Paolo Dalla Pria Lima-Lto, Flagogna (UD) - Italy

ABSTRACT

Modern Computer Aided Design (CAD) and Manufacturing (CAM) technologies are used in the design and production of prosthetic implants. These allow for a greater study of the optimum shapes as weil as the simulation of the mechanical coupling of components and their interaction with the surrounding bone tissue. For this last point, the use of the Finite Elements Method (FEM) allow for the deter­mination of the stress distribution fjeld on both the implant and the biological host structure, with good accuracy. Moreover, it becomes possible to visualize the effects induced on the bone cells activity, as weil as the deformation of the implant.

INTRODUCTION

Joint prostheses are still the source of complex problems of a technical, biological,

social and ethical nature. They are, in fact, called upon to restore joint function

almost completely with no major restriction of the subject's day to day activities, and

for increasingly longer periods, thanks to the progress of medicine and improved

structures and social conditions. They are also being employed on an ever wider

scale in revision surgery. The wealth of clinical experience that has now been ac­

quired, together with numerous studies of bone biology and mechanics, has led to the

identification of many biological and mechanical parameters and hence the creation

and continuous development of a large number of models. Definitive solutions, of

course, have not been attained. This, indeed, is a goal that may never be reached.

Great importance is naturally attached to the methods used in the design and manu­

facture of prostheses with a view to improving their reliability, service life and toler­

ability. Abrief description will now be given of the main steps in the design and

creation of a prosthesis. A product development flow-chart is shown in Fig.l.

423

PROSTHESES DEVELOPMENT FLOW-CHART

CLINICAL REQUIREMENTS TYPE OF PROSTHESIS-SURO. TECHNIQUE

! PROSTHESIS PROJECT I I MATERIAI.S ·1

! A

I r~T'=lJr's f - - - - ~ ]MANUFACTURINOI 20 CAO TECHNIQUES

~---- TOOLS

9 9 ff==~ SHAPES .~ DIMENSIONS f--o COUPLINOS

I : f~ I I I I I I 3D CAO I I ...,.

I COUPLINQS I I I FEM SURFACES I I ~ I I

I

-------01

I CAM I b..- 1 I ! 9 I I

I PROTOTYPES I lcJ- - MECHANICAL TESTS I PROVISIONAL I

IcJ-- T -1--- MACHININO CHECKS ____ ~ ILUEPRINTS FINISHINO ANALISYS

I I WORKSHOPS

I I

I : PRE-SERIES I I I I I I I lcJ- --I ..... ---"1----

CLINICAL TESTS

I I I PRODUCTION ..... --- ENGINEERING - -01

FINAL CHC PROORAHS FINAL BLUEPRIHTS OC PROCEDURES

PROOUCTION

Fig. 1. Prosthesis development f1ow-chart. Solid Iines show the design paths and dashed Iines the revision loops.

424

BASIC DESIGN FACfORS

A prosthesis must clearly be designed to meet certain clinical requirements in func­

tion of its use and the type of condition of derangement involved. At this stage,

therefore, attention must essentially be directed to its biomechanical principles and

the surgical technique needed for its implantation. An assessment of its industrial

feasibility is of equal importance. Indeed, it may often be necessary to determine

which materials will be used and make an approximate estimate of the production

costs.

1WO-DIMENSIONAL CAD

The next step is to determine the general measurements of the prosthesis and the

number of sizes required. CAD technologies are used to provide a graphic illustration

of the solutions adopted. The measurements are then parametrized to ensure opti­

mum dimensioning of the set of sizes and any couplings between components.

Two-dimensional representation of a plurality of views of the object is usually suffi­

cient at this stage. In addition to dimensioning and shape analysis, attention must be

directed to definition of both the technology required to produce the prosthesis, and

the main surgical instruments that will be used for its implantation.

THREE-DIMENSIONAL CAD

Computerised simulation of the components is now unfertaken. The 2D design of the

prosthesis is displayed in space by means of a set of boundary surfaces (Fig.2).

If required, these can be coloured, iIIuminated and shaded. Graphic models of this

kind, therefore, show what the finished product will look like (Fig.3).

Shading can also be employed for 3D simulation of the couplings between the com­

ponents of a prosthesis and its insertion into abone.

Computerised surface modelling - a CAM technique - is then used for extremely

precise determination of the machine tool paths in function of the workpiece material

and the known technological parameters (machining stage, tool types, cutting speeds,

etc.) (Fig.4). The computer is also used to write the machining program that will be

read by a Computer Numerical Control (CNC) machine (Fig.5).

425

Fig. 2. Hip prosthesis modular Neck designed using 3D CAD surface development.

Fig. 3. Computer assembling trial between the component parts of a modular hip prosthesis. Perspective and surfaces shading are used in order to obtain an image

close to the reality.

426

Fig. 4. CAM Technique . Tools paths are traced in the computer before machining. Tool paths and machine parameters (type of tools, speed, et cetera) are recorded in

a file wh ich is loaded by the computerized tooling machine.

Fig. 5. CAM Technique. The Computer Numerical ControUed machine (CNC) reads the previously recorded data file and works the piece to the desired shape.

427

STRESS ANALYSIS

Simulation of the stresses to which a prosthesis will be subjected when actually in use is of extreme importance in preventing the possibility of its yielding due to fatigue. Structural analysis can be performed numerically or experimentally. In the past, it was necessary to carry out many tests to see how components behaved under loads

and then correct any defects that were found. This was naturally expensive in terms of models, tooling and time. Today, computerised simulation is very widely employed. The method most com­monly used is called the Finite Elements Method (FEM). (Fig.6) In FEM simulation, the computer generates a structure called mesh composed of a very large number of basic components or elements. Each element or family of ele­ments is assigned the mechanical and physical characteristics of the prosthesis mate­rial needed for calculation purposes (e.g. its Young's modulus and Poisson's ratio). Constraint and loading conditions as close as possible to those that will be met in

practice are then assigned. (Fig.7)

A specific software, known as the calculation code, now mathematically simulates the behaviour of the structure under the physical and geometrical conditions imposed. This gives an 'a priori' indication of the stress~s and strains a prosthesis will be subjected to during use (Fig.8). The same method can also be employed to assesss the behaviour of the surrounding bone structures, though here the extreme variability of the geometry and physical parameter values of human bones means that no more than a qualitative evaluation can be expected.

FEM greatly reduces the need for experimental tests. It cannot entirely replace them, however, since it does not permit prediction of the effect of the implant environment and cyclic stresses on the fatigue strength of the components of a prosthesis, nor their tribological behaviour.

428

Fig. 6. Finite Element Model (shaded) of a threaded hip prosthesis in a femur. The sectioned femur allows for the visualization of the contact between prosthesis

and bone.

429

Fig. 7. Finite Element Model of the same threaded hip prosthesis (non shaded) . The picture shows the element meshes (metallic prosthesis, ceramic head, cortex

bone and spongeous bone) and the applied loads.

430

Fig. 8. Stress map of the prosthesis. This is a very useful way of representing stress distribution values. The stresses are shown as grey islands, and to each grey value corresponds a stress value, from white (Iowest stress) to black (highest stress) .

Stress-color maps are preferred to B&W for better interpretation.

431

PROTOTYPES

FEM stress distribution analysis and optimisation is followed by the preparation of

prototypes for use in checking the work of the machine tools, and the precision of the

machining and surface finishing of critical parts. The finish of prosthesis compo­

nents, in fact, can have a great influence on their fatigue strength. This parameter

cannot be simulated on a computer.

A certain number of prototypes are thus used for simulated dynamic testing in the

laboratory under the extremely severe loading and constraint conditions for the pa­

tients established in the international standards (e.g. ISO 7206/4). This is the most

critical design stage in terms of strength, since a prosthesis will only be regarded as

reliable if all the specimens pass the cyclic loading endurance tests.

Prototypes are also used to simulate surgery and hence assess the efficacy of the

instruments designed to insert and possibly remove the prosthesis

CLINICAL EXPERIMENTATION

Following the introduction of any modifications suggested by the outcome of the

previous stage, a pre-series of prostheses is made for clinical experimentation, whose

results will be obviously be of fundamental importance in deciding whether or not the

product will be placed on the market.

PRODUCT ENGINEERING AND VOLUME PRODUCTION

At the end of the period needed for adequate clinical experimentation, the whole

project is reviewed and such modifications as may be necessary are introduced.

These may involve several stages of the design process. Lastly, before volume pro­

duction is commenced, all the product quality control checking stages needed to meet

GMP (Good Manufacturing Procedures) standards, are established.

432

BIOMECHANICAL PRINCIPLES OF THE SURGICAL TREATMENT OF THE LONG BONES NON UNIONS

G.F. Zinghi, L. Specchia, A. Moroni, G. Galli, G. Rollo, C. Sabato 3fd Department of Orthopaedie Surgery, Rizzoli Institute, Bologna,Italy

ABSTRACT

Basieally the biomeehanical principles 0/ the surgieal treatment 0/ the long hones non unions are thejixation o/the non union site and the enhancement o/the osteogenetie process. In our non unions series hoth rigidjixation and intramedullary nailing were used. In this paper the results 0/ the surgical treatment perjormed in aseries 0/548 long bone non unions are reported. We classify the long bone non union as primary and secondary. Aecording to our classijication the primary /orms are the non unions which have been previously treated conservatively. The seeondary /orms are the non unions which have been previously operated on. From the radiographie point 0/ view the primary non unions are usually closed, without bone necrosis. The seeondary non unions are the/orms due to an inadequate previous surgical treatment. These /orms are usually lapses with a large amount 0/ hone necrosis.

HUMERUS NON UNION

From 1968 to 1988, 147 humerus non unions have been operated on. Thirty-seven (25.1 %) non unions were primary, 110 (74.9%) secondary. The primary fOnDS have been operated on by internal rigid fIxation by plating. The secondary forms have been operated on by resection of the non union site and fIxation by plating and cortical bone grafting. In these fOnDS it is very important that the resection plane is perpendicular to the longitudinal bone axis in order to load the non union site with compression stress. (1-2-3)

The cortical bone graft is used to avoid rotating deplacements and to obtain a very fInD and stable fIxation that is very difftcult to achieve because the large amount of porotic bone. Homologous bone grafts were used in all the cases. (2-34)

Union was observed in all the primary non unions (100%) and in 106 (96.3%) secondary non unions. The 4 (3.7%) faHures were due in three cases to an insufftcient removal of the necrotic bone and in one case to an inadequate fIxation because a too short plate.

As post-surgical complications six infections were observed. Infection healed in four of these cases after a surgical debridement. In the two other infected cases healing of the infection was obtained after the removal of the plate and mounting of an external ftxator.

433

FOREARM NON UNION

The foreann non unions modify the anatomo-functional units of the foreann that are radius, ulna, radio-ulnar joints, interosseous membrane and ligamentus quadratus. The integrity of these units is very important for the prono-supination range of motion. In these cases the aim of the treatment must be both the union and the salvage of the prono-supination range of motion. 1-2-3-4

From 1968 to 1988, 112 foreann non unions were operated on (fig. 1). Fifty-tree (47.3%) cases followed both radius and ulna fracture, 29 (25.9%) radius fracture, 17 (15.1 %) ulna fracture, 11 (9.9%) Monteggia fracture and 2 (1.8%) Galeazzi fracture.

Internal fixation by plating combined to cortical bone grafting was used. Fixation was performed with the foreann in intermediate prono-supination position. Thin AISI 316 L Sherman design plate was utilized. This plate is easily adaptable to the bone surface incongruencies. 3.5 diameter screws were used. An opposite cortical bone graft was implanted to achieve a very fmnstability.

Union was achieved in all the patients. Time to union ranged from 3 to 5 months. In 2 cases the onset of a post-surgical infection was observed but, despite the infection, union was finally obtained.

Figure 1. a) severe secondary radius non union in a 32 years old male patient operated on by plating and bone grafting; b) radiographie result at a 18 months follow-up.

434

FEMUR NON UNION

From 1968 to 1988, 138 femur non unions have been operated on (fig. 2-3). 62 (45%) were primary forms and 76 (55%) secondary. The primary forms were due to open fraetures 3 (4%) or to an inadequate conservative treatment 59 (96%). The secondary forms were due to an inadequate surgieal treatment 69 (91 %) or to a previous infection 7 (9%).

Twentyseven (20%) non unions were in the proximal third ofthe femoral diaphysis, 98 (71 %) in the medial third and 13 (9%) in the distal third. .

The primary non unions of the proximal third were operated on by plating with eompression of the non union site, the seeondary non unions of the proximal third by resection of all the necrotie bone and Muller plating eombined to opposite eortica1 bone graft The femoral medial third non unions both primary and seeondary were operated on by Kuntseher nailing or Grosse-Kempf interlocking nailing. In one case arevision procedure with apposition of bone grafts was necessary. 1-2-3-4-5-6

The femoral distal third primary non unions were operated on by tour Eiffel internal fixation with 2 crossing Rush nails, the femoral distal third secondary non unions by Muller plating.6-7-8

Union was observed in alt the eases. In 3 patients shonening of the limb more than 2 eentimeter was present

In alt the medial third forms treated by Kuntscher nailing union was achieved. Healing was observed in the very severe forms treated by Grosse-Kempf interlocking nailing also. In these cases the time to union ranged from 4 months to 12 months.

Figure 2. a) primary non union of the femoral distal third in a 44 years old male patient; b) the radiographie result 8 months after Tour Eiffel internal fixation by two crossing Rush nails.

435

Figure 3. a) secondary femur non union in a 27 years old male patient operated on by plating and later by extemal fixator; b) healing of the non union 6 months after the Grosse-Kempf

imerlocking nailing; e) the radiographie result 12 months after the removal of the nail.

436

TIBIA NON UNION

From 1968 to 1988, 151 tibia non unions were operated on (fig. 3). Primary forms were 59 (39%), secondary 92 (61 %). The surgieal treatment was the same as in the femur non unions aeeording to the anatomicallevel of the non union. In the non unions of the tibial proximal third tour Eiffel internal fixation by two erossing Rush nails was performed, in the non unions of the tibial medial third Kuntscher or Grosse-Kempf locking nailing and in the forms of the tibial distal third Grosse-Kempf locking nailing or tour Eiffel fixation were used. 1·2-3-4-5-6

The primary forms with eorregible malalignement of the tibial mechanical axis have a very positive prognosis. In these forms, in ease of associated varus procurvation deformity, a fibula osteotomy has to be performed.

In the primary forms with no eorregible malalignement of the tibial meehanical axis a metaphysea1 tibial osteotomy must be associated. 4-5-6-7-8

The more frequent deformity is the varus procurvation malalignement. In ease a varus deformity is present a fibula osteotomy must be eombined.

In the infeeted forms reseetion of all the osteomyelitie bone, mounting of the Ilizarov fixator, metaphyseal eortieotomy and bone transportation at the rate of 1 mm per day was performed.

Results of all these techniques were very positive: in all the eases union was finally aehieved.

As post-surgieal eomplieations 1 infeetion in one ease operated on by nailing, 3 shortening more than 1 em of the limb and 8 malalignement were observed.

Figure 4. a) seeondary tibia non union in a 26 years old male patient; b) radiographie result 5 months after Kunthseher nailing and fibular osteotomy.

437

REFERENCES

1. Ruggieri F., G.F. Zinghi, Lanfranchi R.: Le pseudoartrosi diafisarie. Aulo Gaggi Editore, Bologna 1977.

2. Ruggieri F.: 11 trattamento chirurgico focale delle pseudoartrosi delle ossa lunghe. Chir. arg. Mov., 58, 358, 1969.

3. Judet 1., Judet R.: L'osteogenese et les retards de consolidasion et les pseudoarthroses des os longs. Atti SICOT, New Yotk 1960.

4. Pawels F.: Biomecanique de la greffe osseuse. Acta Orthop. Belg. 37,701, 1971.

5. Ruggieri F., Sgarbi G.: 11 trattamento delle pseudoartrosi e dei ritardi di consolidazione della diafisi femorale con chiodo endomidollare di Kuntscher. Chir. arg. Mov., 57, 342, 1968.

6. Ruggieri F., Zinghi G.F., Soncini G.: Sull'uso differito deI chiodo di Kuntscher in a1cune fratture deI femore edella tibia. Atti SERTOT, 16, I, 1974.

7. Muller M.E., Allgower M., Willenegger H.: Technigue of internal fixation of fractures. Springer, Berlin 1963.

8. Kuntscher G.: Practice of intramedullaty nailing. Charles C. Thomas, Springfield 1967.

438

QUALITY CON'l'roL AND SOVIET STANDARDS FDR BIOCERAMICS

VITALI A. DUOOK, LEONID L. SUHIH Institute for Problems of Materials Science

Ukranian Acaderny of Sciences 252180, Kiev-180, UkrSSR

ABSTRACI'

Standardized alumina bioceramics properties are reviewed as well as succession of developIeI1t of such standards in the USSR. Fbr rational choice of bioceramics properties to standardize it is necessary to consider a standard as SUffi total of our knowledge on bioceramics functioning in human body. Such approach enables to pick up for each bioceramics kind the IOOSt essential properties that must be standardized and sare other properties of minor iroportance. The above is illustrated by correlation of strength lass nechanisrns and characteristics of several bioceramics. Of all possible quality contral nethods should be preferred those which to the greatest degree would reproduce main processes to affect bioceramics in human body.

INTRODUCl'ION

Ceramic :iIrq?lants have been used in the USSR for roore than 30 years. Last

decade this activity has greatly intensified and since 1989 special

project und.er the title "Bioceramics" is included into gove:rnmental

program "Advanced materials".

Soviet surgeons have at their disposal roore than 80 types of dental

and orthopedic :iIrq?lants though its output is far frau sufficient. M:>st of

these :iIrq?lants are made of alumina biocerarnics and single crystals which

are standardized by Health Ministry. Sarwa standardized properties of

soviet and sare other alumina based bioceramics are listed in Table 1.

Though standardized properties of soviet bioceramics are lCMer than

in international standards laboratory data attained in the USSR are on

the sarre level and it would not take a long tine to reproduce than on a

439

greater scale. Rapid progress in alumina par.ders and sintering technology

gives good outlook as to considerable :irlproving these prqlerties

thmughout the world in the next ferN years.

TABLE 1 Standardized properties of alumina based biooeramics.

Properties Units of Standards CXltlook measure-ments ISO DIN USSR

Ceramics Single crystal

Density g/an3 3.90 3.94 - 3.97 3.96 Bending strength MPa 400 450 300 500 600 Co'Ipressive strength MPa 4000 4500 2500 3000 4500 Grain size ...m 7 - - 2 IIlpact

Jjm2 strength 4000 - - - -Fracture

MPa*ln1/ 2 touglmess - - - - 5 Total :ilIplrities % mass 0.5 0.3 0.5 0.01 0.1 Solubility

gjm2*day in bicm:rlia 0.1 0.1 - - 0.05

In the USSR besides state standards there are also sane other kinds

of standards which give the same rights to produce and to put materials

in evel:}'da.y practice. State standards establish prqlerties of materials

and nethods of measurenents for a lot of materials which are produced 00

a !arge scale and for a loog t:ime, e. g. steels, Sem;! alloys, netal

par.ders etc.For the recently developed materials as a rule are used so

called teclmical cooditioos and for m:rlical materials and devices -

medico-teclmical requirEmm.ts. There is also special standard to

establish oIder of elaboration and to obtain pennissioo for producing of

medical wares (N15. 013 - 86).

In CCIIpliance with the standard all m:rlical developnents and testing

are carried out by special Health Ministry organization - Institute of

Research and Testing of Medical Appliance or under its direct guidance

and IlDJSt be awroved by Health Ministry. SUch m::ncp:>ly results in

440

insufficient c::cupetence in special devices or materials.

SCIENTIFIC FOUNIlM'Irn OF STANDAlIDIZATIrn

Today's order draw no attentiCXl to elaboration of scientific

foundations of standat:ds i. e. of ratiooal choice of paraIteters to standardize and nethods of their measurements CXl the base of up-to-date

knowledge of physical and other processes which deteDlline bioceramics

functiooing in human body. It is of major :inportance for bioceramics as

it becanes inseparable part of human body and interacts with it very

int.imately. There are two main aspects in this interactiCXl:

1) ceramics influence CXl human o:rganism and

2) human o:rganism influence CXl bioceramics.

The first of these problens is investigated by biological methods

and depends mainly 00 ceramics che.mical CClIp)Sition. The ooe problen is

studied by material science methods and includes large scope of tasks in

physical chanistry, physics and chemistry of strength etc.

ekle of the main points of this problen is bioceramics strength loss

as a result of bianedi.a influence.

There are different processes to lead to the strength loss rot we

can pick out several causes to launch the processes:

1) thenoodynamical nonstability of bioceramics or one of its ~ts

in bianedi.a;

2) thenoodynamical nonstability of scma bioceramics local parts

stipulated by mechanical stress, chemical or phase inhaoogeneity,

i.nhaoogeneous distrib.ltioo of illperfections etc.;

3) bioresorption of ceramics in human body;

4) ceramics intercalation in bianedi.a.

Let us examine in short these processes in scma ceramics.

1. <:ne of the IOOSt frequent causes of strength loss process is

hydroxide fonnatioo of bioceramics main carp:>nents and :inplrities because

of their thenoodynamical nonstability. For instance in alumina based

ceramics magnesium hydroxide fonnation usually takes place as magnesia is

added to restrict grain growth and to increase ceramics strength. In

zirccnia based yttria stabilized ceramics yttrium hydroxide fonnation was

obsel:ved [1]. Investigation of hydratation process in different liquids

reveals that the IOOSt active liquid medium m:>lecules are those with lone-

441

pair electran across a protcn donor site. Besides water hydroxide

fonnatioo. is pnm:>ted by annatia and hydrazine blt it is prohibited by

hydroca.rbal roolecules.

2. In the zircalia based ceramics strength lass is stipulated by

tetragonal - m:noclinic phase transfonnation. This transfonnation

usually takes place at high teuperatures or at considerable tensile

strains blt the coo.ditions cf the transfonnatioo. are greatly affected by

chem.ically active media [2]. For exanple when zirconia powCer is placed

into fluorhydric acid tetragonal phase transfonns into IOOnOClinic during

few minutes [3].

3. All ceramic materials never have chem.ical or energetic haoogeneity.

This is cormected both with crystal and chanical inhaoogeneities at grain

boundaries as well as IOOChanical strains coo.centrated at grain boundaries

and at the cracs tips. These inhaoogeneities greatly enhance phase

tranfonnations, hydratation and other physico-chem.ical processes which

lead to strength lass [4].

4. Bioactive calcium plx>sphate based ceramics undergoes bioresoIptioo.

processes in human body at a rate which within several omers of

magnitude depends on ceramics activity that is crystal perfection and

energetic haoogeneity. Insufficient contral of the latter parameters

results in extraooly poor reproducing of phosphate based ceramics

resoIption experiments [5, 6]. Of all phosphate based ceramics

hydroxyapatite reproduces the bone mineral ~ition to a greater

degree and this would be a reasoo. of its minimum resoIptioo. rate. FUrther

developnents in this field should be directed to zoore precise reproducing

of boo.e phase and chanical ~ition daNn to the smallest ilrg;mrity

coo.tent.

5. carlx:n based bioceramics strength can be decreased by intercalatioo.

i. e. by penetratioo. of atans or even large roolecules between carbon

layers in graphite crystal structure. For carbon ilIplants such process

has not been investigated.

BIClCERAMICS CRUCIAL PIa'ERI'IES

This short review shows that bioceramics properties which deteIInine

strength degradation are different for different bioceramics. The sarre

character has intemependence between all bioceramics functional

442

prcperties such as 1Near resistance, solubility, fracture t.cughness and

usually measured structural, chemical and other characteristics ought to

be divided into two parts:

1) main characteristics which are crucial for bioceramics functialing in

human body and

2) less inp>rtant characteristics depending cn the above or not so

significant for bioceramics functicning.

Sane properties such as hardness, microhardness, CXlIpressive

strength grea.tly exceed those which are necessary for bioceramics

functicning .

Stardan:lized characteristics ImlSt embrace tull set of crucial

parameters which is variable for different bioceramics kinds. Other

characteristics can be cited in the standani as auxiliary or for

infonnaticn .

AB an exanple, for alumina based bioceramics can be used a set cf

crucial pa.raneters as follows:

1. Chemical CCIIpOSiticn

2. Total inplrity ccntents

3. Toxic inplrity ccntents

4. Bending strength

5. Fracture tougtmess

6. Mass loss :in bianedia

7 :' strength loss in bianedia

8. Wear resistance in bicmedia

AuXiliary alumina pa.raneters may include such characteristics as

density, grain size, phase CCIIpOSiticn, CXlIpressive strength, pore

characteristics - total and open porosity, pore norplx>logy and pore size

distrihlticn .

For another kind of bioceramics set of crucial pa.raneters ought to

be altered.

StandaJ:dizaticn of bioceramics pa.raneters are no obstac1es to design

and use variety of :i.nplants even if sane parts cf which have elevated

porosity and ttlerefore lower strength or special chemical CCIIpOSiticn

provided that main part of the :i.nplant is made of standardized bioceramic

material.

443

1. sate, T. and Shimada, M., Transfonnatien of yttria-doped tetragooal Z202 polycrystals by annealing in water. J. Am. Cer. SOc., 1985, 68, 356-59.

2. Dnmlald, I.L., In vitro aging of yttria - stabilized zirccnia. ~ Am. Cer. SOc., 1989, 72, 675-76.

3. Ra6aHOBa. M. H •• ,lly60K. B. A •• OKOPOXO.D;, B. B •• C»aaoBhle HSMeHemm B nopollIKax Ha OCHOBe ,ItKOKCH.D;a :mqlKOmm IIpH paCTBopemm: B IJJIammOBO:l lCHCJIOTe. .uOKJIa,IlbI .AB roop, cepHß B, 1989, N6, 38-41.

4. SclJnauder, S., Schubert, H., Significance cf intemal stresses for the martensitic transfonnatien in yttria-stabilized tetragooal zirccnia polycrystals during degradaticn. J. Am. Cer. SOc., 1986, 69, 534-40.

5. Klein, C.P.A.T., van der ü.1bbe, H.B.H., Driessen, A.A., de Groot, K., van der Roof, A., Biodegradatien behaviour of various calcium-phosphate

materials in subcutaneous tissue. In Ceram i es in SUrgery, ed. P. Vincenzini, Elservier Scientific Publishing catpany, Amsterdam, 1983, W. 105-114.

6. waisbrod, H., Gerbershagen, H.U., A pilot study of the value cf ceramies for 1:x:ne replacanent. Arch. Orthop. Trauma Surg., 1986, 105, 298-301.

444

PHYSICO-CHEMICAL TECHNIQUES TO CHARACTERIZE STRUCTURE ARD COMPOSITION OF INTERFACES INVOLVING BIOCERAMICS:

POTENTIALITY ARD LIMITS

FABIO GARBASSI and ERNESTO OCCHIELLO Istituto Guido Donegani S.p.A.

Via Fauser, 4 - 28100 Novara (Italy)

ABSTRACT

The potentiality of several spectroscopic techniques to give information on ceramicsjbiosystem interfaces is reviewed. Emphasis is placed on the necessity of crossing different techniques and of investigating (or simulating) "in situ" conditions. In this frame, a special mention will be done microscopies, a family of microscopies which allow to study real interactions between substrates and biological materials.

INTRODUCTION

on tunnelling are likely to

non-biological

Bioceramic materials categories considering time interval in which be observed ( 1) .

can be classified in different their reactivity and consequently the modifications due to interactions can

Materials like alumina, isotropic carbon and amorphous carbon are NEARLY INERT: their reactivity is very low and the consequences of interactions are observed only after years. Hydroxylapatite and some special glasses are SURFACE ACTlVE: modifications are limited to a surface layer (a thin layer) and occur in aperiod of months. Finally, salts like calcium sulphate and trisodium phosphate, having high reactivity, are fully RESORBABLE and the effect of such a reactivity are evident after a few weeks or days. A special category are COMPOSITES, where reactivity is mainly determined by moieties present at the surface. Two main families can be considered, ceramic coatings on metals or polymers and reinforcing materials in a polymer matrix, like carbon fibers. Bioceramic materials are usually polycrystalline, sometimes single crystals or amorphous. Chemical bonds are ionic or covalent, that is strong and directional: consequently the most of them are good thermal and electrical insulators.

Their chemical inertness, if so, determines a good corrosion resistance, while wear resistance 1S connected to hardness. Negative properties are brittleness and difficult processability. Surface properties can dramatically differ, in terms of structure and/or chemical composition, from bulk properties. Furthermore, the peculiarity of surfaces can be enhanced by the presence of heterogeneities, contaminants and adsorbed species. The characteristics of the interface zone between the implant and the host biosystem is strongly dependent on the nature of both. In fact, considering the reactivity of the implant, only a thin layer of it can be involved, as weIl as the whole object, when it is completely resorbed. On the other side, the biological moiety can react to the presence of the implant weIl far from it. As a consequence, the interface zone can be easily observed using conventional techniques like optical microscopy. However, in this way many details, occurring in the microscopic scale, are lost and only their macroscopic consequences are revealed. Examples are adsorption/desorption and catalytic phenomena, dissolution of the material and protein denaturation. In order to gain more details on the interface zone, many techniques have been developed in the last 20 years. Most of them are, strictly speaking, surface techniques. Their are used to characterize the non-biological moiety be fore and after the contact with the biosystem, indirectly obtaining information on phenomena occurring at the interface. only a few techniques are available to study directly the latter and, as specified in the following, only in special situations.

SURFACE SPECTROSCOPIC TECHNIQUES

spectroscopic techniques able to characterize solid surfaces have been developed mostly around a basic idea: to irradiate the specimen with a primary entity, then analysing the information carried by the emitted secondary entities. 80th entities can be photons or particles (electrons, ions, neutrals). The surface specificity is connected to the mean free path of the secondary entity in the solid, the shorter is the latter, the higher is the probability to come from a very superficial layer. In this sense, and considering a suitable energy range (10-3000 eV), surface specificity is high for techniques using ions as secondary particles, medium in the case of electrons and low for photons. In Table 1, relevant characteristics of techniques taken into consideration are collected. Background theory, characteristics, highlights and drawbacks of most of them are described in more detail in Ref. (2).

Ion spectroscopies The most popular techniques based on ion beams are Ion Scattering Spectroscopy (ISS) and Secondary Ion Mass Spectroscopy (SIMS), both of them need to operate in

TABL! 1 Surface spectroscopic techniques.

Technique Information ~p~ Observed Sensitivity SampIe (*) Qual Quant Profile Layer (#) Damaqe

-------------------------------------------------------------ISS E s~i No 0.2 nm -4/-6 low SSIMS G No No < 1 nm -7 vI XPS E+G s~i Yes 3-5 nm -2/-3 vI US E s~i Yes 3-6 nm -2/-3 low I~

ir V.S. Semi Yes 1-10 ~m -2 no Raman V.S. Semi Yes 10-500 nm -1/-3 no

D~ ir V.S. Semi No ....... -3/-4 no

PM ir V.S. No Yes 1-50 ~m -1/-2 no

-------------------------------------------------------------(*) Meaninq of acronyms qiven in the text (#) Power of ten E: surface elements; G: surface chemical qroups; V.S.: vibrational spectrum; Semi: semiquantitative; vI: very low.

ultra-hiqh vacuum (UHV), in order to limit ~e interference between ions and atmosphere • ISS is based on ~e measurement of ion enerqies after collision of the primary ion beam (usually noble qas ions) with ~e specimen surface. The enerqy after cOllision, that with qood approximation can be considered elastic, depends on the incidence angle, primary enerqy and ~e ratio of the atomic masses present at ~e surface and of ~e primary beam. Thus, measurinq ~e scattered beam enerqy, atoms present at the surface can be recoqnized, ~rouqh ~eir mass value. since ~e scatterinq events are localized at the first atomic layer of the solid specimen, ISS is very surface-sensitive. On the o~er hand, its diffusion is limited wi~ respect to o~er techniques because qa~ered information is not so rich considerinq its complexity and cost. SIMS consists in ~e mass-analysinq of particles emitted from ~e specimen upon ion or atom bombardment. Two types of experiments are possible, in static (in this ca se we speak about SSIMS) and dynamic mode, dependinq on primary ion current, ~at is much lower in the first case. SSIMS is used for surface characterization, due to its surface sensitivity, while dynamic SIMS qives a profile of ions concentration qoinq to ~e bulk, since ~e irradiated area is rapidly eroded. SSIMS is a complex technique both for experimental probl~s, like sampIe charqinq, and interpretation of results, which is not easy and stronqly sensitive to surface contamination, however it is becominq a very popular technique.

Electron Spectroscopie. X-ray Photoelectron Spectroscopy (XPS) and Auqer Electron

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Spectroscopy (AES) belong to the family spectroscopies. Again UHV is necessary experiments, in order to minimize the absorption electrons.

of electron during the of secondary

XPS is now the most popular spectroscopi~ technique for studying solid surfaces. The spectrum 1S obtained by irradiating the sample with a nearly monochromatic X-ray beam, usually emitted from a Mg or Al source. Irradiated atoms emit electrons ("photoelectric effect") having an energy given by the difference between that of the incident photon and its original energy level. If the emitting atom is not far from the surface, electrons can escape from the solid, being collected by an energy analyser. Energies are typical of every atoms, easily recognized in the spectrum. Peak intensities are related to the atomic concentration, allowing a semiquantitative analysis. Energy level perturbations caused by the presence of chemical bonds are reflected in "chemical shifts" in the spectrum, so giving additional information on the oxidation number and chemical state of detected atoms. The richness of data, together with the possibility of analysing samples of whatever nature and shape (powders, fibers, sheets, pellets etc.) and detecting all atoms but Hand He contributed to the enormous diffusion of XPS. On the other hand, it shows a limited lateral resolution (1 mm in conventional instruments, now improved to the 50 ~m scale) and is not so specific towards surfaces like ion spectroscopies, since the emitted electrons are collected from a layer of 20-40 nm depth. AES is based on the effect . discovered by Pierre Auger in 1926. When an electron is emitted by photoelectric effect, a hole is created in an electronic level, which can be filled by another electron belonging to a more external level of the same atom. The transition energy must be dissipated, following one of two preferred and alternative paths, that is the emission of a photon (X-ray fluorescence) or the emission of a third electron from the same or a more external level with respect to the second one: this is the Auger electron. Even if the Auger transitions are visible also in a photoemission spectrum, together with XPS peaks, independent or combined spectrometers have been developed, using an electron gun as primary source. Due to the huge background of secondary electrons, Auger electrons are better evidenced by signal derivatization. UHV system, analysis chamber and electron energy analyser are the same of XPS, one of the main differences is that in AES the position of peaks is independent from primary beam energy, while in XPS it is dependent. Using an electron beam to irradiate the sample can cause serious problems of sample charging, as a consequence insulators are difficult to analyse, problems connected to sample reduction or decomposition have been occasionally observed. Qualitative and quantitative analysis and surface specificity are again analogous to those of XPS, while information connected to chemical bonds is more scarce and difficult to interpret. Since electrons are easily focused,the lateral resolution can be pushed down in the submicronic scale. Instruments have been developed able to scan the electron beam trough the sample, giving Auger maps.

Such technique is known as Scanninq Auqer Microscopy (SAM). A review on AES can be found in Ref. (3).

Vibrationa! Spectro8copie8 A number of techniques have been developed based on the analysis of electromaqnetic radiation (from radiofrequency to near-ultraviolet) which are able to observe a layer of limited thickness with sufficient sensitivity. Generally speakinq, these techniques do not require UHV facilities, however their specificity for surfaces is rather limited. Internal Reflection spectroscopy (IRS), called also Attenuated Total Reflectance (ATR), is based on the total reflection of an incident electromaqnetic wave occurrinq at the interface of two different media, havinq the appropriate refractive indexes. The experiment needs the use of a suitably cut sinqle crystal (for instance, sapphire), which acts as interna 1 reflection element (IRE), in infrared, Raman or fluorescence modes. Consequently, the vibrational spectrum is collected, allowinq the identification of functional qroups in the examined layer or, in the latter case, the spectrum and lifetimes of excited states, that is the electronic spectrum. Normally, a layer of material of several micrometers is analysed, however the penetration of the infrared radiation can be decreased below the limits allowed by variations of anqle of incidence and refractive index of IRE, by depositinq on the examined surface a non-absorbinq film of controlled and uniform thickness. Obviously, in this way sensitivity is also stronqly affected. Diffuse Reflectance Spectroscopy (DRS) has lonq been used, mainly in the UV/visible reqion, but recently became popular also in the IR mode, followinq the introduction of Fourier-transform instruments. In the latter ca se it is often called DRIFT (Diffuse Reflectance Infrared Fourier-transform spectroscopy). DRIFT is made by rather complex instruments and onerous data handlinq, but does not requires particular sampIe treatments and is rather sensitive. Diffuse reflectance occurs when the spectrum is a function of both absorption and scatterinq events, with stronqly scatterinq sampIes like powders. DRS is essentially a bulk characterization technique, becominq more superficial if a phase is deposited on a substrate not qivinq peaks in the same reqion. The use of a shallow layer of a hiqhly reflectinq powder, like KBr, has been also reported to enhance the scatterinq at the KBr/sample interface. DRS has been widely used to study silanes on qlass fibers and silica powders, detectinq coveraqes less than a monolayer. Also orqanic functional qroups on silica qel have been studied, with acceptable semiquantitative results. PhotoAcoustic Spectroscopy (PAS) is based on the detection and evaluation of acoustic siqnals produced by convertinq to thermal enerqy the electromaqnetic enerqy transferred to a solid by irradiation. The technique is suitable for thick and opaque solid sampIes, introduced in a qas-filled cell havinq a wall transparent to the incident radiation; the detector is essentially a microphone. The thickness of the sampIe layer contributinq to the photoacoustic siqnal depends on

M9

the thermal diffusion length, therefore on thermal conductivity. Since in ceramics thermal diffusion lengths are in the ~ ranqe, surface selectivity is equal or less of that of infrared IRS. The photoacoustic cell is the heart of the instrument, its construction is a very important task; the purity of the fillinq qas is also important, since traces of water and carbon dioxide, havinq stronq i.r. absorption, influence the spectrum. An advantaqe is that samples can be analysed in any form, like powders and rouqh surfaces. Aqain silica samples have been heavily studied.

Transmission spectroacopiea Transmission spectroscopies observe the whole material, however, in the presence of a diffuse interface (or hiqh area surface) they can be usefully applied. An example is Raman spectroscopy of powdered bioactive qlass after immersion for a variable time in a physioloqical medium. The formation of a surface species identified as Ca carbonate was observed (4). Also NMR can be used, when a clear-cut distinction exists between interface and bulk siqnals. Several NMR studies have been devoted to the bondinq of silanes on silica qel, usinq these system as a model for qlass fiber composites.

Suitability of Spectroacopic Techniquea From the above survey, it appears that two qroups of techniques are qrossly available. The first one, formed by ion and electron spectroscopies is very surface sensitive but cannot be used in realistic conditions, since UHV is needed; the second one, formed by remaininq techniques, does not respond to required surface selectivity. Speakinq about surface characterization, two other techniques must at least be cited, Scanninq Electron Microscopy (SEM), to observe surface morpholoqy, and contact anqle measurements, to measure surface enerqetics. Previous experience on complex systems suqqests that the combined use of different techniques on the same specimen can qive a better picture. For instance, combininq SSIMS and XPS renders easier the interpretation of SSIMS data and qives a chance to qain quantitative information. On the other hand, very superficial phenomena can be missed in XPS; in this case, contact anqle measurements, which are sensitive to the top layer, allow to check if somethinq occurred. Since contact anqle measurements are sensitive to morpholoqy and chemistry, its combination with XPS and SEM is very useful. Photon spectroscopies can be used in this frame to make on the same system deeper measurements (so performinq a deep profilinq) and to exploit their ability to work in a more realistic environment, i.e. air atmosphere, or better, liquids.

SCANNING TUNNELLING MICROSCOPY AND RELATED TECHNIQUES

In order to find a technique able of a of the microscopic interface between

direct visualization a bioimplant and a

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biological medium, recent progress in scanninq Tunnellinq Microscopy (STM) and related techniques seems very promisinq. STM was developed in the early 80's by Binniq and Rohrer at IBM Labs in Zurich (5), who later won the Nobel prize for such a discovery. The basis of STH lies in the tunnellinq phenomenon, i.e. the current observed between two closed electrodes when a voltaqe is applied. One of the electrodes is constituted by the sampIe, the second one by a metallic tip, in principle as thin that a sinqle atom lies on its top. The experimental arranqement is sketched in Fiq. 1.

DRIVE

z~ X

CONTROL UNIT

IMAGE

Figure 1. Sketch of the experimental STM apparatus.

Movinq in the xy plane constituted by the specimen, the tip measures the tunnellinq current variations due to the distance between the two electrodes. Since the current is extremely dependent on such a distance, a maximum value is measured when the tip is in front of an atom, a minimum when it is in front of interstices between atoms. The horizontal and vertical resolutions are 0.1 nm and 0.01 nm, respectively, thus the topoqraphy of tunnellinq current represent the situation in atomic scale. Two scanninq modes are used, constant heiqht for very flat surfaces (a 10x10 nm to 100x100 nm area is scanned), constant current for rouqh surfaces (in this case the scanninq dimensions are in the ~m ranqe). Since solvent or air molecules are not detected, presumably for their fast movements, STH experiments can be carried out in air and liquids, and this is very appealinq in the field of interactions between solids and bioloqical fluids. A limit is represented by the necessity that the electrodes can be conductive, thus only metal or semiconductor surfaces can be examined. Insulators can be visualized only as thin (less than 2 nm) layers on conductinq substrates. Insulatinq substrates must be previously metallized. A sister technique has been developed, Atomic Force

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Microscopy (AFM), which overcomes the above limitations (6). The physical principle, sketched in Fig. 2, is similar to profilometry.

DRIVE

STMnp

AFM T1P

SAMPLE

z y

CANTllEVEA

CONTAOl UNIT

Figure 2. Sketch of the experimental AFM apparatus.

With respect to STH, the tip is he re placed on a canti lever beam, which is sensitive to the force between the tip itself and the substrate, so cOllecting an image of attractive, repulsive and frictional forces, with aresolution analogous to STH. To maX1m1ze the obtainable information, combined STH/AFM instruments are now in the market. The first applications of STH and AFM were on metal and semiconductor surfaces, demonstrating the powerful ability of visualizing single atoms. More complex systems have been then examined, like single chain polymers and monomolecular Lanqmuir-Blodgett films. Two ways are available to study non-conductive substances: native uncoated samples on inert conductive substrates like gold or highly oriented pyrolytic graphite, or metal coated samples, suitably prepared on glass or mi ca substrates. Each method of preparation has its advantages and dis advantages (7). Metal-coated samples are physically stable, tip-sample perturbations, and give raise to interpreted and quantified, also in height. hand, samples must be first air-dried or freeze dried, then they are exposed to vacuum, thermal stress due to hot metal gases.

minimizing the images easily On the other

frozen, and/or contamination,

Uncoated samples do not request delicate sample preparation procedures, produce direct images of the molecules, also in their hydrated state. The presence of liquids and the possibility to carry out electrochemical experiments is allowed with special cells developed for these tasks. Disadvantages consist in the instability of molecules, due

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also to interactions with the tip, and the difficulty to interpret the obtained images, S1nce theory is not well established for non-conductive samples. Measuring sample heights is indeed a problem: va lues measured in STM are generally far less than those obtained by independent means. This circumstance is tentatively attributed to the higher electron work function of the molecule with respect to the substrate. In spite of above difficulties and limitations, the number of biological studies using STM exploded in the last three years, with a lot of published work concerning nucleic acids, polypeptides, proteins, bacterial cell walls, biological membranes, viruses, etc. A list of references is reported in (7) • Molecular dimensions and shapes have been observed, as well as ability to form, in given conditions, preferential adsorbate layers on substrates or molecular associations between proteins. Soft biological samples have been observed too, an example is a closed stoma with guard cells of a leaf of philodendron cordatum (8). The AFM image well overlaps that obtained by optical microscopy. Special instruments have been designed for biological studies, including a light microscope and the possibility to vary the scanned area fron nanometres to micrometers (9). Besides STM and AFM, several other scanning probe microscopes are in development, all able to detect phenomena at atomic resolution and based on the same control of the probe movement. They detect light, magnetic field, ion conductance, capacitance, thermal variations and so on and so forth. In conclusion, tunnelling spectroscopies seem good candidates for the observation of microscopic interactions between ceramic materials and biosystems in a rea1istic situation. A previous characterization of the surface is necessary, in order to allow an interpretation of the observed phenomena.

REFERENCES

1. Hulbert, S.F., Bokros, J.C., Hench, L.L., Wilson, J. and Heimke, G., Ceramics in clinical applications, past, present and future. In High Tech Ceramics, ed. P. Vincenzini, Elsevier Science Publishers B.V., Amsterdam, 1987, pp. 3-27.

2. Garbassi, F. and Occhiello, E., Spectroscopic techniques for the analysis of polymer surfaces and interfaces. Analytica Chimica Acta, 1987, 197, 1-42.

3. MCGuire, G.E. and Holloway, P.H., Applications of Auger spectroscopy in materials analysis. In Electron Spectroscopy, eds. C.R. Brundle and A.D. Baker, Academic Press, Lendon, 1981, vol. 4, pp. 1-84.

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4. Ravaqlioli, A., Krajewski, A., Ponti, P., Valmori, R., Messori, M. and Moroni, A., Physico-chemical bShaviour of a bioactive qlass coatinq for metal prostheses. In High Tech Ceramics, ed. P. Vincenzini, Elsevier Science Publishers B.V., Amsterdam, 1987, pp. 91-8.

5. Binniq, G., Rohrer, H., Gerber, C. and weibel, E., 7x7 reconstruction on Si(111) resolved in real space. Physical Review Letters, 1983, 50, 120-3.

6. Binniq, G., Quate, F.C. and Gerber, C., Atomic Force Microscope. Physical Review Letters, 1986, 56, 930-3.

7. Fisher, K.A., Yanaqimoto, K.C., Whitfield, S.L., Thomson, R. E., Gustafsson, M.G.L. and Clarke, J., scanninq Tunnellinq Microscopy of planar biomembranes. Ultramicroscopy, 1990, 33, 117-26.

8. Gould, S.A.C., Drake, B., Prater, C.B., Weisenhorn, A.L., Manne, S., Kelderman, G.L., Butt, H.-J., Hansma, H., Hansma, P.K., Maqonov, S. and Cantow, H.J., The atomic force microscope: a toolo for science and industry. Ultramicroscopy, 1990, 33, 93-8.

9. Emch, R., Descouts, P. and Niedermann, Ph., A small scanninq tunnellinq microscope with larqe scan ranqe for bioloqical studies. Journal of Microscopy, 1988, 152, 85-92.

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PROßLEMS CONCERNING THE INDUSTRIAI, PROnUCTION OF ALUMINA CERAMIC

COMPONENTS FOR HIP JOINT PROSTHESES

ERHARn DöRRE Feldmühle AG. Werk Plochingen

7310 Plochingen / Germany

ABSTRACT

High - purity alumina is the only ceramic material which has successfully proved for major components of hip joint pro­theses.

Since 1974 a total number of more than 400.000 joints with ceramic components have been implanted. At present an annual number of about 100.000 ceramic prostheses have to be added. Considering these quantities, every effort has to be made to ensure an extremely reliable, safe and reproducible produc­tion.

Production and inspection processes of alumina ceramic compo­nents are described in detail, emphasizing quality control, avoidance of defects, identification and documentation as weIl as aspects of specification and standardiation.

INDTRODUCTION

The term "High-Tech-Ceramics" comprises the oxide-ceramic materials aluminum oxide (AlzO)) and zirconium oxide (ZrOZ) as weIl as the non-oxide-ceramic materials silicon carbide (SiC) and silicon nitride (Si3N4). They all meet the mecha­nical requirements of high-load bearing joints. However, SiC und SisN4 show a poor biocompatibiltiy, whereas the biocompa­tibility of alumina and zirconia is beyond dispute. Zirconia, however, undergoes a strength reduction under the influence of body fluids as it could be confirmed again [1].

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In addition to this, zirconia is always accompanied by dangerous radioactive elements with a very long half-life period, such as uranium, thorium and actinium. These elements belong to the same column of the periodic system as zirconium and are, therefore, almost not to be seperated fram zirconia. For this reason, the most of the zirconia raw materials show a radioactivity which is considerably higher than the tolora­ted levels. In contrary to former results of radioactivity measurements which were restricted to the less dangerous gamma radiation, other papers reported obout high amounts of alpha radiation, when measuring zirconia animal test samples and zirconia femoral heads which were free available on the market [2]. It is well known that even low-energy alpha radi­ation inside the body will cause cancer. One single alpha­particle decay can already release adegeneration of soft and hard tissue cells.

Therefore, among the High-Tech-Ceramic-materials, high-purity alumina is the only one which was successfully proved for major components of high-load bearing hip joint prostheses due to its biological, chemical and mechanical properties. Since 1974 a total number of more than 400.000 hip joints with components of the high-purity alumina grade BIOLOX have been implanted worldwide. This success is based on a 40-years experience with wear and corrossion resistant components of this material in the field of mechanical engineering [3].

Ceramic-Metal Hip prostheses

Although ceramic materials generally and alumina ceramies particularly have been proposed as biomaterials already in 1934, the realization of this idea took almost 40 years. After extensive simulator and animal tests, the first clini­cal application of high-purity alumina ceramies took place in the form of ceramic-metal hip prostheses in the early seven­tieth. The stern which has to endure considerable dynamic fatigue loads remained metal. All attempts of using ceramic sterns failed completely. The articulation was either ceramics on ceramics or ceramics on polyethylene (PE). This concept did not change basically up to now.

The mounting problem of the metal stern and the ceramic femo­ral head was solved by means of a taper fixation. It turned out to be absolutely stable against rotation. This has to be assured at any rate, considering the metal wear debris which would be developed when articulating with a material 15 times higher in hardness. As it was demonstrated by extensive tests, the frictional torque between the metal taper and the ceramic head is 40 times the frictional torque between head and acetabulum component [4]. And, indeed, a relative motion between ceramies and meta 1 did never happen among the 400.000 clinical cases mentioned before.

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Due to the high meehanieal strength of the eeramie material used here and due to the extremely elose toleranees of this taper fit, the eeramie heads withstand statie loads of up to 100 kN and dynamie loads of up to 60 kN [5]. This material does not undergo any statie fatigue, as it eould be demonstrated with aging tests in Ringer's solution and in animal tissue with preloaded eeramie heads. After an aging period of one year the heads preloaded with 30 kN did not show any load-depending strength reduetion due to this proeedure, as it turned out at subsequent load-to-fraeture tests [6].

This was also a preeondition of the FDA for the approval whieh was given to this material in 1982. The approval of the Japanese Health Ministry followed 1985.

One of the most important properties of the articulating eomponents is the tribology during a long period of time, i.e., the wear and friction behaviour. By means of compara­tive tests in hip joint simulators the tribologie superiority of high-purity alumina ceramic materials could be evidenced [7,8]. These results were confirmed by 16 years of clinical experience, whieh showed the following annual average wear rates [9].

Metal/PE: approx. 200 microns Alumina/PE: " 20 to 130 microns Alumina/Alumina: " 2 microns.

In the case of Alumina/PE the figure of 20 microns refers to an evaluation of explanted components. It includes only wear. The figure of 130 microns refers to penetration measurements between head and aeetabulum component by means of evaluating radiographs. It includes, therefore, wear and plastic flow.

The reason for the superior tribologic behaviour of alumina ceramics are adsorption processes on the eeramic surface resulting in a better adhesion of the synovia lubrication film [10]. This can also be confirmed by measurements of the wetting angle whieh demonstrate a considerably smaller angle for alumina as for metals and polymers.

The superior wetting behaviour of the alumina surface can be reeognized already without any scientific measurement, just by comparing the behaviour of water drops on metal and ceram­ic surfaces. The drops form little spheres on a metal sur­face, wheras they spread out distinctly on an alumina sur­face.

An important prerequisite for the superior tribologic results of alumina, particularly in case of an alumina/alumina artic­ulation, is an extremely smooth polished surface with an average roughness of 0.01 micron and an extreme congruence of the sliding faces with a roundness deviation between 0.1 and 1 micron.

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STANDARDISATION

An international standard for "Implants for Surgery - Ceram­ic Materials based on Alumina" has been established in 1981 (ISO 6474). National standards exist in USA, UK, France and Germany. The national standards are all in full conformity with the ISO standard and still valid. They specify the major properties, such as purity, density, grain size and strength as weIl as wear and corrosion resistance.

Meanwhile the demands on the properties of ceramic components for high load bearing joints increased. Therfore, it was necessary to adapt the standard to the new requirements. A new standard is now in preparation, prescribing the density to be increased from 3.90 to 3.94 gJcm 3 and the strength from 400 to 450 MPa. The average grain size will be reduced from 7 to 4.5 microns.

Regarding the purity, the present standard restricts the content of Si02 and alcaline oxides lower than 0.1 %. This is necessary because of the detrimental influence of these con­taminations on the sintering process. They impede the densi­fication and promote the grain growth. The new standard will also include CaO into the group of detrimental impurities, because it prornotes the static fatigue properties of the sintered product by deteriorating the corrosion resistance [11].

INDUSTRIAL PRODUCTION

The production process of high-purity alumina ceramic compo­nents for high-load bearing hip prostheses is carried out according to powder metallurgical procedures. The starting material is a very pure and fine-grained powder of alpha­alumina. High-Tech applications such as orthopaedic implants require a powder with a purity of at least 99.9 % aluminum oxide, a weIl defined grain size distribution and a high surface activity.

The chemical composition and the mineral phases are analysed by means of X-ray fluorescence and X-ray diffraction methods. Various forming processes require variuos grades of compres­sibility, and accordingly, various kinds of powder preparation.

During all steps of production prior to sintering, clean environmental conditions are required to avoid detrimental contaminations which could react with the alumina during the sintering process. The prepared powder is precompressed with apressure in the range of 2000 bar. A homogeneous desifica­tion is required to assure a symmetrical shrinkage during sintering. Shrinkage differences in the same body can result in interna I stresses. In the unfired stage the pressed bodies are in a condition similar to chalk and can, therfore, be machined easily by turning, milling or grinding. However,

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this requires a sufficient strength and a good machinability of the pressed bodies to avoid defects.

The rules of GMP require a full identification for all major components of orthopaedic devices, such as hip joints. To meet this requierement, it is necessary to mark the parts individually. This is done by means of a computer-controlled engraving machine.

All steps of production and all results of the intermediate and final inspections are documentet. All documents are kept in duplicate on microfilm for 20 years. Thereby, every par­ticular component can be identified and traced back to the raw material. This was also a requirement of the FDA to ob­tain the approval.

The most important step of the production is the sintering process which is carried out in air atmosphere. The tempera­ture ranges, from 1600 and 1800 o C, depending on the proper­ties of the raw material.

Astriet control of the sintering process and small additions of MgO as a grain growth inhibitor are essential in order to achieve a fully dense sintered body with a fine grained microstructure [12]. Besides the computerized temperature control by means of thermal couples and pyrometers, 100 heads which are symmetrically distributed all over the interior of the furnace are used for adetermination of density and grain size, providing another control of the sintering process.

Due to the thermal densification during the sintering pro­gress, the body to be sintered undergoes a linear shrinkage which is of the order of 20 %. This has to be taken into consideration during the green maching process. The course of shrinkage takes place under maintenance of the geometrical shape. The shrinkage is always directed towards the so-called "neutral fiber".

After sintering, the first crack inspection is carried out by means of a penetration method with a fluorescent additive. This way microcracks with a width of about 1 micron can be still detected. Three more creek inspections follow.

A so-called "hard machining" after sintering is necessary if the required final dimensions cannot be attained by the form­ing and sintering process alone or if there are special sur­face requirements, and this is always the case for ceramic hip joint components. Due to their extreme hardness in the sintered condition, high alumina eeramic products can be machined only with tools harder than aluminat such as diamond tools or diamond grit.

Grinding the bore hole of the femoral heads is very critical and has to be done very earefully and precisely because of the extremely elose toleranees which are necessary for a proper function of the taper fixation in order to obtain a

459

high load bearing capacity. The ta per angle tolerance is in the order of 2 minutes and the tolerance of the straightness is 2 microns along an average length of 20 mm.

Whereas the grinding operation gives the heads the exact geo­metry, the subsequent polishing procedure provides the neces­sary smoothness of the sliding faces which is about 0.01 mi­cron, measured as the average roughness.

INSPECTION

Greatest emphasis is put on the intermediate and final in­spection operations. 20 various steps of quality and dimen­sional control are carried out. All parts ar inspected at 100 %. 5 % of the whole production are tested by means of destructive methods. This way also statistical data of the fracture strength can be obtained.

Usually, a surface roughness is measured with a surfindicator by using an inductive-controlled diamond stylus. But in case of such an extreme smoothness, the diamond-stylus would leave a recognizable trace which could be rejected as a surface de­feet. Therefore the inspection of the sliding faces is per­formed by means of microscopes. Ceramic heads with typical surface defects, such as holes, inclusions or colored impuri­ties of microscopical size have to be rejected, although these are only cosmetic defects which do not influcence the strength nor the tribological properties.

The dimensions of the bore hole in the head control the func­tion of the taper fixation and have to be, therefore, inspec­ted with particular care and accuracy. The quality assurance repuires - among others - an audit of random controls and a periodical inspection of all measuring devices.

CONCLUSION

Considering the importance of ceramic orthopaedic implants and the possible consequences of areoperation due to a failure, every precaution according to the present state of the art has to be carried out to assure the quality and to avoid any risk of a malfunction, particularly in view of the large production quanities which will be in the order of 100.000 components per year at the time being.

Coming back to the term "problems" in the title of this paper it can be stated now: The industrial production of alumina ceramic components for hip joint prostheses will be running without problems, provided all measures mentioned before have been taken.

RBFBRBNCBS

1. Thomson, I. and Rawlings, R.D., Mechanical behaviour of zirconia and z~rconia-toughened alumina in a simulated body environment. Biomaterials. 1990, 11, 505-508.

2. Cieur, S., Heindl, R. and Robert, A., Radioactivity of Zirconia Ceramics. Presented at the Third International Symposium on Ceramics in Medicine. Nov. 1990 Rose-Hulman Inst. of Technologiy, Terre-Haute, Ind. USA.

3. Dörre, E., Processing, Properties and Application of Alumina Implants. MRS Int'l. on Adv.Mats., 1989, 357-367.

4. Dawihl, W., Dörre, E. and Altmeyer, G., Lebensdauer­bestimmungen einer kraftschlüssigen Reibverbindung an keramischen HÜftgelenkendoprothesen. Biomed. Techn .• 1980, 25, 311-315.

5. Dawihl, W., Altmeyer, G. and Dörre, E., Statische und dynamische Dauerfestigkeit von Aluminiumoxid-Sinter­körpern. Z. Werkstoff technik. 1977, 8, 328-330.

6. Dörre, E., Dawihl, W., Krohn, U., Altmeyer,G. and Semlitisch, M., Do Ceramic components of hip joints maintain their strength in human bodies? In Cermics in Surgery ed. P. Vincenzini, Elsevier Scientific Publ. Co., Amsterdam, 1983, pp. 61-72.

7. Dörre, E., Beutler, H. and Geduldig, D., Anforderungen an oxidkeramische Werkstoffe as Biomaterial für künstliche Gelenke. Arch. orthop. Unfall-Chrir., 1978, 83, 269-278.

8. Dawihl, W., Mittelmeier, H., Dörre, E., Altmeyer, G. and Hanser, U., Zur Triboligie von HÜftgelenk-Endoprothesen aus Aluminiumoxidkeramik. Med.-Orthop. Techn., 1979, 99, 114-118.

9. Dörre, E., Retrieval and Analysis of Ceramic Hip Joint Components. Presented at the Symposium on Retrieval and Analysis of Surgical Implants and Biomaterials.

10. Dawihl, W. and Dörre, E., Adsoprtion Behaviour of High­Density Alumina Ceramics Exposed to Fluids. In Evaluation of Biomaterials, ed. G.D. Winter, J.L. Leray, K.de Groot, John Wiley & Sons Ltd., 1980, 239-245.

11. Willmann, G., Die Bedeutung der ISO-Norm 6474 für Implantate aus Aluminiumoxid. Zahnärztliche Praxis, 1990, 41, 286-290.

12. DÖrre, E. and Hübner, H., Alumina, Springer-Verlag, Berlin, Heidelberg, New York, Tokyo, 1984.

461

BULK ANI). TBIN FILM CARBON MATERIALS FOK BIOMEDICAL APPLICATIONS:

QUALITY CONTROL CRITERIA AND PROCEDURES

F. VALLANA, P. ARRU, M. SANTI Sorin Biomedica S.p.A.

Cardiovascular Prostheses Division Saluggia, Vercelli, Italy

ABSTRACT

The deposition of biocompatible carbons in the form of bulk coatings and thin films is widely used in the manufacture of biomedical devices and involves very complex high technology processes. The development of suitable Quality Control procedures for these materials is of fundamental importance to achieve the best performance of the end products. This paper summarizes the main criteria and experimental techniques adopted by Sorin Biomedica for the Quality Control of biocompatible turbostratic carbons in the manufacture of implantable prostheses.

INTRODUCTION

The selection of materials suitable for the replacement of parts of the living human body is a critical step in the development of any implantable biomedical device. In fact these materials must comply with a great number of strict requirements, in terms of physical, chemical and biological properties. In particular when the device to be manufactured is a cardiovascular prosthesis, the materials to come in contact with blood should have the highest degree of atoxicity and haemocompatibility, as weIl as resistance to corrosion, mechanical stresses and fatigue. The interest in carbons as biomaterials dates back to the late 1960s, when it was discovered that turbostratic carbons developed for use in nuclear reactors possessed all the above mentioned properties (1, 2, 3). This finding led at first to the development of bulk materials such as

462

isotropic pyrolytic carbon and more recently of thin turbostratic carbon films manufactured specifically for biomedical applications. At present these biocompatible carbons are the materials most widely used in cardiovascular prostheses, particularly artificial heart valves. In recent years many applications in other fields of surgery were studied; thin films of turbostratic carbon have been deposited for instance on orthopaedic prostheses and devices for reconstructive surgery (4, 5). The synthesis of biocompatible turbostratic carbons involves two different techniques depending on the applications. A high temperature chemical vapour deposition process (6) is used to obtain bulk sampIes or thick coatings of isotropic pyrolytic carbon when the material has to per form a mechanical function, as in the case of heart valve occluders. On the other hand, a room temperature physical vapour deposition process is used when the carbon material is to be deposited as thin film in order to give a substrate the optimal haemocompatibility without affecting its mechanical properties (7). Both these processes are very complex, due to the great number of parameters involved which can affect the structure and the properties of the resulting materials; therefore particular care must be devoted to the development of appropriate quality control criteria and procedures in order to guarantee the best material performance. This paper is aimed at giving a description of the main concepts and experimental methods adopted in Sorin Biomedica Laboratories for the Quality Control of turbostratic carbons used in the manufacture of heart valve prostheses and other biomedical devices.

QUALITY CORTROL CRITBRIA Alm PROCBDUUS

The main purpose of the Quality Control of critical cardiovascular devices is to guarantee the compliance of the end products with two fundamental requirements: reliability and haemocompatibility. In order to achieve this result all the materials and components involved in the manufacturing processes must be inspected following suitable Quality Control procedures. As far as the carbon biomaterials are concerned, a careful examination of mechanical, dimensional and physical properties is required to ensure the reliability of the prostheses for extremely long periods of time. Moreover the haemocompatibility is strictly related to the crystalline structure, the chemical purity and the surface morphological properties, that can be assessed by means of structural analysis and morphology testing. A general list of the main test methods used in Sorin Biomedica Laboratories is reported in table 1. It must be pointed out that the same procedures are often not applicable to both bulk and thin film coatings, depending on the great difference in thickness and in the amount of material available for the tests.

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464

QUALITY CORTROL PROCEDURBS FOR BULK PYROLlTIC CARBON

The deposition of bulk pyrolytic ca rb on coatings has found its main application in the manufacture of artificial heart valve occluders, where high mechanical resistance to wear, stresses and fatigue are required in addition to the haemocompatibility. Since the CVD process involves the pyrolysis of hydrocarbons at very high temperature, the material can be deposited only onto low porosity graphite substrates, that have very poor mechanical behaviour. Therefore the durability and reliability of the coated component are entirely committed to the pyrolytic carbon layer, whose mechanical properties must be carefully checked. To this purpose the evaluation of the ultimate flexural strength and elasticity modulus is directly performed through the following bending test procedure. Pure pyrolytic carbon slices, obtained from a flat coated disc by means of mechanical machining to remove the graphite substrate, are cut in rectangular shapes. The test samples are then bent with a three-point testing device fitted on a material testing equipment until breakage occurs (Fig. 1). The deformation of the beam and the corresponding load are measured, recorded and converted into stress-strain data, the cross-sectional area of the beams being known from previous dimensional measurements.

Figure 1. Mechanical testing apparatus (left) and detail (right) of the device used for the ultimate flexural strenght measurement according to

the bending test procedure.

~5

Sinee the meehanieal resistanee of the heart valve oeeluders depends on both the ultimate flexural strength and the thiekness of the pyrolytie earbon eoating, thiekness measurements are performed on all the eoated eomponents after the final polishing step by using high magnifieation X-ray inspeetion (Fig. 2).

Figure 2. High magnifieation radiographie inspeetion of abileaflet valve oeeluder.

Furthermore, the meehanieal resistanee of the oeeluder and the reliability of the prostheses eould deerease in time, depending on the reduction of the reslstant cross seetlon of pyrolytic earbon coatlng due to wear phenomena. Therefore abrasion tests are performed by means of a modlfied pin-on-dise equipment to get a direet evaluation of the wear behaviour of the material; nevertheless, this proeedure is quite time-eonsuming, and eannot be used for every eoating lot approval. In order to overeome this limitation, Viekers mierohardness (DPH) measurements are routinely performed on samples drawn from every coating, sinee a good eorrelation exists between the hardness and the wear resistanee of pyrolytie earbon (8). The Quality Control proeedure involves the penetration in the material of a little d1amond pyramid tip loaded with a 50 g weight; the test sample eonsists of a polished metallographie mount embedding a seetioned oeeluder. Sinee pyrolytie earbon is so resilient and elastie to return to its original state after the indentation (9), a thin polymerie film 1s deposited onto the surfaee of the test sample; when the diamond indenter penetrates in the pyrolytie earbon, it leaves on the polymerie film a square perforation,

466

whose dimensions (about 15 ~m) are measured by means of a microscopic grid and then converted into Vickers hardness units. Several indentations evenly distributed along the whole thickness of the pyrolytic carbon layer are required to guarantee the uniform wear resistance of the material. It is weIl known that the optimal behaviour of pyrolytic carbon in terms of haemocompatibility and mechanical reliability is achieved when its crystalline structure is isotropie. The degree of preferred orientation has often been measured by means of X-ray diffractometry techniques(10); nevertheless this method leads to an evaluation of the average characteristics of the coating. Due to the optical biaxial structure of the crystallites, the anisotropy of the material becomes apparent under microscopic polarized light observation (11); therefore metallographie inspection has been found to be an extremely useful tool for the structural evaluation of pyrolytic carbon with high local resolution (12). The Quality Control procedure allows the detection of highly oriented growth features, which could give raise to the fracture of the pyrolytic carbon coating along preferred directions. Other structural defects such as stratifications and amorphous carbon inclusions can be detected on the metallografie sampies. Finally, a very careful evaluation of the surface morphology of the finished occluders is required to guarantee the highest degree of haemocompatibility, the interaction between blood and carbon surfaces being strongly affected by polishing processes (3, 13). Therefore the assessment of the occluder surface morphology is performed on a regular basis through optical and scanning electron microscopy and roughness measurement procedures.

QUALITY COIITROL PROCBDUBBS FOR TBIN CARBON FILMS

The Quality Control procedures used to assess the physical and chemical properties of thin turbostratic carbon films (such as the Sorin Carbofilm™ coating) are significantly different from those previously described for bulk pyrolytic carbon. This is due to the fact that the thickness of these films is very much lower, being in the range 0.3-0.5 microns instead of 0.3-0.5 millimetres; consequently the quantity of carbon available for the Quality Control tests is often less than 0.1 mg/cm2 • Therefore many important features of these materials can be evaluated only by means of sophisticated instrumental techniques. An interesting example is the assessment of the chemical purity of these films, whose haemocompatibility could be affected by the presence of elements other than carbon, coming from accidental contaminations of the deposition environment. The elementary chemical composition can be evaluated through very sensitive methods such as X-ray Photoelectron and Auger Rlectron Spectrometries (XPS, ARS) (14). Nevertheless, these techniques cannot easily give information about the distribution in

467

depth of the impurities, and moreover are very complex and time-consuming. Rutherford Backscattering Spectrometry (RBS) has been found to be the most suitable method for the chemical analysis of CarbofilmTH (5, 7). Basically the technique consists of placing the carbon film as a target in a collimated monoenergetic beam of 2HeV helium ions (Fig. 3, left side) .

ELECTROST A TIC ANALYZER

ION OETECTOR

THIN FILM

SUBSTRATE

::~::::: ::~:::::

::~:::::

AE - E-E <Xdl I 2

g I­

oU

-I: o

Figure 3. Schematic representation of RBS analysis. The left drawing shows the main features of the experimental arrangement. On the right side the basic relationships between the test parameters are pointed out (m = helium ion mass, K - target atom mass, t = film thickness,

d = film density).

Particles scattered from the test sampie as a result of its interaction with the beam are detected and analyzed; their energy distribution depends on the atomic mass of target atoms and their position inside the film (Fig. 3, right side). Therefore RBS is a useful method to identify the atomic masses of elements and their distribution in depth within the first micron under the surface of the sampie (15). A typical RBS spectrum of a high purity carbon film deposited onto a silicon substrate is shown in Fig. 4.

carbon film, this purpose,

and an uncoated interferometric

Furthermore RBS allows to measure the density of the provided the thickness of the coating is known. To measurements of the height of the step between a coated area are performed on special test sampies according to or profilometric techniques.

468

Energy IMeV)

14r _____ ~O,~5------------r-----------~IT5~-------,

E, ESo"-_____ ...

100 Chonnel

Figure 4. RBS spectrum of a thin carbon film on silicon substrate. The signal corresponding to carbon atoms goes from E1 (surface atoms) to E2

(interface atoms). ~E is proportional to film density and thickness. Es~ corresponds to the silicon atoms at the interface. Heavy element contaminations in the film would appear as sharp peaks in the high energy region (> 1 MeV), while light elements would be superimposed to the silicon and carbon signals at lower energy.

A very critical parameter, that must be carefully controlled, is the adhesion strength between the carbon film and the substrate. Many standard procedures are available for its evaluation, but few of them allow to obtain quantitative results. Bond area pull test has been chosen by Sorin Laboratories as Quality Control procedure; this method involves the bonding of nail shaped pul 1 studs to the carbon coating surface by means of a high strength adhesive (16). The test is fundamentally simple, provided that the substrate is rigidly supported and the stud is freely pulled from the test surface in precise perpendicular fashion. The primary information derived is the load to failure. Important secondary information involves the examination of the par ted surfaces as to define the exact plane and the nature of the failure (Fig. 5). The detection of internal stresses in the carbon film is also important, since they can affect the adhesion strenght between the coating and the substrate (17). A simple evaluation of the internal stresses is performed by coating a thin polymerie test sampIe of known thickness and mechanical properties, then measuring its bending radius by means of an optical comparator (18). Finally optical and scanning electron microscopy are currently used to check the evenness of the carbon coating and its surface morphology.

•

trate leier

A

469

B

Figure 5. A: Schematic representation of bond area pull test. B: Test samples and studs after pull test procedure. The parting of carbon film from the substrate (left side) at low loads indicates poor adhesion strength; highly adherent films cause the failure of the glue at high

loads (right side).

REFERENCES

1. Gott, V.L., Viffen, J.D., Dutton, R.C., Koepke, D.E., Daggett, R.L., and Young, V.P., The anticlot properties of graphite coatings on artificial heart valves. Carbon, 1964, 1, 378.

2. Bockros, J.C., Gott, V.L., La Grange, L.D., Fadali, A.M., Ramos, M.D., and Vos, K.D., Heparin sorptivity and blood compatibility of carbon surfaces. J. Biomed. Mater.Res., 1970, 4, 145.

3. Bockros, J.C. , La Grange, L.D., and Schoen, F.J. , Control of structure of carbon for use in bioengineering. In Chemistry and Physic of Carbon, ed. Valker, P.L. Jr, and Thrower, P.A. , Dekker, New York, 1972, 9, 103.

4. Vallana, F., Development of new materials for orthopaedic prostheses. Proceedings of Italy-USA joint meeting on advances in orthopaedic surgery and traumatology, Pavia (Italy), May 26-31, 1986.

5. Paccagnella, A., Majni, G., Ottaviani, G., Arru, P., Santi, M., and Vallana, F., A new pyrolytic carbon film for biomedical applications. Ceramics in clinical applications: Satelite Symposium of the VI Vorld Congress on High Tech Ceramics, Milan (Italy), June 23-27, 1986.

470

6. Akins, R.J., and Bockros, J.C., The deposition of pure and alloyed isotropie earbons in steady-state fluidized beds. Carbon. 1974, 12, 439.

7. Paeeagnella, A., Majni G., Ottaviani, G., Arru, P., Santi, M., and Vallana, F., Properties of a new carbon film for biomedieal applieation. Int. J. Artif. Org., 1986, 9, 115.

8. Shim, B.S., and Sehoen, F.J., The silieon-alloyed isotropie earbon. Organs, 1974, 2, 103.

wear resistance of pure and Biomater. Med. Deviees Artif.

9. Coons, V.C., A rapid method for polishing pyrolytie graphite. Metal Progr •• 1962, 81, 83.

10. Bacon, G.E., A method for determining the degree of orientation of graphite. J. Appl. Chem., 1956, 6, 477.

11. Koizlik, K., and Koss, P., Optieal determination of the anisotropy of pyroearbon eoatings. In 11 th Biennal Conferenee on Carbon, Springfield, Virginia (USA), June 4-8, 1973, Conf. 730601, p. 169.

12. TombreI, F., pyroearbones.

and Rappeneau, J., Preparation at strueture des In Les Carbons, Masson et C.ie ed., Paris, 1965, 25.

13. Feng, L., and Andrade, J.D., Surfaee atomie struetures of biomedieal earbons observed by Seanning Tunneling Mieroseopy (STM). Submitted to J. Biomed. Mater. Res •• 1991.

14. Czanderna, A.V., ed., Methods of surfaee analysis. Elsevier Publ. Comp., Amsterdam, 1975, 4-5.

15. Chu, V.K., Mayer, J.V., and Nieolet, M.A., speetrometry. Aeademy Press, New York, 1978.

Backscattering

16. Riekerby, D.S., A review of the methods for the measurement of eoating-substrate adhesion. Surface and Coatings Technology. 1988, 36, 541.

17. Mattox, D.M., Thin-film adhesion and adhesive failure-A perspective. Adhesion Measurement of thin films, thiek films, and bulk eoatings, AST" STP 640. Mittal, K.L., Ed., American Soeiety for Testing and Materials, 197.8, p. 54.

18. Roark, R.J., and Young, V.C., ed., Formulas for stress and International Student Edition, MeGraw-Bill Kogakusha, LTD, 1965, 6, 73.

strain. Tokyo,

471

DEVF.LOPM~:N1· 0.' AN HUMAN BON.: MARROW CELL CULTIIIU; TO 'fEST TIIE CITOCOMPATIBILITY OF BUIJK HYDROXYAPATITF. MATF.RIAIJS.

WILKE, A.; ORTH, J.; GklSS, P. Orthopaedic Department, Philipps ~niversity, Marburg, Germany

NEHLS, V.: DRFNCKHAHN, D. Institute of Anatom\, Philipps University, Marburq, Germanv

ABSTRACT

SEM. light - and episcopic microseopy was used to Investigate the phenomenoloqieal behaviour of a human bone marrow eells eultured on sintered hydroxyapatite (HA) eeramics. The human bone marrow cells were harvested by a femoral osteotomy at hip artbroplasty operation. The eells were seeded on surfaee treated and untreated porous dense HA eeramics (OSPROVITR,. Tbe cell <'IlHure W,iS stimulated by addition of growth taetors, namely Interleukin J (IL 3) and Granuloeyte - Monocyte - Colonisation - Stimulatlng - Faetor (GMCSF) over two to tour weeks. Thc eells attaehed aod adhered onto the HA surtaees. proliferated and segreqated extraeellular matrix (SEM). No signs of toxieity appeared.

INTRODUC'\'T ON

The bioeompatibility of materials used in the operative medieine is one of the main problems ot sueeessful therapy. Especially the eytotoxieity ot some of the biomaterials redueed their applieation despite of their good physieal and ehemieal properties. The last two deeades of biomaterial deve]opment have seen inereasing attempts to introduee materials, whieh were tissue eompatible (1). One of these special materials is HA. a so ealled surfaee -reaetive material. It induees an aetive interfaeial interaetion between implant and hast tissue. Sinee there are no physieo - ehemieal differenees to the hast tissue (2, 3, 4) sueh surfaee reaetive materials as HA ean form integrative bonds with surrounding tissues if implanted into bone (5, 6, 7).

The formation of such an intimate bond implies eellular. moleeular and ionic aetivity between the hast tissue and the implant surfaees.

Hast tissue eells have mainly two ways to reaet due to a eontaet to biomaterials. On one hand adesion, proliferation and differentiation is induced by an implant and on the other hand dedifferentiation and eellular death is eaused by eytotoxieity (8, 9, 10).

472

Adbesion of cells onto biomaterials is aprerequisite for tbe osseointegration of prostbesis. The osseomorphogenesis seems to be closely associated with the phenomenological behaviour of the anchoraged cells (lI, 12). No cytotoxic materials, which favour the segregation of extracellular matrix are responsible for this phenomenon.

The aim of this study was to develop a cell culture to be used for testing the cytotoxicity and biocompatibility of biomaterials on a phenomenological basis. Since human bone marrow is the most important tissue, which is in contact to biomaterials - especially in prosthesis and osseosynthesis - we established a human bone marrow cell culture, that is supported by Behring Company Marburg. As a testing material two different hydroxyapatite - ceramies (HA), namly OSPROVIT, donated by The Feldmüble Company, were choosen.

MATERIALS !ND NETHODS

Specimen 1 was a sintered, round shaped, porous bulk HA - ceramic with a size of 26 x 2 mm. It had a higb cristallinity and a pureness of more than 98%. The ceramies were sterilized by autoclaving at 1340 C for 15 minutes. Specimen 2 was an analogon to specimen 1 but surface treated.

Scanning electron microscopy pictures of the different ceramies demonstrate tbe treated and untreated surface qualities (Figure 1, 2).

The human bone marrow was harvested at operations tor hip arthroplasties. At the side of the femural osteotomy we removed a marrow containing spongious bone block, which was of a size of app. 0.7 x 1 x 2 cm.

The bone blocks were put into phosphate buffered solution containing Liquemin. After mechanical crushing the material was sieved (porous size: 0.1 mm).

The harvested particles were moved on to a Ficoll Hypaque Density gradient. After eentrifugation at 800 9 for 20 mi nut es the interphase band was removed and five hund red thousand cells were seeded per dish.

Figure 1. SEM picture of a surface of an untreated HA ceramie (magnification X 50).

473

The medium, which we used for culturing, contained the following ingredients: Iscoves medium with 12.5 % fetal calf serum, 12.5 % horse serum, 5 x 10-6 mol hydrocortison, 0.05 mg/mI Streptomycin and 50 IU Penicillin G. The incubation was carried out by 37°C and 12.5 % oxygen and 5% carbon dioxide. The medium was changed each fourth day (13).

Figure 2. SEM picture of a surface of a treated HA ceramic (magnification X 62) •

The following growth factors were added per milliliter culture medium: Between 10 to 100 ng of Interleukin 111 and Granulocyte-Monocyte Stimulating Factor. Particularly the cells of granulopoesis are interesting for our question since they are able to adhere onto surfaces.

Adding of different growth factors to the medium, esp. GMCSF and IL 3, make it possible to stimulate human bone marrow stem cells to differentiate into mature granulocytes within a few days.

The following techniques were used to investigate the adhered cells.

Light .icroscopy The attached cells were removed from the HA - ceramics either mechanically or with 2% Trypsin and then centrifugated at 800 g. A smear of the cell pellets was prepared, air dried and Pappenheim stained.

Episcopic light .icroscopy In order to study the attached cells on the surface of the ceramic they were incubated with Toluidine blue (5 minutes), rinsed with medium and then evaluated in episcopic microscopical technique.

Episcopic fluorescence light .icroscopy In order to study the attached cells on the surface of the ceramic they were incubated with Acridine orange (10 minutes), rinsed with medium and then evaluated in episcopic fluorescence microscopical technique.

474

Scanning electron .icroscopy The attached cells were fixed with 2 , Glutaraldehyd (1 h). After rinsing with Cacodylat buffer solution, incubation with 2' of Osmium - tetraoxide followed (1 h) . After rinsing with Cacodylat buffer solution, tbe contrasted cells were dehydrated in an increasing aceton concentration (10 min. in 50, 75, 90, 96 and 100 , aceton) . Drying in a critical point chamber was annexed. After coating with gold in a sputter coater, microscopical evaluation was done .

RESULTS

The seeded cells were cultured in disbes for one week until a monolayer was built. A smear of these Pappenbeim stained buman bone marrow cells, which were cultured for one week without contact to HA were taken as a control group and demonstrated a physiological population of normoblasts, granulocytes and myelocytes .

Cells, which were cultured for two weeks just beside the HA ceramic were removed and a smear was Pappenheim stained. No cytomorphological differences between the cells removed from tbe surrounding of tbe different HA ceramics and tbe control group could be seen.

SEM demonstrated weIl attached human bone marrow cells cultured two weeks on both HA surfaces. The cells adhered not only on the plane surface but were also seen inside the pores. Tbe cells spreaded out, proliferated and showed morpbological signs of extracellular matrix independently of the applied HA.

After culturing tbe cells for two weeks on tbe different HA ceramics they were removed, smeared and Pappenbeim stained . A physiological population of granulopoetic cells could be demonstrated. No cytomorphological differences compared to the cells, that grew beside the HA and the control group could be seen .

Figure 3. SEM picture of human bone marrow cells cultured on untreated HA ceramic (magnification X 190).

475

After incubation of the attached cells on the ceramic with Toluidine blue, episcopic light microscopy demonstrated vital, attached and spreading cells. The same results were seen with episcopic fluorescence light microscopy after incubation with Acridin orange. As a specific sign tor vitality we can see incorporated Acridine orange particles in lysosomes.

We developed a standardized and reproduceable human bone marrow cell culture. Dedifferentiation of the cells started after 4 weeks. The dominating cell lines were depended on the different applicated growth factors.

An uninfluenced culture of the cells besides and on the different specimens was possible and no cytomorphological difference between the two applied HA's and the control group appeared. No toxicity of the soluble HA substances could be seen. All smears ot Pappenheim stained cells demonstrated physiological populations of granulopoetic cells. No cytomorphological differences compared to the control culture could be seen.

Both HA ceramics demonstrated no signs of soluble toxic substances. The two surface qualities did not show any influence onto the cell differentiation. No material depending dedifferentiation was obvious. The biocompatibility of the ceramics did not differ from each other.

Figure 4. SEM picture of human bone marrow cells cultured on surface treated HA ceramic (magnification X 158)

DISCUSSION

It is possible to cultivate abuman bone marrow cell culture especially of tbe granulopoetic line on tbe examined HA - ceramics. Tbis experimental set up is able to investigate all biomaterials, tbat contacted human bone marrow in vivo and can answer the question of biocompatibility.

During the last decade various cell cultures were introduced in order to test tbe biocompatibility of materials. Nevertbeless it is negotiable if tbe results are always transferable into clinical application. Especially tbe testing witb animal cell cultures and cell cultures of tissues, tbat usually do not bave any contact to biomaterials is questionable.

476

Particularly on the problem of differentiation ot long term human bone marrow cell culture more research has to be done. Atter 4 wee":; cultured human bone marrow cells were differentiated into fibroblastic l1ke cells.

Caplan and Goldberq (14) demonstrated CE>lls t rom bone marrow cu] rurE> of rats, that were morphologically similar to the tibroblaBtic like cel18 in our experiment. He demonstrated a high cellular activltv of alkaline phosphatase. These cells cultured on HA and implanted into a rat toqether with the ceramic can produce osteoid.

REn:RENCES

1. Hench, L.L. and Ethridge, E.C., Biomaterials - file intertacial problem., Adv. Biomed. En~, 1975, 5, 35 - 150.

2. Hench, L.L. and Ethrldge. E.C., Biomaterials: An Interfacial Approach, Biophysics and Bioengingeering Series-, _ V_QJ_L J. Academic Press, ~ew Yurk. USA, 1982, pp. 279 - 288.

3. Hench, L.L., Splinter. R.J., Allen. tI.C. and Greenlee. T.K .• Bonding mechanisms at the interface of ceramic prosthetic materials. J. Biomed. Mater. Res. Symp., 1971, 2. 117 -141. ---------.

4. Hench. L.L. and Paschali, H.A., Direct chemical bond of bioactive glas8 - ceramic materials to bone and muscle. <L._~iQmed., ___ M~~gr_, __ g_e~_!._.[y!ll.l!, ' 1973, 4, 25 - 42.

5. Jarcho. M., Gumaer, K.I., Doremus, R.H. and Orobeck, H.P. , ['i88ue. cellular, and subcellular events at a bone - ceramic interface, J. !lioeng., 1977, 1, 79 - 92. ----

6. Oqiso, M., Kaneda, H., Arasaki, J. and Tabata, T., t:pitheilal attaehDlent of bone tissue formation to the surface of hydroxyapatite ceramics dental implants, in Advances in Biomaterials, Vol. 3, Wilev, Chichester, U.K., 1982, pp. 50 -- 64. ----- -

7. Bagambisa, F.ß., Die biologischen Reaktion~D_~yt de~ Knochenersatzwerkstoff HydroxvlaDatit - Keramik, PhO Thesis, Munich University, FRG, 1986.

8. Stoker, M., O'niel, C., Berryman, S. and Waxsmann, V., Anchorage alld growth regulation in normal and virus transformed cells, lnt. J. Cancer, 1968, 3, 683 - 693. ---------- -

9. Maroudas, N.G., Growth of fibroblasts on linear and planar anchoralles and limiting dimensions, Cell Res., 1973, 81, 104 - 110.

10. Folkman, J. and Mascona, A., Role of cell shape in growth cont.rol, Nature, 1978, 273, 345 - 349.

11. Weiss, R.E. and Reddi, A.H., Appearence of fibronectin during the differentiation of Cartilage, bone and bone marrow, J. Cell Bio~, 1981, 88, 630 - 636.

12. Stutzmann, J. and Petrovic, A., Bone cell histiogenisis: The skeletoblast as a stem - cell for preosteoblasts and for secondary type prechondroblasts, in Factors and Mecha~isms Influenct~one and Gro~th, Alan R. Liss, NY, USA, 1982, pp. 29 - 42

13. Dexter, T.M., Spooncer, E., Simmons, P. and Allen, Long Term Bone Marrow Culture: An Overview of Teehniques and Experienee, Long Term Marrow Culture, 1984, 57 - 96.

14. Goshima, J., Goldberg, V.M. and Caplan, A.J., The Osteogenic Potential of Cultured - Expanded Rat Marrow Mesenehymal Cells Assayed in Vivo in Calcium Phosphate Ceramie Blocks, The Clinical Orthopaedics, 1991, 262, 298 - 311.

477

FINITE ELEMENT ANALYSIS OF A CERAHIC HIP-JOINT HEAD AND ITS FAllURE HODE DUE TO A CRACK IN THE MATERIAL

E. RAVAGLI - ENEA, Centro E. Clementel - Bologna E. MAGGIORE - Centro di Bioingegneria, Universita' Cattolica - Roma

ABSTRACT

The aim of this work is the mechanical analysis of an alumina ceramic head for artificial hip joints,in order to evaluate the possible modes of failure. The geometry se Jec ted i s based on the one of an • in vi vo" fa i Jed head. Modules from the french CASTEH Finite Elements Code have been used for mecha­nical analysis within the elastic field. Calculations have been performed looking for the rupture load in a first step, under the hypothesis of no error in the conical matching between stern and head. Then, the influence of deviation from nominal dimensions of the matching surfaces was evaluated. Finally the possibil ity of failures, due to cracks in the material, was evaluated using the linear elastic fracture mechanics theory.

INTRODUCTION

Within the cooperation agreement on materials for biomedical applications between ENEA and Rome Catholic University, we intended to study the possible causes of a alumina ceramic head failure. The head was part of an artificial hip joint implanted in a patient which failed "in vivo". The schematic section of the head is shown in fig. 1, whereas fig. 2 shows one photograph of the broken head.

1. Basic hypothesis We did not have the exact dimensions of the conical stern portion to match the head (in particular the exact angle of the cone) because this stem remained in the patient's femur medullary canal. The stern has been supposed to be rigid, indeformable, whereas the following mechanical properties have been assumed for alumina: Young's module = 380 GPa; Poisson's coefficient = 0.26 ; Ultimate tensile stress (U.T.S.) = 350 HPa; Ultimate compressive stress (U.C.S.) = 2150 HPa.

For simplicity's sake, we have assumed an axial symmetry for the geometry and the loading pattern. Therefore the axis of the stern, of the head and of the acetabulum were assumed to coincide with the P load 1 ine of action. The stem-head taper fit has been assumed to be characterized by a 0,1 frict­ion coefficient (1). Finally, the head failure has been considered possible if the U.T.S. (350 HPa) is reached in connection with a P load value fram 3

I I

I--J. ~.L ~ 1~

I

Figur~ I. H~ad sch~matic s~ction

478

Figur~ 2. R~tri~v~d sampl~s of th~

brok~n h~ad

to 7 tim~s th~ body I'\/~ight; in fact, this is th~ ord~r of magnitud~ of the loads th~ hip-joints ar~ subj~ct~d to, und~r normal walking conditions (2).

2. Calculation Codes Used Th~ analysis has b~~n carri~d out by using th~ Fin i t~ El~m~nt Calculation Cod~s of th~ CASTEt1 Syshm, in th~

~lastic fi~ld and in 2D g~om~try (3), (4),<5), (6).

3. Hesh G~neration On th~ basis of th~ axial symm~tr>'

hypoth~sis, W~ consid~r~d th~ 2D structur~ shown in fig. 3, inshad of the space structure. By means of the COCO module, we hau~ divid~d the structur~ into m~sh~s made up of quadrangles, ~ach having 8 nodes. By this m~thod w~ hav~ obtain~d a network of 440 meshes and 1423 nodes.

4. First Step of the Hechanical Analysis The mechanical analysis of the structure has be~n carried out by means of th~ INCA module, under th~

hypothesis of no ~rror in the st~m-h~ad tap~r fit. As a consequ~nc~ th~ P load, appl i~d according to th~ stern axis, is transmi t ted to a h~ad annul a,r. area, corresponding to a length of generatrix I = 13.2 mm , as a pr~ssure that, if

r

Fig.3. Mesh of the 2D structure and th~ nodes 1 to 8 next to th~ conical hole of the head

479

there is no friction, we indlcate by Q (see fig.4). If there is friction (f = tg'f= 0.1), ~~e shall have to consider a Q' = q/cos'f , whose component, according the head axis, is:

Q'v Q' sin(o(+'j') = q cos~ <tgoi+tg'f) q cos .. (1/15+1/10) = q/6 * cos~

~ being the opening semiangle of the head conical hole (corresponding to the angle of the cone on the stem) and ~ the friction angle. The connection between the P compression load appl ied to the stem and the q pres­sure, found by considering the force balance according to the head ax is, is:

q = P/97.55 (q in t1Paj P in N) (1] Fig.4. Distributed load, applied

to the head, i f there i s no taper fit error

At this stag~,to the 1 generatrix of the structure we have appl ied a distri­buted load (in N/mm) , still expressed by [I] and of variable value, until, by means of se'Jeral attempts, a stress equal to the U.T .S. (350 MPa) has been reached in the structure it self. To be prec ise, three possible failure crite­ria have been taken into account, each referring to:

a) maximum stress; b) VON t11SES max imum stressj c) TRESCA maximum stress.

As for qual it y , the mechanical analysis has led to the follOl}/ing main results (see fig.5): - the greatest stresses are tangen­

tial stresses and are reached in nodes next to thO? hole inlet of the headjin particular the maximum stress is reached in node I (see figures 3 and 4);

- a relative maximum stress coinci­des with the bottom edge of the head hole (node 2): it is an axial stress.

As for quantity, the follow ing table shows the most significant results obtained by using the three failure criter i a. 1f we compare these results, it is evident that the c) criterion is the most conservat ive, in that

kg) • t imes

'" --

=f i / " 100 ... ,

:::t -. ------'==;::=;:=-::::-:=:;=+--T' ':=.!-

Fig. 5.<\,l)z,6'~for nodes next to the head hole (criterion a)

the head failure occurs in coinci­dence with the lower compression load appl ied to the stem (P=1159 HOIHever, also this load is 17f18 the body weight, therefore it is actual failure.

too heavy to propose the hypothesis of an

Consequently, we come to the con­clusion that if there is no error in the stem-head taper fit. it is not possible to propose the hypo­thesis of a failure in the head itself caused only bY bodY weioht in normal walking. Besides it is to be noticed that the ratio between the stresses in node 1 and the ones in node 2 ranges from 1.5 to 1.9, according to the failure criterion used.

480

5. Second Step of the Mechanical Analysis At this second stage we have carried out the structure mechan ical analysis, working on the hypothesis of an error in the stem-head taper fit. Two possibil ities have been taken into account: 1) top angle of the stern larger than

the one in the head hole; 2) top angle of the stern smaller than

the one in the head hol e. At this stage the following ~Iypotheses have been proposed: - the greater the error i n the stem­

head taper fit, the smaller the width of their contact annular area;

TABLE

.1

q (Mh) 173

P (kaI 1121

b".node 1 'NP.) ~ 6"V,M . node 1 (NP.) '60

OT . no4. 1 ("Pa) 020

O"':r node 2 ("Pa) 237

<S"'V,M . node 2 (NP.) ... UT . nod. 2 ("Pa) 272

G}""". 1 1 . 48

~ U """. 1

V.M

<S" "ode 2 1~87

V. M.

u;.. "ode 1 1. ln

Ü. node 2 T.

dl.pl. node 1 (./" M' 7.8

., cl

132 117

1308 11'"

266 23.

~ 310

39. ~

180 160

187 188

207 183

1 •• 8 1.,48

1 . 1J1 1. 87

1. 91 1.91

'.' '.3

- this annular area is loaded with a triangular load, whose maximum value coinc i des with the theoretical con ­tact eirele ( node ! in figs.3 and 6, ease 1; node 3, ease 2), then i t deereases 1 i near 1:, un ti 1 i t van i shes in coineidenee with the othe r end of the eontact annular area.

Fig.6.Triangular load applied to the head in ease of taper fit error

5.1 Top angle of the stern larger than the one in the head hole Aecording to the hypotheses proposed, the P load appl ied to the stern is

transmitted to a head annular area by means of a triangular load assuming its ma x imum q value in eoincidenee with the hole inlet edge (node 1). These load effeets are equivalent to the ones of a qeq pressure aeting on the whole annular area and related to q by: qe~ = q/2. Proeeeding in the same way as in paragraph 4, we ean find that tHe relat ion between P and the maximum ualue q of the triangular load is:

q = 57.55 PI 1(225.5-1) (q in MPaj P in N; I in mm)

At this stage, to the I length generatri x of the strueture shololln in fig. 6, we have appl ied a triangular load, whose ma x imum q value (in Nimm) is still

481

expressed bY[2]. By means of several attempts, we have changed the q value until a stress equal to U.T.S. (350 MPa) has been reached in the structure. ]n order to analyse also the influence of a taper error on the failure resi­stance, two different s izes of the contact annular area have been considered too: - case 1: the contact occurs between node 1 and node 4 (see figures 3 and 6)

namely for a length of generatrix 1 = 4.811 mmj - case 2: the contact occurs between node 1 and node 5, namely for a length

of generatrix l' = 1.203 mm. As for quality, the mechanical analysis has led to the following main results (see figures 7 and 8):

the maximum stress (traction tangential stress, in case 1, and cornpression radial stress, in case 2) is reached in node Ij

- in case 2, the maximum traction tangential stress is reached in node 5; it is definitely lower than the maximum radial cornpression stress reached in node 1, nevertheless its value is sufficient to allow U.T.S. (350 MPa) to be reached in node 5, when the compression stress in node 1 is much lover than U.C.S. (2150 MPa): therefore, also in case 2 the head failure is a consequence of the traction tangent i al stress;

- a relative maximum stress coincides with the bottorn edge of the head hole (node 2): it is an axial stress. However this stress is much lower than the maximum stress in node 1 (and also, in case 2, of the maximum traction tan­gential stress of node 5).

IOD ._ U"f' ,., ......... I

300 f' ... i ............ G"~

i

IOD - --:: -- ,V!----·JCtO

------_1'

-~oo •. ~---;~' --y--.... - .... ---,--.;r-.ur----;·[-!-='-,'

Fig.7.~,~,6'~fol" nodes next to the head hole (case 1 - critel"ion a)

1.0'c-:::-,....,.,,~,..-, --=.::;; .. ":-':.:-0:,.:----.--, _---,.------,.-----r----r----,

10 ~ r ...... ~ ... i ~ --= __ . _ .. __ . .!.f. ---------f.J

· 1 ...

~'r--x--...--r--r--I-"l.-"

Fig.8.~,~,().".. for nodes next to the head hole (case 2 - cl"iter ion a)

As for quantity, the tables 2 and 3 (case 1 and case 2) show the most signi­ficant I"esults obtained by us ing the thl"ee failul"e critel"ia. If we compal"e the pr'evious I"esults, it is evident that the c) cl"iterion is the most consel"vative, in that the head failul"e occurs in coincidence with the lowel" stern cornpl"ession load (P=330 kg in case 1; P=408 kg in case 2). ]n the two cases considel"ed, this load is respectively 5 and 6 times the body weight, and therefore it falls within the load values nOl"rnally foresee­ablej therefol"e the hypothesis of an actual head failure can be proposed in both cases. Consequently, we cerne to the conclus i on that if there is an er­ror in the stern-head taper fit, with the top angle of the stern larger than the one in the head hole, it is possible to propose the hypothesis of a failure in the head itself. this failure being the conseguence of body weight in normal waiking.

482

TABLE 2 - Case I TABLE 3 - Case 2

.. ., cl .. ., cl 11 ,-., ... n.' 0<'

11 ,.,., JA' :po, "t " (111 1 1 ... .' .. <r. "..... , ,..,.) l2!l: "" ''B

• 1111 _' '" ... -6:f1" .. ::: ~ ~:: I .. " ·9" .... .. , 'On ...

a.-; ..... , , .... , 107) "" , ...

G", • ..... 11_. 1 - l:!!! ,

(J t ..... \ C". I - - l2!l:

6"~ ...... ~ t!I"'al !..~ ". ,.. r.., ..... .,(" .. , >1' l2!l: >1' , . ,., l2!l: G": ..... ",_. ) , ..

T

<r, ..... J ' ..... l 'l Ol ,. 'v ..... t , ... , "

., ,. , . 6 ",... öl tlllP.' ,. .. ...

T.

(). ...,2'''''. ' .. • t .. a: ,. .... , ,..,... .. .. .. ,. .. ..

G"T .-.. , ..... , "

--(j"~-' •• t '.' '.' I~I 11.0 '0.0 '9 .. 0

'" .. "'" 1 (f .. ...... ~ C" -' 21 .. 1

-'-' -- oll. " ~1 • .,

c;:.,.,..,. t '.' t.· ... G: -' , .. by . lIII ....., 1

5" -' U . I T n.II' U . I:I

bf~. 1 ., .. , ' . \ ~ 6"T

-;;;;;;:;-> 6'* nod. ~ '.' l.t ' ..

..... pl _., 1,..1 , . , ... .-

6"~ G" -., '. , 1.' '. , ~,-._._-

r -' , . 6"T"" '".t '.' '.' , .. UT noM 2

.. '."l . .... ' .,..' 0.' .. , I . t

However if we compare the results now obtained to the ones obtained in the first step (no taper error) we ean see that the stresses at the hole bottom (node 2) has a much lower value in comparison with the ones at the hole in­let, and apparently this makes it more difficult to explain the kind of breakage observed in the head.

5.2 Top angle of the stern smaller than the one in the head hole The hypotheses proposed lead to the consequence that the P load, appl ied to

the stern, is transmitted to a head annular area, having 1 width, by means of a triangular load assuming its maximum q value in coincidence with node 3 (see fig.6). Proceeding in the same way as in paragraph 5.1, we ean find that the relation between q and P is expressed by:

q = 57.55 PI 1(199+1) (q i n HP a; P i n N; I i n mm)

At this stage, to the 1 lenght generatrix of the structure shown in fig. 6, we have appl ied a triangular load, whose maximum q value (in Nimm) is still expressed by [3]. By means of several attempts, we have changed the q value until a stress equal to U.T.S. (350 HPa) has been reaehed in the structure. The contact has been supposed to occur from node 3 to node 6: this situation can be considered as representative of a large taper error.

As for qual ity, the mechanical analysis has led to the following main results (see fig.9):

the greatest stresses are eompression radial stresses; they are reached in the nodes of the two elements loaded; in particular the maximum stress is reached in node 7 (see figures 3 and 6); as for the traction stresses, they reach their maximum value simultaneously

483

in node 8 (tangential stress) and in node 2 (radial stress); theyare definitely lower than the maximum radial cornpression stress reached in node 7, nevertheless their value is sufficient to prevent U.C.S. (2150 MPa) frorn being reached (even if it i s nearly so), when U.T .S. (350 MPa) -j

is reached in nodes 2 and 8; ~D

- in particular in node 2 the VON MISES_ and the TRESCA stresses are greater than in node 8, therefore it is this

1151'( M'" )

"7,' ~ __ -=-=-=·-=·-_-·G-~--~-·-=-~~.

\/

node 2, at the hol e bottorn edge of .(_J

the head, tha t can be cons i dered par- -lllU_1 • ~o <t <~ ~, ticularly critical for the failure Fig.9.~,~,Ö~for nodes next to of the head itself. the head hole (criterion a)

As for quantity, the most significant results obtained by using the three above mentioned failure criteria, are reported in the table 4. If we cornpare the results obtained, it is evident that the c) criterion is the

most conservative, in that the head failure occurs in coincidence with the lower stern cornpression load (P=1062 kg). However, also this load is about 6 times the body weight, therefore it is too heavy to propose the hypothesis of an actual head failure. Consequently, we corne to the conclusion that if there is an error in the stem-head taper fit. with the top angle of the stern smaller than the one in the head hole, it is not possible to propose the hypothesis of a failure in the head itself caused only bY bodY weight in normal walking. However this latter analysis is the one, among all the others previously carried out, that appears to have the stress di­stribution most suitable to explain the Kind of break occurred in the head: now in fact it is just the hole bottorn area (node 2) where the most critical conditions are reached. Critica!'conditions (almost at a failure point) are also found simultaneous­ly in node 8 and in node 7, with a small area of transition frorn the greatest traction stresses to the greatest cornpres­sion stresses. The above rnentioned reasons have induced us to continue to step 3 of the study, to analyze

q (MPa)

P (ka'

6"R "ode 7 (NPa) 0. ROde 7 (MPa)

V.M. 6': node 7 (HPa)

T.

Ö.,. ROde 8 (MPa)

(r, node 8 (MPa) V.M. 6': node 8 (MPa)

T.

6",. ROde 2 (NPa)

6': nod. 2 (HPa) V.M. 6' node 2 (HPa)

T.

e- Rode 7 V.M.

G;.M.node 2

V. nod. 7 T.

Ü T• Rode 2

Ö .. node 8

() R Rode 2

6"T node 8

(;'1: Rode 2

dbpl. node 7 ~I

TABlE 4

01 bl cl

3028 2834 250' 1285 1202 1062

-2118 -1982 -1752

1466 1372 1213

1690 , .... 1395

~ 328 290

343 321 28'

423 3.3 311

~ 328 2.0

374 ~ 309

311 3 •• ~

•• 0 •• 0 ..0

3 •• 3 •• 3 ••

4.' '.0 ',0

0.92 0.92 0.92

1,1 0.B9

•• 0 8.'

if the concentration of stress caused by a side crack, which is supposed to be around node 2, can be a possible explanation for the failure occurred.

6. Third Step of the Hechanical Analysis A side crack at the hole bottorn edge (node 2) has been taken into considera­tion. The position of the crack is shown in figures 10 and 11 (the crack front is in node 9). This position airns to explain a break of the same

Figur~ 10. 2D head mesh with the s i de crack at node 2

484

Figure 11. Crack detail

kind as the one obserued in the retrieved specimen. The hypotheses suggested at thls stage of the study (which is based on the elastic linear fracture theory) are considered to be conservative and they consist in: -suppos I ng tha t the P charge app lied to the s tem has a va I ue of 500 kg

(namel y it exceeds the limit of the value supposed, under normal walking conditions);

-assuming as fracture toughness the value K1C = 2.7 MPa m1/ 2 • In the case of no taper error in stem-head matching, we obtain,by means of successive steps of chain COCO-INCA-MAYA, that the crack critical size is: a = 1.945 mm. Still, this size seems too big to account for the fact that the fault has not been noticed during the tests carried out on the head fragments therefore, in the case of no taper error. the presence of a crack cannot be the possible cause of the fracture. The case of a taper error, where the stern angle is larger than the one in the head hole, has been immediately discarded, because the modules chain of the CASTEM provides values of more than 3 mm for the cr i tical crack. In case of taper error, where the stern angle is smaller than the one in the head hole, the modules chain of the CASTEM provides a critical crack, whose s i ze is: a = 0.670 mm. This size seems reasonable (taking into account the sensitivity of the test method used) and therefore it is possible to sug­gest that the failure occurred in the head was due to the occurrence of a crack. in the position exam ined before. and a taper error (stern head angle< hole angle) in the stem-head matchina.

7. Conclusions It is necessary to complete the third step of the mechanical analysis, by studying the behaviour of the head with the crack located in other positions of the structure. However, at this preliminar stage of the study, we have already found out an acceptable hypothesis for the head failure observed.

REFERENCES

1. Andrisano A.D., Dragoni E., Strozzi A., Anal isi meccanica di testine in ceramica per protesi d'anca. Ceramica Informazione, Dttobre 1988, N. 271.

2. Paul J.P., Force Actions Transmitted by Joints in the Human Body. Proc. R. Soc. Lond. B. 1976; 192: 163-172.

485

3. SYSTEME C.A.S.T.E.M. - eoco - Compagnie Internationale de Services en Informatique, Paris, Avril 1980.

4. SYSTEME C.A.S.T.E.M. - INCA - Manuel d'util isation, 15 Octobre 1985.

5. SYSTEME C.A.S.T.E.M. - ALICE - Manuel d'utilisation, Mars 1986.

6. SYSTEME C.A.S.T.E.M. - MAYA - Manuel d'util isation, Janvier 1984.

486

RESEARCH, PLANNING, AND DESIGN OF EAR OSSICLE PROTOTYPESBASED

ON Al~3' HYDROXYAPATITE, AND zr02

P. Laudadio (Div. Otorinolaringologica, OspedaleMaggiore, Bologna, Italy)

L. Presutti (Div. Otorinolaringologica, OspedaleMaggiore, Bologna, Italy)

A. Ravaglioli (IRTEC-CNR, National Council of Research, Faenza, Italy)

R. Martinetti (Finceramica s.r.l., Faenza, Italy)

Despite enormous progress over the past few years, middle-ear microsurgery still presents controversial problems solved by different methods depending on the different authors.

The two main types of intervention in this sector are those to remedy chronic inflammation and otosclerosis.

Chronic purulent otitis media is an affeetion which causes various degrees of damage to both the tympanum-ossicles system and the bone walls of the middle ear. In partieular, choleastatomatic ehronic otitis is a slowly developing illness which can lead to very serious eomplieations due to the presenee nearby of delicate struetures elosely related both to the tympanie and the mastoid cavity. In this ease, treatment must have two fundamental aims: 1) elimination of the inflam­matory proeess; 2) restoration of anatomieal and funetional eonditions ae similar as poseible to the original ones. Both of these objeetives ean be reaehed only by recouree to eom­plex microsurgical techniques, and often more than one opera­tion may become neeessary. In general, two types of inter­vention are carried out: one purely for treatment, and a second one some months later for functional purposes. These operations are called two-stage tympanoplasties and were introduced quite recently. Until a few years ago, in fact, the eommonest surgical intervention was radical mastoidect­omy, whieh basieally eonsisted in eonstruction of one large

487

tympanie-mastoid eavity eonneeted to the external auditory meatus. The hearing loss, however, was eonsiderable. The trend today is to abandon radieal mastoideetomy in favour of tympanoplasties of the middle ear, with the aim of restoring normal eonditions both anatomieally and funetionally.

The applieation of biocompatible materials in surgery of chronic inflam~ation ean either be: 1) As fillers of the mastoid eavity in the surgieal reeon­struetion of previously existing radieal cavities, or, with regard to tympanoplasties, in open interventions by oblitera­tive techniques. In both eases it is neeessary to fill a bone eavity with a material able to guarantee stability in time, not liable to give rise to phenomena of ejection or migration, weIl tOlerable, and preferably also able to promote normal bone development. These characteristics are ensured by a mixture of 60% hydroxyanatite and 40% tricalcium phosphate in the form of granules sized 0.5-1 mm in diameter and presenting a micro- and macroporosity for efficaeeous osteocoaleseenee. Over the past four years we have utilized this bone substitute in some 50 surgical operations, obtaining very favourable results in a high percentage of eases. All results are evalu­ated elinically and radiologically (by C.T. scanning). 2) In the form of ossieular prostheses.

The current tendeney of mierosurgery of chronie otitis is to try to restore a satsfactory hearing funetion. Hence the need to have at disposal materials able to replace one or more than one middle-ear ossicle in the event of damage

488

caused by the affection. The orientation of a number of authors today is to use either autologous or homologous ossicles (remodelIed, if needed). Other authors prefer to use biocompatible synthetic products. If in fact on the one hand it is true that autologous or homologous materials ensure fairly satisfactory results, on the other hand these products are not always easy to find, or they may be available in forms unfit to adapt to certain specific requirements.

Experiments have been carried out over the last few years on numerous materials for ossicu1ar urostheses, such as ~lastipore, polycel, ceravital, proplast, etc. Prostheses made from such materials are substantial1y similar in shape and can be schematica11y classified according to two models: 1) The TORP, resembling an urobrella in shape and fit to re­p1ace the entire ossicu1ar chain. It functions in much the

489

same way as the columella in birds in that it establishes direct contact between the tym~anic membrane and the oval window, eliminating the amplifying mechanism represented by the incus. Actually, this system is very efficient and can produce results approximating to normal conditions. 2) The PORP, a ~artial nrosthesis for substitution of the incus alone. It is therefore utilized in the presence of an undamaged stapes, which can thereby be put into contact with the malleus.

oVe are currently carrying out a study for development of TORP and PORP ~rostheses made from the fol1owing materials:

zr02 ZIROONIUJ',' OXIDE

A 94.6% pure oxide stabilized by addition of 5.27% Y2~ (yttrium oxide) of polycrystalline tetragonal structure. The planning of nrototy~es was made possible by an optimum plasti­city of the mixture, obtained by the use of po1yacrylates. Pre-firing was conducted at 100000 (thermal cyc1e 1000Ch ) and final sintering at 150000 (thermal cyc1e 20000 h). Shrinkage is equal to 28%, with qq.3~ relative density.

A1 203 .\LmrINA

A q9.7% pure oxide with an alpha-A1203 structure. The planning of nrototypes Nas made possible by an optimum plasticity of the mixture, obtained using sodium lignin sulfonate. ~re­-firing was at 115000 (thermal cycle 10000h ) and sintering at 15000C (thermal cycle 1000Ch ). 3hrinkage was of 18%, with a relative density of Q8.1%. ~echanica1 strength under compress­ion: 3596 ! 305 ~Pa.

HA h~OROXY~P~TITE

A QO.5% pure hydroxyapatite, appronriately calcined -- before treatment -- with sodium lignin sulfonate to render it suit­ably plastic for the realization of prototypes. Pre-firing at 11000C (thermal cyc1e lOOOCh ) and sintering at 1300°0. 20% shrinkage and q8.5% relative density.

~11 the mentioned ~~terials were used, but the greater part of the exneriments was on hydroxyapatite.

I\n ossicular prosthesis is exrected to answer the follow­ing main requirements: a shane nerfect1y adaptab1e to sur­roundin? str~ctures and easy to maneuver for nositioning;good biofunctionality, to ensure adequate sound transmission to inner-ear structures; good biocompatibility, to ensure stable results in time and avert the risk of ejection; the possibi1ity to be modified in the intra-operative stage to be adapted to the different needs. Strength under pressure is not very im-

490

portant, because vibration of the tympanum-ossicles system is obtained by application of low-energy compression waves and rarefaction waves.

Microsurgery of otosclerosis has to face slightly differ­ent problems. As is known, otosclerosis produces more or less serious deafness due to cancellous bone ingrowth along the borders of the oval window, with consequent ankylosis of the stapes on the oval window. Treatment consists in removing the stapes and replacing it with a prosthesis. Such prosthesis is generally positioned onto the oval window after this has been covered with a portion of vein to protect the delicate inner­-ear structures nearby. It is therefore necessary that the prosthesis should be somehow fixed to the incus on one side and that it should rest on the oval window on the other side. If performed skilfully, this intervention gives very good results in a high percentage of cases, enabling total resto­ration of the hearing function. There is, though, a small number of unsuccessful outcomes due to a variety of factors, including the type of prosthesis utilized. For example, a prosthesis exerting too tight a grip on the long process of the incus may cause it to undergo necrosis; an incorrect fixation may result in detachment in the event of traumas; a non-compatible material may be the cause of reactions from an extraneous bOdy, or it may result in granulation or ex­pulsion; an excessive weight of the prosthesis may provoke bedsores, dislocations, etc.

The hydroxyanatitic stapedial kind of prosthesis that we currently use avoids many of the mentioned problems, because in addition to utilizing biocompatible materials it is design­ed in such a way as to be perfectly adaptable to the lenticu­lar process of the incus. There is therefore no necessity to apply metal hooks. Furthermore, hydroxyapatite allows post­-operatively to perform radiological controls by high-resolu­tion C.T. with the aim of detecting the position of the prosthesis with respect to the oval window and the incus.

In conclusion, we believe that the introduction of ceramies represents a decisive step forward for middle ear microsur~ery and that their usa will become even more wide­spread in the futureonce current clinical and laboratory studies are adequately verified and confirmed -- a process which in this sector must take a minimum of five years.

491

PRELIMINARY TESTS '1'0 DETERMINE '1'BE INFLUENCE OF S'1'ERILIZA'1'ION AHD STORAGE

ON COMPRESSIVE STRENGTB OF BYDROXYAPA'1'I'1'E CYLINDERS

B. GASSER, W. MÜLLER and R. MATHYS JR.

MATHYS LTD Bettlach, CH-2544 Bettlach, Switzerland

ABSTRACT

The clinical application of calcium phosphate ceramics as block material demands for definition and determination of mechanical properties. The aim of this study was to characterize the mechanical behaviour of porous hydroxyapatite cylinders and to evaluate the influence of sterilization and storage on their compressive strength. Cylinders with a porosity of 60% and with 14 mm in diameter and length each were axially loaded to failure after sintering, after sterilization in the autoclave, and after storage under simulated in vivo conditions for 1, 3, 6 and 12 months. The material behaviour of the cylinders was brittle. The original compressive strength of ab out 19.0 N/mm2 after sintering was reduced by more than 10% and 30% after sterilization and storage for 6 months respectively. This decrease in compressive strength for these two steps was found to be significant (w=O. 95) .

I NT RODUC'1' ION

Resorbable and nonresorbable bone substitutes based on calcium

phosphate ceramics become of more and more interest. These

artificial "bone grafts" may have the form of granules or

blocks of different shape. The porosity, the size of the pores

and the size of the granules can be varied (1). Granules are

mainly used to fill up defects in dental or orthopaedic

surgery. Ceramic blocks additionally fulfill a mechanical func­

tion and have therefore, depending on their application, to be

492

adapted to the required shape and/or to withstand the acting

loads . These requirements demand for clear and reliable

definitions and determinations of the mechanical properties of

ceramic block material. With respect to the clinical

application, e.g. in the spine (2) or for reconstructive

surgery (3), the behaviour under compressive load is most

important.

The aim of this study was to characterize the mechanical

behaviour of porous hydroxyapatite cylinders under compression.

Additionally, the influences of heat sterilization and storage,

as a function of time, under simulated in vivo conditions on

the compressive strength should be evaluated.

MATERIAL ABD METBODS

Hydroxyapatite (Mat. I; pore diam. 200-400 ~m) and hydroxyapa­

tite with a content of about 10% tricalcium phosphate (Mat.

11; pore diam. 400-800 ~m) were investigated. Both materials

tested had a porosity of 60%. Manufacturing of the test speci-

Figure 1. Examples of the porous hydroxyapatite cylin­ders studied under compressive load

493

mens was done under clean room conditions by pressing,

isostatical condensation, cutting to cylinders and sintering.

The shape of the cylinders was finally 14 mm in diameter and

length each (Fig . 1) . The compressive strength of these

cylinders was then determined after sintering (sint), after

sterilization in the autoclave (no vacuum) at 134 0c (ster),

after sterilization and storage in Ringer's solution (39 oC)

for 1 month (1m), 3 months (3m), 6 months (6m) and 12 months

(12m) . 2 to 6 specimens of the two materials were investigated

in all groups . The specimens were axially loaded on a materials

testing machine with the speed of 0 . 5 mm/min. The

Figure 2 . Broken cylinder in the loading device of the materials testing machine

loading device (Fig. 2) was made with a ball bearing on one

side to correctly load the specimens . The resolution of the

load cell was 10 N leading to aresolution for the compressive

strength of 0 . 07 N/mm2 when considering the diameter of the

ceramic cylinders . The analysis for assessment of the

significance of the differences in compressive strengths

between the material groups treated and stored in a different

way was based on the Wilcoxon - test . This type of statistics

494

should take into consideration the small number of samples

taken at random and the unknown distribution of the average

values.

RESUL'l'S

Under the compressive loads applied no deformation of the

specimens could be measured, their behaviour was fully brittle.

Failure occurred as 2 or more cracks in a slightly oblique and

vertical direction (Fig. 2), indicating that shear stresses in

an oblique plane were the reason for failure. After sintering

the compressive strength of Mat. I (HA) and Mat. II (HA/TCP)

was 19.0 and 17.5 N/mm2 , respectively. A decrease in strength

Mat. Type

Mat I'

sint ster 1m 3m 6m 12m

Mat. Ir' sint ster 1m 3m 6m 12m

TABLE 1

Compressive strengths of Mat. I and II

No. of Specimens

6 6 4 3 3 4

2 5 4 4 4 4

Average [N/mm2]

19.0 17.2 17.0 16.3 12.9 15.1

17.5 14.9 14.6 13.4 10.0 13.1

S.D.

± 1.9 ± 1.3 ± 1.8 ± 0.6 ± 3.9 ± 1.4

± 2.5 ± 3.3 ± 3.6 ± 3.7 ± 0.9

Decrease in Strength

10% 11% 14% 32% 21%

15% 17% 23% 43% 25%

after autoclaving and with increasing duration of storage could

be found (Tab. 1). This decrease in compressive strength for

Mat. I (100% HA) was significant (w=0.95) between sintering and

495

sterilization and between sterilization and immersion in Rin­

ger' s solution for 6 months (Fig. 3). No significant diffe­

rences between any steps of treatment or storage could be found

for Mat. 11 (90% HA/10% TCP). The measured compressive strength

of Mat. 11 was always smaller when compared to Mat. I. However,

this difference was only significant comparing the two

different materials after autoclaving.

Significance of Differences: Mat. I

N/mm 2 (100% HA)

20

15

10

5

,.-.-.-t8:iS"'~ r-.

r'" " 1\ llj ~ ., Wilcoxon - test . I significant • not significant - - ~

si nt ster 1m 3m 6m 12m

Figure 3. Significance in compressive strength of Mat. I between the different kinds of treatment

DISCOSSIOH

These tests demonstrate the type and amount of strength

decrease of HA cylinders. The reduction due to sterilization

and storage under simulated in vivo conditions for different

length of time could be characterized as moderate. The in-

crease in strength between 6 and 12 months of storage in

Ringer's solution remains unclear. Maybe an additional study

with a larger number of specimens could lead to clarification.

The slightly smaller compressive strengths of Mat. 11 may be

496

caused either by the 10% content of TCP in the hydroxyapatite

or by the larger pore diameter of this material.

It is known that determination of mechanical properties of

ceramic materials demands for a large number of specimens. This

is demonstrated by some quite large values for the standard

deviation of the compressive strengths found by means of the

statistical analysis. Nevertheless, the results found are

considered to give a reasonable overview on the decrease in the

strength of these calcium phosphate based ceramics.

Autoclaving and storage, as shown, have a moderate influence on

the decrease in strength. However, the porosity of the material

is also a very important parameter. It directly influences the

structural strength and, additionally, allows for bone ingrowth

which increases the compressive strength of the material in

comparison to the original "bone substitute" with empty pores,

as it is known from clinical application.

However, compressive strength is not the only parameter needed

to describe the mechanical behaviour of ceramic block

materials. It remains open what other tests, e.g. notched-bar

impact or bending tests, should be performed to assess calcium

phosphate based ceramics. Maybe that a general guideline is not

possible or useful, and that testing will strongly depend on

the later clinical application of the material.

REI'EREHCES

1. Eggli, P.S., Müller, W. and Schenk, R.K.; Porous Hydroxy­apatite and Tricalcium Phosphate Cylinders with Two Different Pore Size Ranges Implanted in the Cancellous Bone of Rabbits. In Clinical Orthopaedics, July 1988, 232, pp. 127 - 138.

2. Magerl, F., Schenk, R. and Müller, W.; Klinische Erfahrungen mit geformten porösen Hydroxylapatitblöcken. In Conference Proceedings "Biomaterialien und Nahtmaterial", edited by Rettig, H.M., Springer-Verlag, Berlin, Heidelberg, 1984, pp. 53 - 60.

3. Hemprich, A. and Hidding, J.; Secondary Correction of Trau­matogenic Enophthalmos - An Indication for Hydroxyapatite Blocks.ln Clinical Implant Materials. edited by Heimke, G., Soltesz, U. and Lee, A.C.J.; "Advances in Biomaterials, Vol. 9", Elsevier Science Publishers B.V., Amsterdam, 1990, pp. 543 - 548.

497

A STUDY OF THK IIITHODOLOGY FOR TREATIIINT OF TITANIUN SUBSTRATES TO BE COATED WITH HYDROXYAPATrrE.

.. A. KRAJEWSKI. A. RAVAGLIOLI. R. IlARTDiK'lTI. C. IlAlfGANO

Istituto di Ricerche Tecnologiche per la Cen.ica deI C.N.R •• Faenza (Italy)

.. Poliambulatorio Odontoiatrico Venini. Colico - Como (Italy)

SUMMARY

The application of synthetic materials for dental implants is now weIl established. Many of these implants are made with titanium or titanium alloys. It was found that such implants, correctly designed, remain weIl osseointegrated under full load application. A covering with hydroxyapatite of the metallic dental roots was introduced a short time aga with the aim to avoid any mechanical irritation of the gengive/implant interface and to allow for a strong adhesion of bony tissue on the implanted dental prosthe­sis • Wi thout enter upon the subject about the success of this kind of implants, this study deals with the purely ceramic problem on how to impro­ve the bonding between ceramic coating and metal substrate by lowering the sand-blasting residue.

PRELIMINARY MONITORING

A large number 01' industrially produced dental root samples made up

01' a titanium core and a hydroxyapatite coating have been found to give ri­

se to various failures. including the detachment of the root i tself from

the implant site over time. Because this is a serious failure, a study was

started to investigate the reasons for i t. A commercial Dyne sampie was

examined extracted from a middle-aged female patient following detachment

from the implant site. On inspection, the sampie appeared to have lost its

hydroxyapatite coating. An analysis was consequently carried out on virgin

sampies identical to the original one taken from the patient and produced

by the same manufacturer; the aim was to evaluate the situation at the in-

terface and the consistency of the hydroxyapatite deposited on the surface.

Considering that among the main or concomitant causes suspected there was

498

also the occurrence of residues of sand-blas ted materials remained embedded

in the metal surface, a systematic study was started to try to find the

best possible solution for this kind of failure.

TUE AlM OF TUE STUDY

The objective of this study is to identify viable methods for trea-

ting surfaces of metal supports to be coated so that the surfaces in ques-

tion may allow the lowest possible embedment of sand-blasting residues. In

order to classify the different practicable treatments for subsequent se-

lection of the most suitable ones, a number of pure titanium specimens were

produced whose surfaces were treated in different ways. Such treatments

consisted in aseries of procedures which included sand-blasting and chemi-

cal etching (either before or after sand-blasting). Pure titanium was choo-

sen for this study since it is the softest metallic material utilized for

dental roots. A number of industrial companies were involved in this rese-

arch which were interested in improving the quali ty of plasma-sprayed

coatings.

TUE MATERIALS

The materials used :for this study were:

a) for preliminary investigations: a sampIe extracted from a patient and

two brand-new commercial sampIes of the same type (Dyne), both coated with

hydroxyapatite;

b) for conclusive investigations: a number of sampIes which included tita­

nium screws (0=4 mm, length=12 mm) and smooth sampIes (about 2x2 cm).

All the metallic materials found in the composition of the different sam-

pIes studied during the preliminary. and the conclusive stage were of 99.98%

TAB. I - Average composition of the spherular sand-blasting material called "Crystal". The nature of such material show to be prevalently glassy rather than castor-quartz as indicated by commercial specifications; the spherules have a nearly uniform grain di ameter of 250-3001»".

COMPONENT Si02 Cao Na20 MgO AlZ03 FC203 K20

wt% 75 13 13 5 1 0.5 0.5 (:l:10%wt%)

499

pure ti tanium of the same commercial biomedical grade type. The materials

used for sand-blasting were synthetic aluminas, corundum and "Crystal" sphe-

res whose composition - determined by both chemical and X-ray fluorescence

analyses - is ~eported in Table 1. The chemical substances utilized were all

Merk of RPE grade.

TUE PROCEDURES

Chemico-physical analyses were conducted by m~croprobe scanning elec-

tron microscopy and X-ray diffraction.

Except for the Dyne commercial sampIes, the surfaces of metallic substrates

were subjected to chemical etching in the ways indicated in Table 2. These

treatments were carried out using solutions of nitric acid and hydrofluoric

acid.

I AB. Z - Serie. ofireilmenlslbilibe sarfl<e oflillniam ......... 1 .. bl ............. n. ....... oflrell ...... comspo .... I .. Ihll ... ported ... rti". rro.loft. Sa ....... i .. condiliollS ...... : ...... dlsta ... • 40 .!Bö I ...... nt li .. • 10 mla; , ........ -3Alm.

SAMPLE SAND- CLEAN. !!~KING DIST. :t~~t w~~h ~Ubi 3O;%HJW' WITH SODAIt WATER CODE BLASTED C,H,CI, WATER WASH. AGENT CLEAN. SOL. SOL.

A - Y Y Y - - - -8 Y Y Y Y AI Y - -C Y Y Y Y CS Y - -D Y Y Y Y AI Y Y -E Y Y Y Y COR Y - -F Y Y Y Y COR Y - Y G Y ACd_.nd - Y AI Y - Y I:.lhanoi

H Y YII) Y Y AI Y - Y

TAB. 3 - A .. 1ysis of I. titul •• sarfl<es .. rrI .. 0lIl11)' ............. joIaed 10 SEM. Except Ibo case nd DYNA s.m,a.lbll ......... 1 ao Irell_1bo codes oflM 01 .... SImples comspo_ 10 _ reporlnd I. Tahle I.

ELEMENTS ANALIZED SAMPLES (Val .. s in Iloms %) (± stand. deviation) A 8'21 C 0''' E FI F2 G H DYNE

TI 2.61 99.37 90.83 98.10 90.36 90.61 98.33 99.21 94.65 77.32 70.23 AI (2.22) 0.32 9.04 0.84 7.36 8.66 1.02 0.38 5.00 21.29 1453 '51(0.10) 0.15 0.13 1.06 1.24 0.60 0.65 0.41 0.35 1.39 6.16 PCO.IO) - - - 0.21 - - - - - 1.(0)1

OCO.IO) - - - 0.20 0.13 - - - - 0.92"" S 0.10 0.16111 - - - - - - - - 0.1801

F. 0.20 - - - 0.40 - - - - - -Ca (0.20 - - - 0.23 - - - - - 5.92 K(0.20) - - - - - - - - - 1.46°'

TUE RESULTS

a) Results of the preliminary investigations.

C !H1Cl l • trichloroethylenc

(I) After trichloroethylenc clunins il is etched with JO'K. HNO, + 2% HF solution

Lcaenda: AI • AI20 1/60 UOOpm); es • ·Cryst"- spheres: COR - Corundum supercor

56(250 pm) Y • yes

(I) Sulphur tome from o .... nie sludp:s idcntified 15 sman spots or about 'pm diameter.

(2) On Ihc screw neck il i5 not sinaled out presence oe OIher elements lhaß titanium.

(3)> Tbe indicatcd elements may come from biochemieal deposition durißJ the time of implantation in the patient.

(4' AlthoUJh aluminium is detec:tcd on the sunac:e. no stuek AI!Ollrains are siqlcd out.

On examination, the Dyne screw-shaped dental root sampIe extracted

from a patient appeared virtually denuded of i ts original hydroxyapati tic

coating. A dental inspection of the extraction si te inside the patient I s

mouth revealed in fact very few remnants of the hydroxyapatite mantle still

500

adhering to bone tissues. Analyses of the whole surface of the extracted sam-2

pIe showed sparse hydroxyapati tic zones of about 15000 um wi th negligible

thickness. Microprobe analysis of the denuded surface detected a large amount

of Al, as shown in the last column of Table 3. A further analysis carried out

on the surface of the virgin screws showed that stuck A1 203 grains may be

present, taking into account the appreciable number of dark spots displayed

by electron backscattering (Figure 1). The gray area on the bottom-left side

of Photo 1b was ascertained to be a large spot of remnant hydroxyapatite (HA)

Figure 1 - View of an area of a Dyne Hample graf ted by a patient, by SEM. (a) Standard collection; (b) Back-scattering electrons. The dark spots correspond to Alumina grains stuck on the Titanium substrate surface; the left down greish area on the photogram (b) corresponds to a small r...ained piece of the original Hydroxyapatite covering.

Figure 2 - SEM microphotographs taken on a cross section perpendicular to the &Xis of a D~~e type virgin acrew. (a) Standard collection; (b) Dots microprobe analysis on Ca or P.

501

Photographs were taken of cross sections perpendicular to the axis of Dine­

type virgin screws (Figure 2). Microprobe dot analysis on Ca and P distribu-

tion (Figure 2b) allowed to identify the trend regarding the thickness of the

HA coating, ranging from a minimum of 20 um to a maximum of 50 um. Higher ma-

gnifications show a good adhesion of molten HA particles partially crystalli­

zed and a number of small cavities (close porosity). Hemispherical grains of

semi-molten particles impacted on plasma-spraying of some 20 um in radius can

be noted, while the cavities range from 5 to 9 um in diameter. The density of 2

such cavities is of the order of 70000 per cm

b) Results of the conclusive investigations.

The results of microprobing of the sampIes treated in accordance with

the specifications of Table 2 are reported in Table 3. The photographs of the

Figure 3 show the extent of Al203 grains embedded on the surface of each sam­

pIe as can be determined by comparing the real images (side a) with those

(side b) obtained by electron backscattering (EBS).

DISCUSSION AND CONCLUSIONS

The dissolution of the HA coating of the implanted Dyne screw arguably

results from lysis taking place during the post-operative stage, when resor­

bing cells, that cannot distinguish between HA of bone and of ceramic cove-

ring, take action to resorb damaged bone. The average durability of a HA coa-

ting depends on a number of factors connected wi th the quali ty of the HA

layer and depends as well on the kind of treatment performed on the surface

of the substrate before coating. The rate of resorption of it is presumably

a function of its thickness, its chemical nature, its degree of amorphousness

and the type and extent of porosity. Apart from the thickness of the examined

HA layer of the Dyne sampIes (too thin) , among the factors that can reduce

the lifetime of a HA coating is the existence of Al203 grains remained stuck

on the metallic substrate after sand-blasting. The role played by such grains

in weakening the stability of the HA layer derives from the different adhe-

sional energy between the coating and the metal or the Al203 grain respecti­

vely. Such points of discontinuity at the interface make possible the deve-

lopment of microcracks which, in conjunction with the porosity, allow permea-

tion of physiological fluids containing lytic substances that can operate wi-

502

fIa. 3 • Compuative view 0( microphotopaphI collec:ted by SEM on standard concIitIoIl (a serieI) aad on eIec:troa lNIct scatterina c:oIlection (' series). Tbe second capitaI Jetter refen to die umple code repertecI in Tabte I. ( 300x )

503

thin the ceramic layer and notably at the metal/ceramic interface. Presence

of Al203 grains was detected in all the the commercial sampIes examined. The

sand-blasted titanium sampIes prepared for this study show that the embedment

of sand-blasted grains is always detectable but it is possible to lower their

quantity. The Al detected by microprobe analysis is practically due to such

stuck Al203 grains; the photos of Figure 3 show a direct proportionality bet­

ween the Al quantity picked up by microprobing (Table 3) and the density of

the dark spots existing on EBS microphotographs. Such proportionality is cer­

tified by the straight line of the graph of Figure 4. Therefore, on the basis

of the resul ts of Table 2 and of the certi tude that all the Al detected

corresponds to embedded Al203 grains, it follows that the treatments perfor­

med on sampIes C and F are the most suitable. Regarding Q sampIes, use was

made of a quartz-bearing sand-blasting material, but practically no presence

of residual silica was nonetheless observed.

20

'" 15

...., c:: ::l 0

8 cO

10 ..:.: ~ '-'

~ cO ::c

5

"'" UJ III

5 10 15 20 Al? amounl from microprobe analysis

Figure 4 - Diagram showing the straight-li ne obtained by plotting the computed BSE black amount% on the a­lIOunt% of Aluminium coming from microprobe analysis.

In conclusion, to obtain a

substantial lowering of sand-

blas ted grains stuck on the

metal surface, it appears ne-

cessary the dependence on:

a) concerning the sand-blas-

ting material: crystal degree

and habit of grains (spheroi-

dal shapes seem to be better)

chemical nature and purity of

the main component;

b) concering the subs tra te :

the degree of pre-existing

roughness, i • e. imparted by

the machining of the metallic surface (smooth or slightly satinated surfaces

appears to be the best); the extent and quality of etching treatments (espe-

cially after sand-blasting operations) by acids which probably help the em-

bedded granules to come unstuck.

504

STRESS ANALYSES OF PULL·OFF TESTS FOR STRENGTH MEASUREMENTS OF COATINGS

U. SOLTESZ, E. BAUDENDISTEL, R. SCHÄFER Fraunhofer-Institut für Werkstoffmechanik

Wöhlerstraße 11, D-7800 Freiburg

ABSTRACT

For the determination of the mechanical strength of coatings for prostheses materials usually the pull-off test is applied. The results, however, are obviously influenced by the testing procedure. This Finite Element (FE) study shows that stress concentra­tions in the range of more than 10 can cause considerable errors. The effects of dif­ferent geometries and the material parameters is shown.

TEST PROCEDURE

Conventionally, the maximum adhesion strength'of a coating is defmed to be the aver­

age stress ä which can be applied in normal direction to the surface without damaging

the material,

F ä =7\. (1)

assuming that the force F is applied by a contact element (stud glued on the surface)

of cross seetion A and the stress distribution in the area A is uniform. Due to the

geometrie discontinuities and the differences in the elastic properties of the compo­

nents of the test arrangement, stress peaks are produced. These peaks can lead to pre­

mature failure of the coating and thus, to falsified results. In order to analyse the

stress distributions in the sampIe and to determine the height of the stress concen­

trations the Finite Element method was applied.

In this work, the stresses near the surface are considered. The interaction be­

tween the substrate and the coating were not part of the investigation. Therefore, in

the Finite Element models the substrate and the coating could be replaced by a homoge­

neous sampIe. Nevertheless, the results also represent the stress distributions in the

505

upper layer of a eoating. For the models, the elastie properties of a glass, Le.

(E = 63 GPa and v = 0.2) were assumed as an example. All arrangements are axially

symmetrie.

In the diagrams whieh show the stress distributions, the stresses are normalized

to Cf whieh is defmed aeeording to (1). The quality of a test arrangement will be spe­

eified by the stress peak level 0., Le. the ratio

(J

o.=~ Cf (2)

whieh should be c10se to 0. = 1. (J represents the peak value of the stress distribu-max

tion.

Non-symmetrie models

In this test proeedure, a support ring is used to hold the sampie when the foree IS

applied (Fig. 1).

F

ring

c:ontact zone

Figure 1. Non-symmetrie puB-off test (scheme and FE model)

In the eaIculations the dimensions (s = sampIe thiekness, d = stud diameter, I = stud

length, D = support ring diameter) and the meehanieal properties of the support, the

stud, the glue and the pad layer are varied. The FE-model for one of the eonsidered

geometries is also shown in Fig. 1. For reasons ofaxial symmetry, in the Finite

Element model only half of the sampIe has to be eonsidered.

506

Symmetrie models

In order to overcome the geometrie discontinuity at the edge of the support, in the

symmetrie models the support has been replaced by a stud whieh is attached to the

opposite side of the sampIe (Fig. 2a).

In a second modifieation the diameters of the studs and the sampIe are chosen to

be equal in size to study the influenee of the discontinuity in the diameters of the

previous model (Fig. 2b).

F F

F F

a) b)

Figure 2. Symmetrie puIl-off arrangements

RESULTS

Non-symmetrie models

Tbe material of the stud by whieh the force is transmitted to the surfaee of the sam­

pie, influenees the stress distribution. Fig. 3 shows the stress peak which is caused

below the edge of a stud with the same elastie properties as the sampie (E = 63 GPa).

Tbe stress nonna! to the surface is increased by a factor of more than 9 compared 10

the homogeneous stress state.

More eompliant (PMMA) or stiffer (steel) materials produce lower or higher

maximum stresses, respectively (Fig. 4).

Ir) Ir) G) .... -Ir) 0 E .... 0 c:

10

8

6

4

2

0 )

----,/ IIdge of etud ;

-21-~-,--~-r~--.-~~ 0.0 1.0 2.0 3.0 4.0

distance fram center [mm]

507

12

tS 10

Q)

> Q) 8

.::L. 0 6 CU Q.

Ul 4 Ul Q)

2 ..... ..... Ul

, , , • I I

I I

I , •

,

.. ______________ e

,

non symmetrie

I/d - 2 s/d - 0.2

1.0 2.0 3.0

E etud JE somple

4.0

Figure 3. Stress distribution in a non­symmetrie test (glass stud,

s/d = 0.2, Vd = 2)

Figure 4. Influenee of the stud material on the stress peak level a

(s/d = 0.2, l/d = 2)

The thiekness of the sampie and the length of the stud are also parameters which influ­

enee the height of the stress peak. For two l/d ratios and different s/d ratios,

(s = sampie thiekness, d = stud diameter, I = stud length) the stress peak heights a are shown in Fig. 5.

14 tS

12 Q)

> 10 Q)

.::L. 8 0 Q) Q. 6

Ul Ul 4 Q) ..... ..... 2 (J)

0

SBJ I/d - 2.0 tzJ I/d - 0.4

glue thickness ... 0

s/d

s: sam pie thiekness d: stud diameter I: stud length

Figure 5. Stress peak level in the non-symmetrie test (lid = 2 and Vd = 0.4)

In the models described until now, a rubber pad was used between the sampie and

the support. To study the influenee of the meehanieal properties of the support this

layer was removed and direet eontaet between the sampie and the support was assumed.

508

In this case, the stress peaks are reduced. This effect can be shown for the geometry

considered here and for a second arrangement in which the gap between the pull stud and

the support has been reduced, i.e. smaller diameter of the support opening D (Fig. 6).

es Qj > Q.)

~ 0 Q.) Q.

(/) (/) Q.) ~ +' (/)

12 ... compflOnt support

10 ... .tif1 aupport " -4.

,,'

8 '" ". 6 -,- ,--.-----<4 I/d - 2

s/d - 0.2 2

glue thickness ... 0 0 0.0 0.1 0.2 0.3 0.<4

gap width [mm]

0.5

s: sampie thickness

d: stud diameter

1: stud length

Figure 6. Influence of the support stiffness (Young's modulus and gap width) on the stress peak level <X

The characteristics of the glue layer which is used to fix the stud on the surface also

have a strong effect on the stress peak. The thickness of the layer and its Young's

modulus were varied. Differences in the range of more than a factor of 12 can be de­

tected when a short stud and a thin sampie are used (Fig. 7).

1<4

es • • • glu. thickn_ 0.02 mm 12 .... gl ... thiekn ... 0.05 mm

Q.) .... gl ... thickn ... 0.' mm > 10 ~ ~ 8 0

""./'

, • Q.) 6 a. . /

(/) <4 ~ .... ~/,. non symmetrie

(/) ~. " I/d - 0.<4 Q.) . ,-~ 2 s/d- 0.2 +' (/)

0 0.0001 0.001 0.01 0.1 10

E glue/ Esomp'e

Figure 7. Influence of the glue stiffness (Young' s modulus and glue layer thickness) on the stress peak level <X

509

Symmetrie models

All ealeulations show that in the symmetrie test the stress peaks are mueh lower than

in the non-symmetrie test and that the variation of the material and geometrie parame­

ters has a mueh lower effect on the results than in the non-symmetrie procedure. Never­

theless, similar to the non-symmetrie case, the height of the stress maximum is influ­

eneed by the stiffness of the studs and the thiekness of the sampie (Fig. 8).

3.0

Ö ...... s/d=O.8 ..... s/d=O.4

Q) 2.5 .... s/d=O.2

> Q) symmetrie

~ large sampie 0 2.0 Q) a. fIl fIl 1.5 Q) L ....... fIl

1.0+----r---,r-....--~__.--r--.-_i 0.0 1.0 2.0 3.0 4.0

Figure 8. Influenee of the stud material and the sampie thiekness on the stress peak level a (symmetrie arrangement)

1.6

1.4

III 1.2 III

~ .... 1.0 III

0 0.8

E 0.6 L-0 c 0.4

0.2

0.0 I I I I 0.0 0.5 1.0 1.5 2.0 2.5

distance from center [mm]

Figure 9. Stress distribution in the symmetrie arrangement - sam­pIe diameter adapted to stud diameter, steel stud, s/d = 0.4

The reduetion of the sampIe width W to the same diameter as the pull studs resuIts

in a further improvement of the stress homogeneity. For the ratio s/d = 0.4 and steel

studs which is shown in Fig. 9, the stress peak is only in the range of 1.1 and thus,

it is very close to the desired homogeneous stress distribution

CONCLUSIONS

In the evaluation of pull-off experiments, the test method has to be taken into ae­

count. In unfavourable arrangements, eonsiderable stress peaks ean be responsible for

the destruetion of the sampIe and thus, lead to ineorrect (lower) values of the

strength. Improvements can be aehieved by optimized parameters of the test arrangement.

The best results are found in a symmetrie test procedure.

510

CELL ADHESION STRENGTH TO BIOCERAMICS

AND ITS MATHEMATICAL MODEL

Tetsuya TATEISm, Takashi USHIDA Biomechanics Division, Mechanical Engineering Laboratory, Agency oflndustrial

Science & Technology, Namiki 1-2, Tsukuba. Ibaraki 305, Japan

ABSTRACT

The aim of this study is to measure adhesion strength of cultured cells to bioceramics such as alumina, and to adapt a mathematical model of defect growth kinetics to the cell adhesion phenomena. Fibroblasts from mouse (L-929) were cultured on alumina plates and fibronectin­coated alumina plates. The adhesion strength was measured by loading adhered cells with centrifugal force vertical to the material's surface. At the same time, we measured adhered areas and morphology of cells on the 2 types of surfaces by using an image analyzing method. The results show that the cells adhered more tightly to fibronectin coated alumina than to alumina. (50% of the cells were peeled off from the fibronectin-coated alumina under the load of 5000, while 50% of the cells were peeled offfrom the alumina under the load of 1000.) The image processed data show that the average of adhered areas of cells to fibronectin coated alumina was 3 times as large as to alumina 6 hours after seeding. The results show influence of fibronectin-receptor bonds on the cell's adhesion strength and the cell's adhesion phenomena. Considering the influence of fibronectin-receptor bonds, an mathematical model of defect growth kinetics was adapted to the cells adhesion phenomena, where we assumed that the rate of condensation of vacancies or the rate of rupture of fibronectin-receptor bonds at the tip of the crack was proportional to the radius of the crack.

INTRODUCfION

The interface between biomaterials and living cells is becoming one of the most important issues in the fields of orthopedics and bio-industries. Among the subjects concerning the inter­face, adhesion phenomena of cells to biomaterials would be one of the significant subjects in order to evaluate the biocompatibility of biomaterials. We have attempted evaluating this sub­ject from a biomechanical point of view through measurement of adhesion strength of cultured cells to biomaterials1) and quantitative of morphology of cultured cells on biomaterials3). It is inevitable for what is called biomaterials to contact with living cells in vivo. However, accord­ing to the position where the biomaterials are used, they are requested to have either of two different characteristics, one of which is not to make cells adhere to the surfaces such as a frictional surface of artificial hip joint, and the other of which is to make cells actively interact with the surfaces such as a surface of artificial hip joint's stem.

Several kinds of so-called cell adhesion factors such as fibronectin, vitronectin and lam­inin have been found. These cell adhesion factors possibly interact with the surface of bio­materials and their receptors in the cell membrane, and the interaction is said to influence the cytoskeleton by the transmembrane control.

511

In order to measure adhesion strength, a viscometric method was developed where cells were loaded with shear stress. In this study, cells were loaded with vertical stress to material's surfaces. The feature of this method is to be able to quantify the adhesion strength itself, be­cause vertical stress is directly applied to the mass of cells adherent to material's surfaces.

MATERIALS AND METHODS

Specimens Polycrystalline alumina (Kyocera, Japan) was sintered into plates (diameter: 30 mm,

thickness: 1 mm) so as to be put into cell culture dishes (#2500, Corning, U.S.A.). The sur­faces of the specimens were polished so that the surface roughness (Ra) was to be O.lj.1.Ill. The surfaces were also coated with fibronectin (Boehringer Mannheim, Germany) from human plasma. (1 mg I ml, lOOmlI plate)

Apparatus for loadin~ of yertical stress Figure 1 shows an apparatus for

loading of vertical stress, which was made of aluminium. The specimen (alu­mina plate) was mounted up-side down on a ringed spacer in a culture dish. The apparatus was suspended to a centrifuge (Hitachi, Japan) and then centrifugal force was applied to the adhered cells.

Cell cuhure Fibroblasts derived from mouse

connective tissue (L-929) was cultured under the condition of 37"C, 5%C02 con- Spacer centration in air. The culture medium was prepared with supplementing Eagle's MEM (Nissui, Japan) with 10% foetal bovine serum (Flow Laboratories, Austra-lia), 2mM L-glutamine and 18mM NaHC03. The cells were subcultured Figure 1. Apparatus for loading ofvertical stress once a week and seeded on the specimens after 4 days from the subculture.

Ima~e analyzin~ The images of cells through a polarizing microscope were input with a video camera for

low illumination intensity (min. 0.3 lux, Matsushita, Japan). The input images were A/D­convened into 512x512 pixels at 256 gray levels by an image processor (pIAS LA-500, Ja­pan). The A/D-convened images were processed into binary images by selecting an appropri­ate threshold value so that black areas would represent cells and white areas would represent material's surfaces. Then the values such as area, perimeter and maximum diameter of a cell were calculated from the binary images.

RESULTS AND DISCUSSION

Ex.periments Figure 2 shows changes of rates of peeled cells ( number of peeled ceUs after loading

vertical stress I number of peeled cells before loading vertica1 stress) concerning L-929 which had been seeded on alumina and cultured 6 hours before. In this experiment, vertical stress was sequentially applied to the cells on a specimen (3 min. by each strees). Figure 2 shows the

512

results of 2 kinds of sequences where 100 .. ~

each curve represents one sequence. (%) There were found differences of the CI)

rates of peeled cells after applying ~ 80 vertical stress between alumina and fi- U bronectin-coated alumina. i 60 ~

Adhesion areas of L-929 on the Qj 2 kinds of surfaces were measured by ~ processing numerically cell images '0 40 obtained by the CCD image camera. Q)

Figure 3 shows the distribution of ad- ~ 20 hesion areas of L-929 on alumina and fibronectin-coated alumina surfaces. Figure 4 shows means and standard deviations of adhesion areas and shape

o

factor 1 of L-929 on the 2 kinds of sur-

/.-r

o

.- ---- On Alumina

• On Fibronedin-ooalad Alumina

500 1000 (GI 1500 Loading acceleration

faces. The shape factor 1 was obtained . by calculating the ratio of (perimeter Flgure 2. Changes of rates of peeled cells of a cell)2 to 411: (adhesion area of acelI), so that the shape factor 1 would stand for the degree of spread morphology with pseudopods the adhered cells present Figure 3 and Figure 4 mean

~" I I I I

2" OnAlumina

-2v On Alumlna + Fibronectin

that the L-929 cells adhere to the fi­bronectin-coated alumina surface with larger adhesion area and more pseudo­pods in comparison with the alumina surface. Conceming the relation be­tween adhesion area and adhesion strength, adhesion area seems proba­bly to be related with adhesion strength according to the results in this study. In general, cells adhere to mate­rial surfaces with adhesion plaques where fibronectin receptors interact with fibronectin adherent to the sur­faces. Therefore, the results above

15

I" 'v

5

}.J o 2

M rvtA I .

'" " ~"V\. 600

Adhesion Area

~

12 ( .. ml 2) would be acceptable, if adhesion area ,. of a cell has relation to number of ad­

hesion plaques of the cello Figure 3. Distribution of adhesion areas on alumina

and fibronectin-coated alumina

Mathematical model The application of constant or cyclic loads of long duration on incubated cells on biomaterials results in a peel­ing that is a kind of failure of cell ad­hesion. The theoretical research on failure of cell adhesion is directed to­wards the establishment of criteria for extension of an isolated defect inside a closed cell adhesion area under the condition of quasistatic loading.

A model of defect growth kinet­ics at the interface between cells and a biomaterial is investigated in which the mechanism is presumed to be an

Adhesion Area 11 - Shape Fador 1

2000 ~========~=========;3 (11 m 2) 1 2.5

1500

500 r--------I--0.5

Ol...-_ ................ ..-.;.&....-_..L...-_ 0 Alumina Fibronecrtln coatad alumina

Figure 4. Means and standard deviations of adhesion areas and shape factorl

irreversible chemical reaction known in statistics as the pure birth processes.

Suppose that a growing defect at any instant is accumulating void at rate which can be wrinen analogically as a sequence of bimolecular reactions with rate constant: P . (Figure 5)

The distritution function of cracks between cells and the biomaterial f can be wrinen as Eq.(1), where f : crack distribution at the dass g c': crack initiation function P =13 l(g), l(g)=2m: perimeter of penny s~aped crack p: reaction rate of crack growth

dfoldt=c-Pofo

dfddt=ßofo-P1f1

513

y

90a = nr0 2

(go + g)a = nr2

Figure 5. Growing law of penny shaped crack

(1)

If the initiation of crack can be neglected (that is <:=0), then the solution of the equation (I) is given by Eq. (2), whereN: total number of cracks, <r>: average of crack radius, s: disper­sion of radius

(2)

On the other hand, if ~O, then the solution is known as Eq. (3).

Fg= f fg(t-'t~t, ~O (3)

The growth law of the penny shaped crack having radius runder the condition of con­stant load G can be written as Eq. (4) and the dispersion as Eq.(5), where a is the area of a ruptured bundle of fibronectin-receptor bond.

<r>=ro+aßt

02= a log r2-log~) 4Jt

G i i i i i

~penny haped crack

Ce~

(4)

(5)

Figure 6 shows a penny shaped crack in an adhesion cell under the distribution load G. The reaction rate of crack growth and its simplified form are given by Eq.(6).

Finally taking into account the stress intensity factor k l , the radius of the penny shaped crack is given by Eq.(7).

514

(6)

A: material constant, h: Planck's constant. K: Boltzmann's constant. T: absolute temperarure. I1f: free energy with the formation of a crack of radius r. I1F*: activation energy for the sepa­ration of crack surface. S: stress acting on the specimen. So: constant depending upon the shape of crack, v: effective volume. q: stress concentration. E: elastic modulus.

I r=rO+CI(KI>mKT 1, a--*>

where

Cl AaKTexJ I1fl~1dn: h It KT Sof1tEr··

(7)

As was already mentioned. the distribution of cells on a biomaterial takes a form of logarithmic normal distribution as Eq.(8) by the experimental observation.

f= d!cct exr{-(logr-D)2M2] (8)

In case of osteosarcoma (MG-63) D=1.32 ct=O.l1

FlOm the growth law of penny shaped crack Eq.(7) and Eq.(8). number of peeled ceHs from the biomaterials at time t is given by the last equation (Eq.(9».

n=[f(t)dt (9)

Figure 7 shows total number of detached ceHs on a biomaterial when the distributed load G is applied to the cells.

f

r

100 r--------::::: ..... ------,

(%) ..!!l 80

'8 "B 601------1----------1

140 1-----1---------1 'Ci S 20 ....... __ ._ ..... _ ........ _ ..... _ ..... __ ......... _ ......... _ ......................... .

o

<r>=ro+aßt r Figure 7. Total number of peeled ceHs

515

CONCLUSION

Tbe experimental results showed that the cells (L-929) adhered more tightly to fibronectin­coated alumina than to alumina, and at the same time the cells spread on fibronectin-coated alumina with larger adhesion area than on alumina. A mathematical model of cell peeling was considered based on the rupture process of fibronectin-receptor bonds. Tbe constitutive as­sumptions Eq.(I) to Eq.(9) are of course highly idealized. Prediction of peeling behavior of cell based on these assumptions are, however, qualitatively characteristic of real cell adhesion on a biomaterial. Tbe model of peeling mechanism would be useful for evaluation of initial bio­compatibility of biomaterials implanted in biological hard tissue.

REFERENCES

1. Tateishi, T., Continum Tbeory of Cumulative Damage, Bulletin of the JSME, 1976, 19, 1007-18

2. Ushida, T., Tateishi, T., Adhesion strength and adhesion area of cultured cells on bioceram­ics ,Advances in Biomaterials, 1990,9,231-236

3. Ushida, T., Tateishi, T., Evaluation of Cell Morphology on Bioceramics, Bioceramics, 1989, 1,36-41

3. Pierschbacher, M. D., Rouslahti, E., Cell attachment activity of fibronectin can be dupli­cated by small synthetic fragments ofthe molecule,~ 1984,309,30-33

4. Rouslahti, E., Pierschbacher, M. D.,~, 1987,238,491-497

Adar, F., 317 Anderson, 0., 402 Antolotti, N., 195,236 Arm, P., 461 Aurelle, J.-L., 230

Baldet, P., 383 Baquey, A., 345 Baquey, C., 345 Barbezat, G., 156 Barinov, S.M., 206 Baschenko, Yu.V., 206 Baudendistel, E., 504 Bertocchi, G., 265 Bertoluzza, A., 189 Biasini, V., 236 Biasiol, S., 353 Böhler, M., 17 Bolz, A., 326, 360 Bonnel, F., 383 Borghetti, P., 195 Bousfield, B., 217 Bousquet, G., 230 Bueno Lozano, A., 26 Bukat, A., 223, 378 Bungaro, P., 73

Caja, V., 141 Calista, F., 26 Callejas, P., 244 CandeIon, B., 345 Cannas, M., 353 Capannesi, G., 211 Caroli, A., 49 Carrerot, H., 230 Cenni, E., 285 Chao, E.Y., 141 Chen, A., 78, 89 Chen, J., 78, 89

517

INDEX OF CONTRIBUTORS

Chuong, W., 408 Ciapetti, G., 285 Clarke-Smith, E.M.H., 334 Coppola, G., 295

Dalla Pria, P., 422 DeI Bo, M., 101 Denissen, H.W., 130 Dion, 1., 345 Dondi, M., 308 Dörre, E. 454 Drenckhahn,471 Dubok, V.A., 438

Egger, E., 141 Evans, E.J., 334

Farina, C., 62 Fassina, P., 223, 378 Fattori, G., 308 Filmer, H., 156 Fioravanti, S., 236 Freche, M., 181

Gabbi, C., 195 Galli, G., 432 Garbassi, F., 444 Gasser, B., 491 Gatti, A.M., 402 Gatti, M.A., 308 Ghuong, W., 408 Giannini, S., 62, 295, 388 Giunti, A., 26 Gottsauner Wolf, F., 141 Greco, F., 211, 223 Griss, P., 302, 366, 372, 472 Groot, C.G., 340 Grootde,K., 166,340

Gross, U., 275 Gualdrini, G., 73

Hirota, K., 417 Hochstrasser, J., 156 Holmes, R., 317 Hooff van den, A., 130 Huguet, M., 383

Jammet, P., 383 Jianguo, Z., 408

Kalk, W., 130 Klein, C.P.A.T., 166 Knahr, K., 17 Kraft, M., 366 Krajewski, A., 62, 203, 236,

295,388,497

Laudadio, P., 486 Locardi, B., 148 Ludwig, M., 372

Maccauro, G., 378 Macedo, S., 302 Maggiore, E., 477 Mangano, C., 236, 497 Marinoni, E.C., 67 Martinetti, A., 236 Martinetti, R., 62, 203, 486, 497 Masse, A., 353 Mathys Jr., R., 491 Mesana, T., 345 Monari, E., 308 Mon~oqp,R.,256,26O,265,270 Montanari, G., 113 Monticelli, G., 35 Monties, J.-R., 345 Montina, P .P., 26 Morelli, M.A., 189 Moreschini, 0., 35 Moroni, A., 62, 73, 124, 141, 295, 388, 432 Müller, W., 491 Müller-Mai, C., 275

Nehls, V., 471 Nicoll, A.R., 156 Nieuport de, H.M., 130

518

Noem, G., 308 Nucci, C., 256, 260, 270,

Occhiello, E., 444 Orth, J., 302,366,372,471

Pasquino, E., 308 Passi, P., 107 Pavone, S., 124 Piancastelli, A., 203, 236 Piantelli, S., 223, 378 Piconi, C., 211, 223 Piening, W., 372 Pigato, M., 295 Pin, Z., 408 Pitten. S., 195 Pizzoferrato, A., 26, 285 Plenk Jr, H., 17,317 PonzUurl,L.,I24,295 Pourtein, M, 345 Prati, C., 113,256,260,265,270 Presutti, L., 486

Rambert, A., 230 Ravagli, E., 477 Ravaglioli, A., 1,62,67,203,236,

295, 388, 486, 497 Rey, c., 181 Ricon, J.Ma., 244 Rieu, J., 230 Rinaldi, S., 308 Rollo, G., 73, 124, 141,432 Romanini, L., 35 Royer, P., 181 Ruggeri, A., 353

Sabato, C., 73, 432 Salito, A., 156 Salzer, M., 17 Santi, M., 461 Savarino, S., 285 Savino, A., 113,256 Scbacken, H.G., 130 Schäfer, R., 504 Sch~h,M.,326, 360 Schön, R., 118 Sedda, A.F., 211 Semlitsch, M., 118 Soltesz, U., 504

Souyrios, F., 383 Specchia, L., 73, 124, 432 Squarzoni, S., 26 Stea, S., 26, 285 Streicher, R.M., 118 Strocchi, R., 353 Sturlese, S., 236 Sudanese, A., 26 Suhih, L.L., 438 Szarska. S., 396

Tateishi, T., 510 Tinti, A., 189 Toni, A., 26 Toschi, E., 113,256,260,265,270 Trinchese, L., 141 Trotta, F., 156,236

Ushida, T., 510

Valdre, G., 270 Vallana, F., 308, 461 Vendemia, V., 124

519

Venini, G., 67 Venturini, A., 295 Voigt, C., 275 Vrouwenvelder, W.C.A., 340

Walter, A., 17 Wan, D., 78, 89 Wang, S., 78, 89 Weinländer, M., 317 Weiqun, C., 408 Wilke, A., 302, 366, 372, 471 Willmann, G., 250 Wolke, lG.C., 166

Xingdong, Z., 408

Zaffe, D., 295, 388, 402 Zaghini, N., 223 Zaghis, A., 101 Zanasi, S., 49 Zarotti, F., 46 Zhang, X., 78, 89 Zhou,l, 78,89 Zinghi, G.F., 73, 432