J O U R N A L O F M A T E R I A L S S C I E N C E : M A T E R I A L S I N M E D I C I N E 1 6 (2 0 0 5 ) 1165 1171

Titaniumhydroxyapatite porous structuresfor endosseous applications

C. POPA1, V. SIMON2, I . VIDA-SIMITI1, G. BATIN1, V. CANDEA1, S. SIMON21Technical University of Cluj-Napoca, 103-105, Bd. Muncii, 400641 Cluj-Napoca, RomaniaE-mail: catalin.popa@stm.utcluj.ro.2BabesBolyai University of Cluj-Napoca, Romania

Materials for uncemented endosseous implants have to assure an as short as possibleosseointegration time. Thus, a material with both surface bioactivity and a porous outerstructure can become a preferred choice for this type of applications. This paper presents aclass of titanium-base PM composites, reinforced with particulate hydroxyapatite. Rawmaterials were titanium powder, obtained through hydridingmillingdehydriding, withthe grain size of 63100 m, and sol-gel hydroxyapatite (HA) powder, produced by thereaction between Ca(NO3)24H2O and (NH4)2HPO4. Blends with 5 to 50% HA were preparedand pressed in a rigid die, producing single composition or gradual composition samples.The applied pressure was of 400, 500 or 600 MPa. Sintering was performed in vacuum, at1160 C. All samples, although well sintered, displayed swelling during sintering, due todiffusion into the matrix. The increase in volume is more severe for higher amounts of HAin the green compacts and for higher applied compaction pressure. Compacts with agradual increase of the HA content are recommended from the functional and mechanicalpoint of view, but the increase should be slow, not to produce interlayer cracks. The outersurface shows interconnected pores, suitable for the ingrowth of vital new bone.C 2005 Springer Science + Business Media, Inc.

1. IntroductionThe time that passes from the surgical interventiontill the stable fixation of uncemented endosseous im-plants is of major importance for the success of theoperation. An unexpected mobilisation leads to theformation of a thick fibrous capsule around the im-plant, blocking the blood supply to the new boneand causing an early failure. This leads to the ideaof maintaining the classic method of cementing im-plants. Although solving the early mobility problem,this is far from ideal: the method is suitable mostly onlyfor orthopaedic implants; the intimate contact implantmaterialcementbone can produce electrochemicalreactions decreasing the biocompatibility; the generallyused PMMA-base cement can be the place for interfa-cial cracks [1]. Biological fixation, using the growthof newly formed bone into the non-uniformities ofthe implants surface is more suitable from the longterm biomedical point of view. A PM porous surfaceis favouring osseointegration. Titanium is the preferredchoice for the quasi-bioinert material to produce im-plants with biological fixation, as it has proven a bet-ter biocompatibility in comparison to the other metalsand alloys used in implantology [2, 3]. This is due tothe stable oxide layer that insulates titanium surfacesfrom the tissue and is a good basement for the cas-cade cellular processes that precede the formation ofbone, causing almost no inflammatory response after

implantation [4]. Titanium was found the only metalbiomaterial to osseointegrate [5, 6], as resulted from a20 years dental surgery experience. In time, there wereeven assumptions on a possible bioactive behaviour [7],due to the slow growth of hydrated titanium oxide onthe surface of the titanium implant that leads to the in-corporation of calcium and phosphorus. The in timesuccess rate of surgery related to titanium implants isstill largely dependent on the size/shape of the implant,as well as on the general clinical conditions, and, some-times, far from 100% [811]. As osseointegration startswith the cellular stage and continues with the nucle-ation of mineral and the structuring of the new vitalbone, the overall required time is varying in a broadrange. A proposed solution for a better control of os-seointegration is the bioactive fixation. The most usedmaterial is hydroxyapatite, as the basic mineral in ma-ture bone. Coating of metal implants with hydroxya-patite became a usual method, although it has somedrawbacks: the toughness of hydroxyapatite is low (KIC 1 MPa m1/2); during the coating stage (plasma sput-tering, etc.), interfacial cracks can occur; the adher-ence metal/ceramic is low and can diminish in time.Hydroxyapatite was also found to disturb the prolifer-ation and osteocalcin synthesis of human osteoblasts[12].

We propose, in this paper, porous titaniumhydroxyapatite composites as materials for uncemented

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endosseous implants, to summate the biologicaladvantages of both components.

2. Materials and methodsTitanium powder (0.01%Fe; 0.01%Al; 0.001%Si;0.05%Mg) with the grain size of 63100 m was used.The powder was obtained through the hydridingmillingdehydriding process and the grains display aspecific shape. The purity degree was assessed using the transformation temperature, by DTA (DTA/DSCSetaram Labsys DTA16). Alumina crucibles of 100 lwere employed for 100 mg samples. The temperaturewas calibrated according to the Curie temperature ofnickel.

Hydroxyapatite (HA) was obtained through a sol-geltechnique. The used reagents were Ca(NO3)24H2Oand (NH4)2HPO4. Ca(NO3)24H2O was dissolved inwater at 90 C, while (NH4)2HPO4 was dissolved atthe room temperature. The filtrate that resulted aftermixing was washed with water at 40 C, than dried inan oven for 2 h, at 110 C. The gel was heat treated, firstat 900 C, for 3 h, then left to cool slowly, in the oven.A second heat treatment was performed at 1150 C, forone hour, with oven cooling to 650 C, then air-cooling.The obtained HA was milled and screened, in order touse a powder with the grain size of less than 100 m.

Blends of titanium and 5, 10, 15, . . ., 50% HA weremixed in a turbula mixer for 30 min. Samples werepressed in a rigid die with the surface of 0.5 cm2, with-out the use of any lubricant, with pressures of 400, 500and 600 MPa, 3 samples for each experimental condi-tion. Single composition samples, as well as multilay-ered ones were produced. In these cases, the stackingsequences were conceived for a continuous increase ofthe hydroxyapatite content. The green compacts weremeasured (0.01 mm precision micrometer) positionedseparately in zirconia crucibles and subsequently vac-uum sintered (106 torr) at 1160 C for 60 min, withdwelling stages at 200, 600 and 800 C, 30 min each.A part of the samples were cross sectioned by the mid-dle of the generator, ground (180, 240, 320, 400, 600,1000, 2000), polished with alumina (

TABL E I Technologic properties of the employed Ti powder

Size fraction Flow rate Apparent density Compression Fill porosity Tap density Tap compression Tap porosity[m] Vflow [s/50 g] a [kg/m3 103] Ratio Ca [%] Pa [%] v [kg/m3 103] Ratio Cv [%] Pv [%]

63100 82.7 1.93 42.79 57.21 2.35 52.10 47.906380 87.2 1.88 41.68 58.32 2.16 47.89 52.1180100 66.5 1.76 39.02 60.98 1.92 42.57 57.43

Figure 3 General aspect of raw HA powder.

Figure 4 DTA pattern of the Ti powder (calibration by Curie temperatureof Ni).

Figure 5 XRD pattern of raw HA powder (peaks at 2 angles of 25.62,31.13 and 34).

This means that the filling of dies has to be performedthrough a larger nozzle than usual. The compressionratio is better for the finer fraction, compared to thecoarser one. The best density/compression ratio wereobtained for the whole powder. If associated to flowrate, the whole powder was chosen as raw material forthe following experiments. Although vibration inducesa more advanced compaction, due to the problems relat-ing to the filling of dies, we have chosen not to employthe method.

The properties of titanium are strongly dependenton the impurities content (gases are to be considered,mainly). The purity assessment of the raw Ti powderwas performed by the transformation temper-ature (882 C, according to ASM), strongly dependentmainly upon the content of stabiliser impurities. Fig. 4shows the exothermic curve, with the transition temper-ature at the raw value of 907.5 C.

The measured values of the temperature were higherwith 24.34 C than the real ones, due to the thermalinertia, as determined after nickels Curie point. Thus,

Figure 6 Volume variation of 5% HA samples after sintering.

Figure 7 Sintered Ti 5% HA sample (600 MPa compaction pressure),polished, etched (Krolls reagent).

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Figure 8 Surface of sintered Ti 50% HA sample, unpolished.

the calculated transition temperature was of883.61 C. The small difference to the standard 882 Ctemperature (1.61 C) reveals the purity of the Ti rawmaterial.

The diffraction patterns performed on the employedHA powder, Fig. 5, are close to the data in litera-ture, stating that the principal diffraction peaks of HAappear at 2 values of 24.2 for reflection (002), at

Figure 9 EDX elements distribution maps, 5% HA sintered sample, formerly pressed with 400 MPa.

30.8 (triplet) for reflections (211), (112) and (300),and at 34.0 for reflection (200) [13, 14]. The peaksof the HA powder employed in this study were at 2angles of 25.62, 31.13 and 34.

All samples displayed swelling during the sinteringstage. Samples with 5% HA that were compacted usingone of the three values of the applied pressure and sub-sequently sintered in the same batch were measured.Their increase in volume is shown in Fig. 6.

A higher pressure was shown to produce moreswelling. While for 400 and 500 MPa the dispersionof values is broader, the 600 MPa pressure generated amuch more even swelling among samples. The causeof swelling is the diffusion of calcium/phosphorus inthe titanium matrix. The phenomenon was proven bysubsequent EDX analysis. The effect of pressure overswelling shows that, by a higher loading of the grainsin the metal matrix during the compaction stage, thediffusion coefficient of Ca/P in Ti increases. This iscaused by the more numerous lattice defects inducedby a more severe plastic deformation. The dispersion ofthe measured values of diameter/height is broader forsmaller pressures. This is consistent with the fact thatthe pressure is transmitted unevenly through the col-umn of powder inside the die (no added lubricant). Ahigher pressure (600 MPa in this particular case) leads

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to a more even pressure transmission in the entire greencompact.

Fig. 7 presents an example of structure for the 5%HA sintered compacts.

Samples are well sintered, with a compactnessincreasing with the compaction pressure. During thegrinding/polishing of samples, the ceramic compo-nent particles in the outer layer were pulled out andleft craters in the surface of samples. This is wellobserved in the EDX elements map around these re-gions. If analysing the surface of samples, without pass-ing through the cuttinggrinding procedure, the posi-tion of the ceramic phase inside the compact can beseen, Fig. 8 (50% HA). The transition layer that can beseen at the interface between the matrix and the ceramicphase is believed to play a positive role in the overallmechanical stability of the compact.

The pores, with their bio-functional role, areimportant in what concerns osseointegration. Thecharacteristic to be considered for a material usedin endosseous implants is not the overall porosity(percentage of pores), but the aspect of pores.

The size of pores in the composite samples wasestimated microscopically. As seen in Figs. 7 and

Figure 10 EDX elements distribution maps, 5% HA sintered sample, formerly pressed with 600 MPa.

8, the pores are interconnected, with very irregularshapes/sizes. Large pores, with the main dimensioncomparable to 100 m (the critical minimum sizeneeded for the formation of a vital new bone), as well assmaller pores, could be seen. A higher compaction pres-sure applied to green samples reduced the sinteredstate porosity.

The EDX analysis on polished samples has shownthe diffusion of elements from the HA particles tothe titanium matrix. Fig. 9 shows the distribution mapof main chemical elements around a crater resultingfrom pulling out a ceramic particle (5% HA sinteredsample, formerly pressed with 400 MPa). The diffu-sion of Ca is much more even inside the titaniumgrains, compared to P , that remained much less dif-fused, in the proximal region around the HA parti-cle. Fig. 10 presents the distribution of Ca and Paround a removed ceramic particle, on a sample thatwas formerly pressed with 600 MPa. Ca atoms remainwell diffused inside the matrix. P was shown to dif-fuse in a more advanced manner compared to Fig. 9.Fig. 11 shows, comparatively, the relative amountsof the main elements in the areas shown in Figs. 9and 10.

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Figure 11 Comparison between quantitative results of EDX on400 MPa. (a) vs. 600 MPa. (b) pressed samples (5% HA).

Figure 12 XRD pattern, Ti HA composite (Ti peaks at 2 angles:38.25, 40.04 and 52.67C).

The amount of diffused P in the Ti matrix is moreimportant as the compaction pressure of samples washigher (600 vs. 400 MPa). This is consistent with theabove conclusion that a higher pressure leads to a moreadvanced diffusion of elements from HA to the Ti ma-trix.

XRD analysis performed on sintered samples,Fig. 12, shows both components, Ti and HA in the finalstructure of the composite. The peaks that occur for 2angles of 38.25, 40.04 and 52.67 can be attributed toTi [15], while the peaks for HA remain virtually thesame as in Fig. 5.

The composition of green samples is the other keyfactor, together with compaction pressure. For a higheramount of HA, the diffusion process of Ca and P oc-curs in more numerous sites (around HA particles).The swelling effect is, then, more obvious, Fig. 13(400 MPa, 30% HA).

For the gradual samples (with a later-to-layerincrease in the HA amount), the swelling effect causesthe limitation of the method. In the same sample (ba-

Figure 13 Effect of the HA content on the samples increase in volume(400 MPa compaction).

sically the same transferred pressure), that consists ofseveral layers with various amounts of HA, the varia-tions of volume are different (higher for the layers withmore HA), causing stresses between layers. The inter-layer stress is higher when the difference between thecompositions of layers is higher. Transverse cracks canoccur for differences of more than 5%.

The gradual amount of HA is of particularimportance for implants. The more HA at the directboneimplant contact surface, the better the expectedbiocompatibility. On the other hand, the presence of HAhas only a functional effect, as it decreases the mechan-ical properties. The more the HA content, the higher thebrittleness. Thus, a structure with an as much as possibleHA content in the areas in direct contact with bone andan as low as possible in the rest would be the most suit-able. Unfortunately, it is not possible, as shown above,to decrease sharply the HA content, due to the risk ofcracking. The samples with more than 30% HA displayan excessive brittleness and, at 50% HA, the materialis virtually unusable.

For every shape/size of implant, the couplepressurestacking sequences can be optimised.

4. ConclusionsPorous PM Ti HA composites (5 to 50% HA), usinga Ti powder obtained by the hydriding-millingdehydriding process and sol-gel HA, were obtainedthrough a blendingpressing-sintering route.

A diffusion process of Ca and P in the matrix duringsintering (1160 C) causes the swelling of samples.The process is controlled by two parameters: com-paction pressure and HA content. A higher pressure(over 500 MPa), as well as a higher amount of HA,lead to significant swelling.

The amount of HA is of major importance and hasto be kept into severe limits: a too high amount (morethan 30%) leads to excessive brittleness; a too low con-tent in contact with bone is insufficient for the requiredbioactivity.

Gradual structures, with a high HA content at thesurface and a low one in the interior, are the most suit-able. The decrease of the HA content one layer to an-other has to be slow (highest rate of 5%), not to producecracks as a result of the different swelling rates.

The outer pores are interconnected and suitable forthe ingrowth of a new vital bone.

AcknowledgmentsThe authors wish to thank the Romanian Programfor Materials, Micro and Nanotechnologies,MATNANTECH, for the generous financial sup-port.

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Received 13 Juneand accepted 19 August 2005

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