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ChapTEr II.5.8 Adhesives and Sealants 889

CHAPTER II.5.8 ADHESIVES AND SEALANTS

David Christopher WattsProfessor of Biomaterials Science, The University of Manchester, School of Dentistry & Photon Science Institute, Manchester, UK

INTRODUCTION

Description and Definition of adhesives, and related Terminology

In the context of biomaterials, adhesion science and tech-nology assume high importance. It is useful to commence with the slightly wider concept of bonding. It is evident that many biomaterial devices are structured from mul-tiple components, and these need to retain their mutual integrity during clinical service. In this context, the bonds in question may include metallic welds, although these would not be considered within the domain of adhesion, any more than screw attachments that have a role in some implants.

The most challenging issue is the creation of a durable interfacial bond between a biomaterial and its host tis-sue. In addition to the function of bonding, per se, an adhesive biomaterial may be required to fulfil a space-filling role – replacing some or all of any lost natural tissue. In the former case, the function may be termed grouting. This is, for example, the primary function of bone cements in orthopedics. A further ideal function of an adhesive is that of sealing; that is, the prevention of ingress of moisture, air, biological fluids, bacteria or other species through the adhesively bonded zone.

Adhesive is a general term, and in specific contexts may be replaced by designations such as cement, glue, paste, fixative, and bonding agent. Some adhesives may be designed to exhibit further functions, such as anti-bacterial action, delivery of drugs or beneficial ions, such as the antibacterial ion Ag+ or fluoride (F−). Fluoride is a component of dental preventative treatments which ini-tiates the partial replacement of hydroxyapatite (HA) – the tooth enamel’s normal crystalline composition – with fluorapatite (a related crystal which incorporates F−). Fluorapatite is more resistant to decay than HA.

Sometimes the converse of adhesion, namely abhe-sion, is necessary in clinical treatment. This “non-stick” feature is required with blood-compatible biomaterials, and is exhibited by non-fouling surfaces covered with polyethylene glycol (PEG).

During the past half-century there has been a strong convergence of the ancient technologies of adhesives and the modern science of adhesion – based especially on advances in surface science and the understanding of molecular adhesive mechanisms. This progress has been stimulated by widespread industrial uses and biomedical applications of adhesives. The former have been driven by advances in polymer science, and the adoption of lightweight alloy and non-metallic composite materials

in the aerospace and automobile industries. Neverthe-less, the stringent biomedical requirements for adhesives have led to some specialized developments.

The diverse biomedical contexts (host environments) of adhesives each require appropriate materials and tech-niques. Biological hosts differ principally as to whether they consist of hard or soft tissues. Connected with this are the timescales required for adhesive/bonding dura-bility. These may vary from days, in the case of wound- closure adhesives, to decades of years, in the cases of bone cements and dental restorations. The aim of this chapter is to give a sound introduction to the physico-chemical and materials science aspects of adhesives and sealants, together with a description of current systems utilized in a range of surgical practice. It does not attempt a detailed study of the molecular biological interactions of adhesives.

More details on the background to adhesion and adhe-sives can be found in a number of texts (Kinloch, 1987; Comyn, 1997; Pocius, 1997; Mittal and Pizzi, 1999; Eliades et al., 2005; Matinlinna and Mittal, 2009). Gen-eral information on wound closure and surgical adhesives is given in Sierra and Salz (1996) and Chu et al. (1997). The proceedings of a conference on adhesion in den-tistry provide an important set of contemporary reviews ( Armstrong et al., 2010; Braga et al., 2010; Mair and Padipatvuthikul, 2010; Marshall et al., 2010; Perdigão, 2010; Scherrer et al., 2010; Söderholm, 2010; Tagami et al., 2010; Van Meerbeek et al., 2010).

THE LOGIC OF ADHESION PROCEDURES

Inspection of a wide range of industrial, domestic, and biomedical adhesives shows the generally common pre-sentation as some form of fluid agent that incorporates a means of conversion into a solid form. The vast majority of such adhesives are designed for application to adher-ends that are fairly rigid solid surfaces. However, there is often an educational gap in the understanding of why fluid adhesives are necessary for application to solid adherends, and why they must undergo solidification reactions. Hence, we shall outline the straightforward logic of such an adhesive design strategy.

Our starting point – in chemistry – is the existence of electromagnetic forces between, as well as within, molecules. Additional to the primary bond types (cova-lent, ionic, and metallic), these secondary intermolecular forces include hydrogen bonds and van der Waals bonds, both dipolar and dispersion types. There is a wide varia-tion in the relative strength of these bonds, covalent bonds being the strongest; even weak bonds can have a massive cumulative effect if they have a high number density, as with the H bonds between water molecules.

However, it is characteristic of intermolecular forces and interaction energies that they diminish rapidly with separation, often depending on the inverse 6th power of the separation distance (r). For example, the net interac-tion energy (U) between two permanent dipole moments

890 SECTION II.5 Applications of Biomaterials

(μA, μB) within neighboring molecules or surfaces is determined by the overall average of the product:

[U(r) × e−U(r) / kT

]

where k is Boltzmann’s constant and T is temperature. When:

A B KT3r

then:

A BUaverager KT

2 2

6

2 1

3

Where ε is the permittivity of the medium.Hence, if two solid surfaces are brought into suffi-

ciently close proximity, at moderate temperatures, there is the possibility of interaction energy between them.

Nevertheless, there is a further topological factor; namely that even macroscopically smooth surfaces are microscopically rough. Hence, two solids in close proxim-ity at the nanoscale may be akin to a superimposed pair of mountain ranges. The actual contact regions between the two solids are restricted to the “mountain peaks,” and a major proportion of each total solid area is not in contact and is held apart beyond the distance zone over which the interaction energies are significant. The effective intimate contact area between the two surfaces is then too small to directly realize the potential of the intermolecular forces. It follows that rigid contact between the two solids fails to deliver an appreciable net interaction force. However, there are some exceptions to this outcome when one or more of the contacting solids exhibits plastic flow under mechanical stress. Thus, pieces of gold foil may become self-adherent under compressive stress. Similarly, waxes can mutually bond when temperature is raised slightly.

To achieve adhesive bonding in the general case of two rigid solids, such as a tooth enamel surface and an orth-odontic bracket, it is necessary to apply a fluid adhesive between them. Moreover, the fluid must be of appropri-ate chemical formulation to initially wet both surfaces, exhibiting a low contact angle. One or both surfaces may have been subjected to some form of pre-treatment or conditioning with an etchant or primer that, inter alia, may have modified surface porosity. In this case, the adhesive fluid may be drawn into the solid surface layers by capillary action.

The presence of a suitable fluid between two solids greatly enhances the potential for intermolecular force interactions at each solid–fluid boundary. Further, if for example, the solid is a calcified tissue and the fluid incor-porates carboxylic acid groups, then primary bonds may be created at the interface. Such primary and/or secondary bonds can immediately generate a measureable resistance to tensile loading of the solid interface. This may be aug-mented by the Laplace pressure difference arising from the fluid surface tension. Nevertheless, such solid pairs weakly bonded by a fluid can be readily disrupted by shear forces.

In the general case, therefore, to achieve an adhesive bond zone stable against shear forces, it is imperative to

convert the adhesive agent from a fluid to a solid. In this process, it is desirable for the adhesive agent to exhibit high dimensional stability in relation to the interface with one or more adherends. There are several physico-chemical mechanisms available for achieving adhesive solidification, or setting, including:

• Phase transformation on cooling; • Solvent evaporation; • Polymerization (and cross-linking) of fluid monomers

or oligomers; • Acid–base reaction.

Phase transformation has not been applied with biomed-ical adhesives, as the requisite temperatures (normally above 100°C) are beyond the biologically-tolerable range. Polymerization of monomer systems is the most widely used approach. However this is susceptible to polymerization shrinkage phenomena and thus stress-development (see the section on Stress-Development in Adhesive Joints due to Polymerization Shrinkage). Acid–base reactions are present in the setting of polycarboxyl-ate and glass–ionomer cements.

The resultant interfacial bonds, generated by the above adhesion strategy, can be analyzed into different micro- or nanostructural contributions. These may include micro-mechanical bonds created by the interlocking of solidi-fied adhesive “tags” (extensions) into surface porosity or roughness of the adherend. Where the microadaptation of the solid adhesive and the adherend remains molecularly intimate, the intermolecular forces can contribute; and where specific primary bonds are generated, especially with metallic adherends, their contribution is significant.

Given the foregoing general scheme of how adhe-sive bonds can be generated, many of the factors that can weaken or compromise these bonds will be obvi-ous. These include: air voids; contaminants; and weak boundary layers. This leads to consideration of surface pretreatments designed to avoid these problems.

ADHEREND SURFACE PRETREATMENTS TO ENHANCE BOND STRENGTH AND DURABILITY

Biomedical adhesive systems for use as, or with, bioma-terials are usually accompanied by detailed instructions for use. Frequently this printed information and advice is supplemented by audiovisual material relating to the mechanisms of action and to detailed applicatory steps. The steps involve meticulous preparation of the adher-end surfaces, especially in the case of hard tissues. In practice, there can be a tension between the complexity and effectiveness of the pretreatment versus the desire for clinical simplicity and overall speed of application.

Preparation may involve cleansing the tissues from contaminants such as blood or saliva, for if these sub-stances remain in situ they can have a negative, abhesive effect. However, the goal is often not merely to remove a contaminant, but also to enhance surface free energy and


thereby promote improved wetting, spreading, and pen-etration of the adhesive agents. In the case of hard dental tissues, a well-established treatment is application of an aqueous phosphoric acid solution to etch and/or deminer-alize the outer tissue layers (as considered in the sections on Hard Tissue Adhesives: Bone and Tooth Cements; and Acid-Etch Bonding to Enamel). A converse experimental approach in former times was to hyper-mineralize cal-cified tissues. This was successful in vitro, but required clinically unfeasible pretreatment times.

When bonding to synthetic biomaterials, such as ceram-ics, roughening of a “fitting” surface – to be bonded – increases the effective surface area. This is often combined with application of chemical primers, such as silane cou-pling agents. In the case of many implants and oral pros-theses, the bonding surface may undergo plasma spraying to deposit a well-attached ceramic layer, perhaps of hydroxyapatite (HA). This radically changes the chemi-cal nature of the surface to which adhesive bonding must then be achieved.

HARD TISSUE ADHESIVES: BONE AND TOOTH CEMENTS

auto-polymerizing pMMa Bone Cement

historical Background. In 1936 it was noted that mix-ing ground polymethylmethacrylate (PMMA) powder with the monomer (MMA) produced a dough which could be manipulated and molded; hence it became one of the first biomaterials (Weber and Chapman, 1984; Donkerwolcke et al., 1998; Kühn, 2000; Webb and Spencer, 2007). Early applications were in dentistry. This was initially employed by orthopedic surgeons as a cement or grout to improve implant fixation in 1953. However, the major breakthrough in the use of PMMA/MMA bone cement in total hip replacement (THR) was proposed by Smith and Charnley, who used it to secure fixation of the acetabular and femoral components and to transfer loads to bone ( Charnley, 1960, 1970) (Figure II.5.8.1). More recent applications include percutaneous vertebroplasty applied to the osteoporotic spine, which can provide significant and prolonged relief of pain (Lewis, 2006; Webb and Spencer, 2007).Mechanism of Setting of pMMa/MMa Dough. Methyl methacrylate (MMA) liquid (Figure II.5.8.2) is rapidly absorbed into PMMA powder, creating a tacky material that progresses to a non-tacky dough stage. These can be considered merely physical changes. However, benzoyl peroxide (BPO) initiator, mixed in with the powder, can be chemically activated to form free radicals, either by heat-ing to above 60°C or by the incorporation of a suitable amine activator into the monomer, to enable room tem-perature auto-polymerization or “cold-cure.” As activa-tion generates free radicals from BPO, these start to react with the MMA molecules. The free radical polymerization reaction thus continues through the propagation stage – where chain growth occurs from multiple radical initia-tion sites – until this reaches the termination stage, where

polymerization is either fully complete or the free radicals are consumed. MMA polymerization is a highly exother-mic process, arising from the conversion of C=C bonds. Hence, in the constrained interfemoral environment there is normally a significant temperature rise, which auto-accelerates the setting process, peaking above 56°C for 2–3 minutes in vivo. Sometimes higher temperatures are reached, up to 90°C, which can volatilize unreacted mono-mer (boiling point of MMA is 101°C). This creates voids that can later lead to mechanical failure.rheological Factors in Cement Delivery. While the physical absorption and chemical polymerization pro-cesses take place, the material at the early dough stage is forcibly injected into the femoral space around the metallic femoral prosthesis. The rheology of the injection process is complex. Bone cements can be formulated to yield dif-ferent viscosity ranges: low; medium; and high. High vis-cosity cements have a short waiting/sticky phase, followed by a long working phase that is more suitable for many surgeons.

FIGURE II.5.8.1 Cross-section of femur and interlocking bone cement showing good adaptation of the cement to the prosthesis (now removed), and a zone of interaction with the inner surface of the compact bone. (with courtesy of J. Charnley, personal com-munication)

FIGURE II.5.8.2 Methyl methacrylate and methyl cyanoacrylate.


Cement in its liquid phase of curing behaves as a non-Newtonian (pseudoplastic) fluid, with viscosity decreas-ing as shear rate is increased. However, the viscosity of all cements increases appreciably during polymerization as the polymer chains lengthen until the material is solid-ified. Setting times range from 5 to 13 minutes from start of mixing (Kühn, 2000).Mechanism of “Bonding” or Grouting. As the cement flows under pressure to penetrate the interstices of can-cellous bone and adapts to the surface of the femoral stem, it can achieve micro-interlock with the bone when solidified. Nevertheless, PMMA bone cement undergoes a degree of shrinkage on polymerization. Hence another benefit of pressurized injection is partial compensation for this effect. Shrinkage is a potential major source of porosity (Gilbert et al., 2000).

alternative Bone-Cements: Calcium phosphate. In orthopedic surgery there are some alternatives to PMMA bone cement, and there is also the option of cementless fixation. Inorganic cements, notably from the Ca–P sys-tem, include HA [Ca10(PO4)6OH2] and TCP [Ca3(PO4)2]. The biocompatibility of HA and its similarities to bone mineral have led to the study of dense HA for the aug-mentation of osseous defects.Classical and Modern Dental-Bonding Cements: Conventional acid–Base Cements. Dental cements are, traditionally, fast-setting pastes obtained by mix-ing solid and liquid components. Most of these materials set by an acid–base reaction, and subsequently devel-oped resin cements harden by polymerization (Smith, 1971, 1991, 1998). The composition of these mate-rials is shown in Table II.5.8.1. The classes and typical mechanical properties are given in Table II.5.8.2. The

Influence of pore Distribution in Bone Cements on the Cement/Implant Interface

Implant fixation has been achieved with acrylic bone cement for many years. Although this is a successful application overall, the bond between the implant and bone cement has been shown to be the most fragile link in the construct in femoral components, with failure likely only a short time after implantation. Interfacial porosity and the formation of microcracks have been identified as factors detrimental to cement/implant performance, with pores serving as nucleation sites for these microcracks.

To decrease interfacial porosity and improve the longevity of the cement/implant bond, several authors have suggested pre-heating implants. Although cements react differently to the curing

TaBLE I I .5.8.1 Classification and Composition of Tissue Adhesives

Type Components Setting Mechanism

hard Tissue adhesives

BoneAcrylic bone cement Methyl methacrylate and polymethylmethacrylate Peroxide – amine initiated polymerization

TeethDental cements:

Zinc phosphateZinc oxide powder, phosphoric acid liquid Acid–base reactions;

Zn complexationZinc polycarboxylate Zinc oxide powder, aqueous poly(acrylic acid) Acid–base reactions;

Zn complexationGlass ionomer

(polyalkenoate)Ca, Sr, Al silicate glass powder aqueous poly(acrylic

acid-itaconic acid)Acid–base reactions;Metal ion complexation

Resin modified glass ionomer

Dimethacrylate monomers. Aqueous poly(acrylic acid- methacrylate) co-monomers. Silicate or other glass fillers

Peroxide-amine or photo-initiated polymerization

Resin-based Aromatic or urethane dimethacrylates, HEMA Photoinitiated addition polymerizationDentin adhesive Etchant: Phosphoric acid (aq.)

Primer: HEMA in ethanol or acetoneBond resin: Dimethacrylate monomers

Photoinitiated addition polymerization

Soft Tissue adhesivesCyanoacrylate Butyl or isobutyl cyanoacrylate Addition polymerizationFibrin sealants A. Fibrinogen, Factor XIII

B. Thrombin, CaC12Fibrin clot formation

GRF glue Gelatin, resorcinol, formaldehyde Condensation polymerizationHydrogel Block copolymers of PEG, poly(lactic acid) and acrylate esters Photoinitiated addition polymerization

environments, the most prevalent trend was increased mechanical properties when cured at 50°C versus room temperature. Pores were shown to gather near the surface of cooler molds, and near the center in warmer molds for all cement brands. Pore size was also influenced. Small pores were more often present in cements cured at cooler temperatures, with higher-temperature molds producing more large pores. The mechanical properties of all cements were above the minimum regulatory standards. This supports the practice of heating cemented implants to influence interfacial porosity.

Pelletier, M. H., Lau, A. C. B., Smitham, P. J., Nielsen, G., & Walsh, W. R. (2010). Pore distribution and material properties of bone cement cured at different temperatures. Acta Biomaterialia, 6, 886–891. Elsevier. DOI: 10.1016/j.actbio.2009.09.016


relationship between principal categories is illustrated in Figure II.5.8.3.

Zinc phosphate cement is the traditional standard. This material is composed primarily of zinc oxide pow-der and a 50% phosphoric acid solution containing aluminum and zinc. The mixed material sets to a hard, rigid cement (Table II.5.8.1) by formation of an amor-phous zinc phosphate binder. Although the cement is gradually soluble in oral fluids and can cause pulpal irritation, it is clinically effective over 10 to 20 year periods. The bonding arises entirely from penetration into mechanically produced irregularities on the surface of the prepared tooth and the fabricated restorative material. Some interfacial leakage occurs because of cement porosity and imperfect adaptation, but this is usually acceptable since the film thickness is generally below 100 μm.poly-Electrolyte Cements: Zinc polycarboxylates and Glass Ionomers. Poly(carboxylic acid) cements were developed in 1967 (Smith, 1967, 1998) to provide mate-rials with properties comparable to those of phosphate cements, but with adhesive properties to calcified tissues. Zinc polyacrylate (polycarboxylate) cements are formed from zinc oxide and aqueous poly(acrylic acid) solution (Figure II.5.8.4). The metal ion cross-links the poly-mer structure via carboxyl groups, and other carboxyl

groups complex to Ca ions in the surface of the tissue (Figure II.5.8.5). The zinc polycarboxylate cements have adequate physical properties, excellent biocompatibility in the tooth, and proven adhesion to enamel and dentin (Smith, 1991), but are opaque. The need for a translu-cent material led to the development of the glass-iono-mer cements (GIC).

GICs are also based on poly(acrylic acid) or its copo-lymers with itaconic or maleic acids, but utilize a calcium aluminosilicate glass powder instead of zinc oxide. GICs set by cross-linking of the polyacid with calcium and alu-minum ions from the glass, together with formations of a silicate gel structure. The set structure and the resid-ual glass particles form a stronger, more rigid cement (Table II.5.8.1), but with similar adhesive properties to the zinc polyacrylate cements. Both cements are widely used clinically.Neutralization of Bacterial acids. Acid–base cement formulations are non-stoichiometric, having an excess of the basic powder. This is beneficial in intra-oral appli-cations as the net basicity of the set cement retains the

TaBLE I I .5.8.2 Properties of Dental Cements and Sealants

Strength

Dental Cement/SealantCompressive(Mpa)

Tensile(Mpa)

Elastic Modulus (Gpa)

Fracture Toughness (MN−1.5)

Zinc phosphate 80–100 5–7 13 ~0.2Zinc polycarboxylate 55–85 8–12 5–6 0.4–0.5Glass ionomer 70–200 6–7 7–8 0.3–0.4Dimethacrylate sealant unfilled 90–100 20–25 2 0.3–0.4Dimethacrylate sealant filled 150 30 5 –Dimethacrylate cement 100–200 30–40 4–6 –Dimethacrylate composite 350–400 45–70 15–20 1.6

FIGURE II.5.8.3 Classification of major types of acid–base dental cement based upon the combinations of basic powders of either zinc oxide or alumino-silicate glass with two aqueous acidic liquids.

FIGURE II.5.8.4 Structure of the polyelectrolyte: poly(acrylic acid), as present in aqueous solution, in polycarboxylate and glass iono-mer cements. The carboxyl anions COO− are bridged by divalent or trivalent metal ions during the setting process, forming a salt matrix.

HO On + ZnO

n

n

Zn2+OO + H2O

O OΘ

Θ

FIGURE II.5.8.5 Neutralization of poly(acrylic acid) by zinc oxide to form zinc polycarboxylate cement, in which Zn2+ functions as a bridging ion between pairs of carboxylate groups.


capability of neutralizing in situ erosive acids produced by bacterial metabolism of dietary sugars.resin-Modified Glass-Ionomer Cements. In these cements, the polyacid molecule contains both ionic carboxylate and polymerizable methacrylate groups (Figure II.5.8.6). It is induced to set by both an acid–base reaction and visible light polymerization. These dual-cure cements are widely used clinically. Adhesive bonding but not complete sealing is obtained, because of imperfect adaptation to the bonded surfaces under prac-tical conditions.Dual-Setting resin-Based Cements. Resin cements are fluid or paste-like monomer systems based on aromatic or urethane dimethacrylates. Silanated ceramic fillers are usually present to yield a composite composition. They are normally two-component materials that are mixed to induce setting. They may also be light-cured. These set materials are strong, hard, rigid, insoluble, cross-linked polymers (Table II.5.8.1). Bonding is achieved by mechanical interlocking to surface roughness.Self-Etching resin-Based Cements. In recent mate-rials, reactive adhesive monomers may also be present to avoid the necessity of using a separate dentine-bonding system, as described in the section on Bonding to Dentin via the Hybrid Inter-Phase).

acid-Etch Bonding to Enamel

Dental enamel is a hard, stiff substrate of generally uniform chemical composition. Bonding to enamel is required either in conjunction with bonding to dentin, for cavity restoration or for bonding veneers or orth-odontic brackets directly to enamel of teeth in the upper or lower dental arch. In both cases, the modern approach was pioneered by Buonocore (Roulet and Degrange, 2000; Eliades et al., 2005). This involves conditioning with an aqueous phosphoric acid etchant for about 60 seconds. Subsequent water-rinsing and air-drying leaves a visibly matt surface on the enamel – now differentially etched, with either the enamel-prism cores or the prism peripheries exposed (Figure II.5.8.7). Application of an unfilled resin (monomer mixture) or a moderate-viscos-ity resin-composite allows this to flow across the enamel, penetrating surface porosity. Subsequent hardening of the resin results in its retention via a multitude of micro-scopic tag-like resin extensions into the enamel surface.

Bonding to Dentin via the hybrid Inter-phase

Intense research Effort and advances in this area. Bonding restorative biomaterials to dentin is much more challenging than to enamel, due to the more complex, hydrated substrate. Not only may this involve caries-weakened dentin, but the cut dentin surface has a weakly attached “smear” layer. This is a layer of mixed com-position, including denatured collagen, generated by the thermo-mechanical process of cutting dentin by high-speed burs. This material is mechanically weak and is smeared all over the cut surface, covering the sound den-tine underneath. This must be either removed or signifi-cantly modified.

Nevertheless, recent decades of intensive worldwide research have made remarkable advances both in under-standing the ultra-structure of the resultant bonds, and then devising improved formulations (Nakabayashi and Pashley, 1998; Roulet and Degrange, 2000; Eliades et al., 2005; Moszner and Salz, 2007; Matinlinna and

AI AICOOH

COOH

COOH

COOH

COOH

COOH

C

C

C

C

hν

AI3+

COO

COO

COO

COO

COO

COO

C

C

C

C

FIGURE II.5.8.6 Formation of the matrix of a “resin modified” glass ionomer cement (RMGIC), by means of C=C bond polymerization, concurrently with ionic salt bridges with metal ions derived from an ion-leachable glass.

FIGURE II.5.8.7 Surface of acid-etched dental enamel. In this image, there are alternating regions where the bundles of prism structures are viewed: (i) transversely (end-on); or (ii) semi-longitudi-nally. In the case of the transverse prisms, the prism peripheries have been preferentially etched, leaving the prism cores visble.


Mittal, 2009). At present, the outstanding challenges are to enhance the long-term durability of adhesive bonds against both intra-oral stresses and biochemical degra-dation (Breschi et al., 2008; Watts et al., 2009). There is also a secondary goal of formulating simpler, and thus time-saving, application protocols.

Early Unsuccessful approaches

It was recognized early that simply applying hydropho-bic resins to dentin gave little benefit. Moreover, for several decades it was deemed clinical malpractice to apply an acid etchant to dentin. Development of adhe-sives proceeded slowly until the critical insights con-cerning dentin hybridization and the hybrid-layer were achieved, as discussed in the section titled Hybrid-Layer Creation via Three-Stage Approach: Etch; Prime; Bond.

Chemistry of Etchants, primers, and Bonding agents

Modern effective dentine bonding systems can be con-sidered as incorporating a three-fold set of agents, although two or even three of these functions may be combined in a simplified system. The etchant or con-ditioner function is provided by either an aqueous phosphoric acid gel or by acidic monomers. The primer function is classically achieved by a solution of a hydro-philic monomer, such as HEMA, in a solvent such as acetone or ethanol. Subsequently, a “bonding agent” consisting of unfilled resin monomers is applied to the dentine.

hybrid-Layer Creation via Three-Stage approach: Etch; prime; Bond

The optimal approach is to use a separate acid etchant, and to apply this to the dentin for about 10 seconds, which removes the smear layer (Figure II.5.8.8). This demineralizes the outer dentin layers, leaving residual type I collagen fibrils and proteoglycans. Washing the dentine removes the acid and salt residues. However, drying results in collapse or compaction of the colla-gen. The second agent, the primer, has the role of re-expanding the collagen, thereby restoring its porosity (Figure II.5.8.9). This permits the third agent – unfilled hydrophobic “bond” monomers – to permeate and seal the collagen. When these are photo-cured they form an entangled interpenetrating network (IPN) with the colla-gen fibers. The resultant interfacial IPN region, typically 4 μm in thickness, is known as the hybrid zone. Effec-tively, the polymerized resin is taking the place of the original mineral phase in this region. The strength of this bonding zone depends on the continuing integrity of its component elements, principally the collagen fibrils and the polymerized monomers.

FIGURE II.5.8.8 Dentin surface following etching with phosphoric acid solution. The dentine tubules are clearly visible.

FIGURE II.5.8.9 Cross-section of acid-etched and demineralized dentin. The collagen fibrils are visible (arrowed) within a tubule.


The Quest for Simplified Clinical procedures

Clinical success depends on understanding the rationale of each step and meticulous observance of manufac-turer’s instructions. The market has led to availability of simplified systems, where two or more steps may be combined into one, including self-etching systems (Figure II.5.8.10). Earlier “one step” systems of this type gave reduced performance, but newer products incor-porating clever delivery devices are more successful. Nevertheless, there are many systems available where a simplification of clinical application procedures is done to the detriment of bonding efficacy.

aging and Stability of the Bonded Interface

When properly applied, good adaptation and bonding of the adhesive and restorative material to dentin can be achieved in the short- and medium-term. Also, a good initial seal of the interfacial zone can be achieved. How-ever, a number of potential weaknesses have been identi-fied (Breschi et al., 2008).

1. Insufficient resin impregnation of dentin: There are circumstances where the demineralized collagen is not fully supported by infiltrated resin, leaving a mechanically weak zone.

2. High permeability of the bonded interface: The hybrid zone is susceptible to “nano-leakage” by small molecules and ions, which may provide a pathway for degradation reactions over time.

3. Sub-optimal polymerization: The efficiency of polymerization of the infiltrated monomer may be

compromised by the presence of water and dissolved oxygen, thereby weakening the IPN structure.

4. Phase separation: Monomer mixtures in situ of vary-ing hydrophilicity may spontaneously phase separate upon polymerization.

5. Activation of endogenous collagenolytic enzymes:Recent studies have revealed that dentin matrices may be slowly degraded over time by dentin-derived proteolytic enzymes, particularly matrix metallo-proteinases (MMPs). Collagen integrity within the hybrid layers may be promoted by incorporation of chlorhexidine, an anti-bacterial agent with MMP-inhibiting properties.

FIGURE II.5.8.10 Components of self-etching enamel-dentin adhesives. (After Moszner and Salz, 2007.) The adhesive systems usually incorporate monofunctional co-monomers and additives, including photo-initiators, solvents, stabilizers, and fillers. (Recent Developments of New Components for Dental Adhesives and Composites. Macromolecular Materials and Engineering. 292, 245–271. 2007. Wiley. DOI: 10.1002/mame.200600414)

Inhibitors for the preservation of the hybrid Interfacial Zone Between adhesives and human Dentin

The outstanding resistance of the collagenous dentin matrix against thermal and proteolytic disruption has been attributed to the high degree of intermolecular cross-linking and tight mechanical weave of this specialized connective tissue. However, great attention to the potential proteolytic activity of dentin has been raised since com-plexed and active forms of matrix metalloproteinases (MMPs) were identified in either non-mineralized or mineralized compartments of human dentin. MMPs belong to a group of zinc- and calcium-dependent enzymes that have been shown to be able to cleave native collagenous tissues at neutral pH in the metabolism of all connective tissues. Galardin is a synthetic MMP-inhibitor with potent activity against MMP-1, -2, -3, -8, and -9. It also partially preserved the mechanical integrity of the hybrid layer created by a two-step etch-and-rinse adhesive after artificial aging.

Breschi, L., Martin, P., Mazzoni, A., Nato, F., & Carrilho, M. et al. (2010). Use of a specific MMP-inhibitor (galardin) for preservation of hybrid layer. Dent. Mater., 26, 571–578. Elsevier: DOI: 10.1016/j.dental.2010.02.007


Incorporation of anti-Bacterial Functionality

Efforts have been made for some time to incorporate anti-bacterial agents in either composites or dentin- adhesives, with the latter showing greater promise (Imazato, et al. 1998, 1999; Imazato, 2003).

Stress-Development in adhesive Joints due to polymerization Shrinkage

Inevitability When Using Methacrylate Matrix Composite Biomaterials. As already noted, adhesive layers, when solidified by polymerization, are accompanied by shrinkage phenomena. In restorative dentistry, tooth cavities – caused by carious decay – are now principally repaired by tooth-coloured polymer–ceramic composites, instead of silver amalgam fillings. These bulk materials also exhibit setting shrinkage which exacerbates the inter-facial stresses on the adhesive joints. This is an inescapable feature of the currently dominant dimethacrylate “resin” composites. Despite numerous attractive properties of these biomaterials, many adverse clinical problems can arise if this shrinkage problem is not adequately managed (Watts et al., 2009). It is a major potential cause of adhe-sive failure.

Molecular Origins of Shrinkage and polymerization-Kinetics

When methacrylate monomers undergo polymeriza-tion, this involves the conversion of C=C double bonds to C–C single bonds. The original monomer molecules are no longer separated by a mean intermolecular (van der Waals) distance, but become connected via the C–C bonds. This implies a closer packing of the molecular units – or in other terms, molecular densification. For example, with methyl methacrylate (MMA), the volume change per mole of methacrylate groups is:

ΔVC = C = 22.5 cm3 mole − 1

This localized molecular shrinkage normally occurs throughout the bulk of the material and so produces a macroscopic shrinkage.

For multi-methacrylate monomers of higher C=C functionality (f), and for a degree of conversion (DC) that may be less than 100%, the volumetric shrinkage is given by:

volm

VDC fV

22.5 100

where: V is volume and Vm is molar volume. This can easily be generalized for a mixture of monomers (Silikas et al., 2005).

Both methacrylate and acrylate polymerization occur by a free radical mechanism. This may be activated by

various means, although photo-activation using vis-ible blue light is the most widespread option in den-tistry, with chemical activation being essential for bone cements (see Mechanism of Setting of PMMA/MMA Dough).

In dimethacrylate dental monomers, there are two C=C polymerizable groups at either end of the molecule. Statistically, both C=C groups react in a high proportion of cases, generating a tightly cross-linked matrix network. However, some monomers only react via one C=C group, leaving the other C=C group unreacted and pendant. This is because following a rapid auto-acceleration phase, the mate-rial vitrifies and the chain-end free radicals become locally trapped within the glassy network. The rate of polymerization then rapidly diminishes (auto-decel-eration) towards zero (Watts, 2005). Thus, at 37°C, the degree of conversion of monomer is only about 65%. Nevertheless, physical properties are otherwise satisfactory.

Factors that affect the Magnitude and Vector Direction of Shrinkage Stress

When an adhesively-bonded resin-composite is placed between two or more opposing cavity walls, then the material is not able to shrink freely upon polym-erization. Therefore, providing adhesive bonding is achieved to the walls, a state of “shrinkage-stress” will be generated in the restored tooth system. However, material at or in proximity to a free, un-bonded surface will be able to shrink freely, with minimal local-stress generation.

Different tooth cavity sizes and shapes involve differ-ing cavity configuration factors (Cf) of bonded to non-bonded interfacial surface areas.

Hence:

Cf = Area bonded / Area non−bonded

other factors being equal, cavities with a high C-factor will be more highly prone to shrinkage stress (Figure II.5.8.11).

Nevertheless, it has been recognized recently that other factors are seldom equal. In particular, the mass (or volume) of setting material commonly varies between different cavity shapes or designs. It is the composite mass that drives the shrinkage stress magnitude, and so if this is reduced appreciably, the shrinkage stress will also be reduced (Watts et al., 2009).

In complex cavity shapes, Finite Element Analysis (FEA) has been fruitfully applied to determine the dis-tribution and vector directions of shrinkage stress. This requires realistic tooth and cavity shapes, and allow-ance being made for the visco-elastic character of both the setting material and dentin tooth tissue. In the interior of a cavity, the repair biomaterial is subject to a state of generalized plane strain, whereas close to


the cavity opening, the material is subject to a state of plane stress.

Strategies to Minimize Shrinkage Stress in Bonded Cavities of Low Compliance

The compliance of the cavity walls is a further significant factor. When the cavity walls are relatively thick, then the compliance will be low and strategies must be sought to minimize shrinkage-stress. One common clinical prac-tice is to apply the material in a succession of layers, either parallel, or at an angle, to the base of the cav-ity. These layers are adhesively bonded to the walls and photo-cured in sequence. This can reduce the net effect of shrinkage stress. The effect of shrinkage stress within restored molar teeth can be directly measured by detec-tion of inward displacement of cavity walls and tooth cusps (Watts et al., 2009).

Development and Chemistry of New Low Shrinkage Monomers, Including Siloranes

The perceived significance of the above shrinkage problem has led to intense efforts to produce improved formula-tions, including novel monomer chemistry. An innovative approach is the development of four branched silorane monomers (Figure II.5.8.12) that are photo-cured via a cationic polymerization mechanism (Weinmann et al., 2005). These materials exhibit less than 1% shrinkage on setting, and are strongly hydrophobic. They require a spe-cialized system adhesive to bond to enamel and dentin.

Soft Tissue adhesives and Sealants

performance requirements. Most soft tissue adhe-sives are intended to be temporary. That is, they are removed or degrade when wound healing is sufficiently advanced for the tissue to maintain its integrity. Effec-tive adhesion can be obtained on dry skin or wound surfaces by using wound dressing strips with acrylate-based adhesives. However, on wound surfaces that are wet with tissue fluid or blood, the adhesive must be able to be spread easily on such a surface, provide adequate working time, develop and maintain adhe-sion, desirably provide hemostasis, facilitate wound healing, and maintain biocompatibility. Positive anti-microbial action would be an additional advantage (Ikada, 1997).historical Overview. Few, if any, systems comply with all these requirements. Currently, there are two principal systems in widespread clinical use – cyano-acrylate esters and fibrin tissue adhesives. Another system, based on a gelatin-resorcinol-formaldehyde (GRF) combination, still receives limited use (Table II.5.8.3). For many years there have been studies under-taken using bioadhesives, especially polypeptides from marine organisms (mussel adhesive) (Sierra and Salz, 1996; Chu et al., 1997). An important new develop-ment by Messersmith and co-workers (Brubaker et al., 2010) has led to synthesis of branched polyethylene glycol (PEG) cores with endgroups derivatized with catechol, a functional group abundant in mussel adhe-sive proteins.

FIGURE II.5.8.11 Schematic cavities in molar teeth. Cavity (C) shows a greater ratio of bonded to non-bonded area than is the case with cavities (A) and (B). This has consequences for the magnitude of shrinkage stress established at the biomaterial–tissue interfaces within each cavity, for a given mass of a specific material.


Cyanoacrylate Esters

Chemistry. These esters are fluid, water-white mono-mers that polymerize rapidly by an anionic mechanism in the presence of weak bases such as water or NH2 groups. Initially, methyl cyanoacrylate (Figure II.5.8.2)

was used, but in the past two decades isobutyl and n-butyl cyanoacrylate have been found more accept-able. The higher cyanoacrylates spread more rapidly on wound surfaces, and polymerize more rapidly in the presence of blood. Furthermore, they degrade more slowly over several weeks, in contrast to the methyl ester, which hydrolyzes rapidly yielding formaldehyde that results in an acute inflammatory response.performance. These materials achieve rapid hemo-stasis, as well as a strong bond to tissue. However, the polymer film is somewhat brittle and can be dislodged on mobile tissue, and the materials can be difficult to apply to large wounds. Because of adverse tissue response and production of tumors in laboratory animals, cyanoac-rylates are not approved for routine clinical use in the United States, although a commercial material based on n-butyl cyanoacrylate is approved by several other countries.

(A) (B)

FIGURE II.5.8.12 (A) Structure of a 4‐branched silorane molecule. The 3‐member functional epoxide groups react via a cationic mecha-nism to give a ring‐opening polymerization. This by itself creates molecular expansion upon setting, and can be used to formulate low‐shrinkage materials. (B) A four arm catechol‐terminated PEG (cPEG) adhesive precursor inspired by the protein glues of marine mussels. This rapidly forms adhesive hydrogels under oxidizing solution conditions.

TaBLE I I .5.8.3 Characteristics of Soft Tissue Adhesives

property Cyanoacrylate Fibrin Glue GrF

Ease of application Poor Excellent PoorSet time Short Medium MediumTissue bonding Good Poor ExcellentPliability Poor Excellent PoorToxicity Medium Low HighResorbability Poor Good PoorCell infiltration Poor Excellent Poor

(After Ikada, 1997.)

The relationship Between Soft Tissue adhesion and Drug Delivery

Soft tissue adhesives are not only required for surgical wound heal-ing, but also for the controlled delivery of drugs to specific sites, of which the mucosal membranes are particularly suitable. The process of mucoadhesion involving a polymeric drug delivery platform is a complex one that includes wetting, adsorption, and interpenetration of polymer chains among various other processes. The success and degree of mucoadhesion bonding is influenced by various polymer-based properties, such as the degree of cross-linking, chain length, and the presence of various functional groupings. The attractiveness of mucosal-targeted controlled drug delivery of active pharmaceuti-cal ingredients has led formulation scientists to engineer numerous polymeric systems for such tasks. Formulation scientists have at their disposal a range of in vitro and in vivo mucoadhesion testing set-ups in order to select candidate adhesive drug delivery platforms. As such, mucoadhesive systems have found wide use throughout many mucosal covered organelles for delivery for local or systemic effect. Evolution of such mucoadhesive formulations has developed from first-generation charged hydrophilic polymer networks to more spe-cific second-generation systems based on lectin, thiol, and various other adhesive functional groups.

Andrews, G. P., Laverty, T. P., & Jones, D. S. (2009). Mucoadhesive polymeric platforms for controlled drug delivery. Eur. J. Pharm. Bio-pharm., 71, 505–518.


The current uses are as a surface wound dressing in dental surgery, especially in periodontics, and in life-threatening applications such as brain arteriovenous malformations. Reports of sarcomas in laboratory ani-mals (Reiter, 1987), evidence of in vitro cytotoxicity (Ciapetti et al., 1994), and lack of regulatory approval have restricted their further use, in spite of work on syn-thesis of new types of cyanoacrylate.

Fibrin Sealants

Formulation, presentation, and Setting processes. Fibrin sealants involve the production of a synthetic fibrin clot as an adhesive and wound-covering agent. The concept of using fibrin dates back to 1909, but was placed on a specific basis by Matras in 1972 (Matras, 1972). The commercial materials first avail-able consisted of two solutions that are mixed imme-diately before application to provide a controlled fibrin deposition. Later a “ready-to-use” formulation (Tisseel Duo) was introduced (Schlag and Redl, 1987). The essential components of these solutions are as follows:

Solution a: Solution B:Fibrinogen ThrombinFactor XIII CaCl2

Fibrinogen is at a much higher concentration (~70 mg/ml) than that in human plasma. On mixing the two solutions, using a device such as a twin syringe with a mixing nozzle, a reaction similar to that of the final stages of blood clotting occurs. Polymerization of the fibrinogen to fibrin monomers and a white fibrin clot is initiated under the action of thrombin and CaCl2. Apro-tinin, an inhibitor of fibrinolysis, may also be included in solution A. The composition may be adjusted to promote hemostasis, for example, or to minimize persistence of the clot to avoid fibrosis.advantages and applications. There are now several commercial products available in Western Europe, Japan, and Canada; but they were not approved for use in the USA until 1998 due to the FDA concerns about viral con-tamination. Fibrinogen for these commercial materials is manufactured from the pooled plasma of selected donors, using processes such as cryoprecipitation. The material is subjected to in process virus activation, and routinely screened for hepatitis and HIV. To minimize these risks recent processes produce the fibrinogen in a “closed” (sin-gle donor) blood bank or utilize the patient’s own blood. This autologous fibrin glue is preferred, but its quality is partly determined by the fibrinogen level in the donor plasma (Ikada, 1997). Other complications include for-mation of antibodies and thrombin inhibitors, as well as potential risks of BSE (bovine spongiform encephalopa-thy) if bovine thrombin is used. More recently, human-derived thrombin has been employed.

Originally developed during World War II to stop bleeding from battle injuries, fibrin sealants are presently used during surgery for several different purposes:

• to control bleeding in the area where the surgeon is operating;

• to speed wound healing; • to seal off hollow body organs or cover holes made

by standard sutures; • to provide slow release delivery of medication to

tissues exposed during surgery.

Fibrin sealants have several advantages over older methods of hemostasis. They speed up the forma-tion of a stable clot; they can be applied to very small blood vessels and to areas that are difficult to reach with conventional sutures; they reduce the amount of blood lost during surgery; they lower the risk of postoperative inflammation or infection; and they are biodegradable by the body during the healing process. They are particularly useful for minimally invasive procedures, and for treating patients with blood clot-ting disorders.

The adhesive strength is not as high as that of cyano-acrylates, but it is adequate for many clinical situations. Thorough mixing of the ingredients and application techniques or devices that allow uniform spreading, are essential to success.

The material has been used in a wide variety of sur-gical techniques for hemostasis and sealing involving thoracic–cardiovascular, neurologic, plastic, and oph-thalmic surgery, and as a biodegradable adhesive scaf-fold for meshed skin grafts in burn patients (Sierra and Salz, 1996; Ikada, 1997; Feldman et al., 1999). A sym-posium has reviewed clinical uses (Spotnitz, 1998). An auxiliary aspect of fibrin sealants is their use as a delivery vehicle at local sites for antibiotics and growth factors (Sierra and Salz, 1996).

Attempts to modify fibrin sealant have been directed toward: (1) improvements in ease of application and con-trol of setting by use of a one-component light- activated product (Scardino et al., 1999); (2) improvements in strength and performance by addition of fibrillar colla-gen (Sierra and Salz, 1996); and (3) development of a formulation containing gelatin (Ikada, 1997).

Gelatin-resorcinol-aldehyde Glues

Formulations. Gelatin is an animal protein with adhe-sive properties and chemical similarity to connective tissue. This glue was developed in the 1960s by Falb and co-workers (Falb and Cooper, 1966; Cooper et al., 1972) as a less toxic material than methyl cyanoacrylate. The material is fabricated by warming a 3:1 mixture of gelatin and resorcina, and adding an 18% formaldehyde solution. (This glue is called “GRF.”) Cross-linking of the gelatin and resorcinol by the formaldehyde takes place in about 30 seconds.


Limited applications. The material was used in a vari-ety of soft tissue applications, but technical problems and toxicity have limited its application in recent years to aortic dissection (Ikada, 1997). In attempts to overcome the toxicity and potential mutagenecity/carcinogenicity of the formaldehyde component, modified formulations have been developed in which other aldehydes such as glutaraldehyde and glyoxal (Ennker et al., 1994a, b) are substituted for the formaldehyde. Favorable results with this material (GR-DIAL) have been reported (Ennker et al., 1994a, b). Concerns over toxicity remain, how-ever, and this material has not received FDA approval for commercial use. Less toxic gelatin cross-linking agents have been investigated in vitro without substan-tial improvement.

The three aforementioned types of adhesives have undergone extensive clinical trials in Europe and Japan. Fibrin sealant is the most widely used material currently. However, all of these systems have significant deficien-cies. This is illustrated by their relative characteristics listed in Table II.5.8.3 (Ikada, 1997).

Bioadhesives

Bioadhesives are involved in cell-to-cell adhesion, adhe-sion between living and nonliving parts of an organ-ism, and adhesion between an organism and foreign surfaces. Adhesives produced by marine organisms, such as the barnacle and the mussel, have been exten-sively investigated over the past 40 years because of their apparent stable adhesion to a variety of surfaces under adverse aqueous conditions. These studies have shown that these organisms secrete a liquid acidic pro-tein adhesive that is cross-linked by a simultaneously secreted enzyme system. The bonding probably involves hydrogen bonding and ionic bonding from the acidic groups (Waite, 1989).

The adhesive from the mussel has been identified as a polyphenolic protein, molecular mass about 130,000 Daltons, which is cross-linked by a catechol oxidase sys-tem in about 3 minutes. A limiting factor in the practical use of this material is the difficulty of extraction from the natural source. The basic unit of the polyphenolic pro-tein has been identified as a specific decapeptide (Waite, 1989; Green, 1996). Recombinant DNA technology and peptide synthesis have been used in attempts to produce an affordable adhesive with superior properties. Lit-tle information has been reported on the performance (including biocompatibility) of these materials.

hydrogel Sealants

A new approach is the development of synthetic sealants based on poly(ethylene glycol) (PEG) hydrogels that are derived from the original work of Sawhney, Pathak, and Hubbell (1993). A family of fully synthetic, implantable, resorbable hydrogels intended for use as surgical sealants

has been developed (Moody et al., 1996; Ranger et al., 1997; Alleyne et al., 1998; Tanaka et al., 1999), for bar-rier coatings (West et al., 1996), and for drug delivery matrices (Lovich et al., 1998). The hydrogels are formed by in situ deposition of aqueous formulations based on specialized dendritic macromolecules, followed by pho-topolymerization to highly cross-linked structures.

The macromers are reactive block copolymers con-sisting of a water-soluble core, such as polyoxyethylene, flanking biodegradable oligomers such as poly(lactic acid) or poly(trimethylene carbonate), and polymerizable end caps such as acrylate esters (Sawhney et al., 1993). Control of the physical properties and degradation rates is achieved by specifying the molecular structures and concentration of the reactants in the formulation (Dai et al., 2011). Photopolymerization can be effected by ultraviolet or visible light using appropriate photoinitia-tors. Typically, the initiator system eosin Y spirit-soluble (EYss)/triethanolamine (TEA) is employed with visible illumination in the 450–550 nm range.

In most applications for these hydrogels, strong bond-ing to tissue is required. This is achieved by use of a two-part sealant system consisting of primer and topcoat. Strong, durable bonding to a wide variety of internal tis-sues has been demonstrated (Coury et al., 1999).

New research Directions: Biomimetic approaches

As a result of the experience of recent decades, the prob-lems involved in developing an adhesive system for both soft and hard tissues have been addressed and identified. A practical limitation in many systems remains ease of manipulation and application. For example, the effec-tiveness of the fibrin sealant is critically dependent on proper mixing of the ingredients and uniform applica-tion. It has proved difficult to reconcile short- and long-term biocompatibility needs with chemical adhesion mechanisms that use a reactive monomer system.

Where relatively temporary (less than 30 days) adhe-sion is required, as in wound healing, systems based on natural models that allow biodegradation of the adhesive and interface, and subsequent normal tissue remodeling, appear to merit further development.

For longer-term (years) durability in both soft and hard tissues, hydrophilic monomers and polymers of low toxic-ity that can both diffuse into the tissue surface, and form ionic bonds across the interface, seem to be a promising approach. Evidence has been obtained of the need for hydrophobic–hydrophilic balance in adhesive monomer systems (Nakabayashi and Pashley, 1998), and the use of hydrophilic monomers such as hydroxyethyl methacrylate in commercial materials has facilitated surface penetration.

Recent development trends in soft tissue adhesives appear to reflect these approaches. Newer materi-als (Chu, 1997) comprise cross-linked collagen mate-rials, light-cured polymerizable and biodegradable


polyethylene glycols, photopolymerized derivatized collagen (DeVore, 1999), a bioresorbable hemostatic collagen-derived matrix with thrombin, and serum albumin cross-linked with a derivatized poly(ethylene glycol) sealant. The development of synthetic pep-tides and materials based on human recombinant components are being investigated. However, the most promising development for hydrogels inspired by mussel adhesives appears to be the approach of the Messersmith group (Brubaker et al., 2010). Rather than using natural polypeptides, the synthesis of catechol-terminated PEG solutions allows a hydo-gel to be formed in about 30 seconds. Upon implan-tation, the cPEG adhesive elicited minimal acute or chronic inflammatory response in C57BL6 mice, and maintained an intact interface with supporting tissue for up to one year. In situ cPEG adhesive formation was shown to efficiently immobilize transplanted islets at the epididymal fat pad and external liver surfaces, permitting normoglycemic recovery and graft revascu-larization. These findings establish the use of synthetic, biologically-inspired adhesives for islet transplantation at extrahepatic sites.

Another significant deployment of PEG block copoly-mers and hyaluronic structures to make photopolymer-izable formulations is exemplified by the Tirelli group (Tirelli et al., 2002; Park et al., 2003).

On calcified surfaces, the use of hydrophilic electro-lytes, such as the polycarboxylates, has demonstrated

that proven ionic bonding in vitro can also be achieved in vivo (Smith, 1998). An advantage of such systems is that surface molecular reorientations can improve bonding with time (Peters et al., 1974). Encouraging preliminary results have been obtained with new glass ionomer hybrid systems, and there is considerable scope for the future developments of modified polyelectrolyte cements.

The development of more efficient adhesives and seal-ants that, in addition to enhancing the durability of cur-rent applications, would permit new applications such as osteogenic bone space fillers, percutaneous and permu-cosal seals, and functional attachment of prostheses is still a challenging problem for the future.

Finally, a major biomimetic initiative is the creation of self-healing polymer composites (Trask et al., 2007). Most materials in nature are themselves self-healing composite materials. The concept of an autonomic self- healing material, where initiation of repair is integral to the material, is now being considered for engineer-ing applications. This bio-inspired concept offers the designer an ability to incorporate secondary functional materials capable of counteracting service degradation, whilst still achieving the primary, usually structural, requirement. So far this research has been mainly directed towards structural materials (Brown et al., 2005a,b), but it is a short step to apply these concepts and designs to the structural adhesives being designed for hard tissue bonding.

Since 2008 the use of dental amalgam to restore teeth has been for-bidden in Norway, and since June 2009 in Sweden also, mainly because of environmental issues and potential health risks related to its mer-cury content. Other countries may soon follow, so that composites are definitely the materials of choice to directly restore teeth in the least invasive way. Although decayed/fractured teeth can be reconstructed minimally invasively and nearly invisibly using adhesive technology, the clinical longevity of composite restorations is still too short. Bond-ing is also indispensable in the treatment of root caries lesions, the current worldwide prevalence of which increases dramatically with age. However, tooth bonding in the relatively aggressive oral envi-ronment is far from perfect. Within a time frame of 3–5 years adhe-sive restorations lose their marginal seal, leading first to unesthetic discoloration, and eventually caries recurrence. This forces dentists to replace restorations too often, leading with each new intervention to further weakening of the patient’s tooth and, naturally, also to higher public-health costs.

Also in orthopedic medicine, one of the main causes of aseptic loosening of cemented hip replacements is the lack of a compound stable against hydrolysis between the hydrophobic bone cement and the hydrophilic acetabular bone stock. Current hydrophobic bone cements cannot chemically bond to the hydrophilic osseous surface, and thus undergo long-term hydrolytic degradation processes, causing the bone–bone cement interface to de-bond. Bone bonding cements containing functional monomers that can directly adhere to bone, and thus can complement the current purely mechanical stabilization of

orthopedic implants, would therefore be of great benefit in orthopedic surgery.

In an attempt to answer questions of direct importance to the mechanisms of bond degradation/stability, the complex biochemical interplay at the adhesive–tooth interface was studied by analyzing the chemical interaction of the functional monomers phenyl-P and 4-MET with HAp, using correlatively solid-state nuclear magnetic resonance (NMR) and X-ray diffraction (XRD). Clinical data confirmed the overall better performance of a 10-MDP-based adhesive over its precursor, a phenyl-P-based adhesive, while a 4-MET-based adhesive performed somewhat in between the other two. Direct comparison of the inter-facial molecular interactions of the functional monomers phenyl-P and 4-MET with similar data obtained before for the best performing 10-MDP was expected to reveal how functional monomers interact at the adhesive interface. The hypothesis was tested that no difference in chemical interaction with HAp was found among the three functional monomers.

A time-dependent molecular interaction at the interface with stable ionic bond formation of the monomer to hydroxyapatite was found, competing in time with the deposition of less stable calcium phosphate salts. The advanced tooth–biomaterial interaction model gave not only an insight into the mechanisms of bond degradation, but also provided a basis to develop functional monomers for more durable tooth reconstruction.

1Yoshihara et al. (2010). Acta Biomaterialia, doi:10.1016/j.actbio.2010.03.024

Case Study Nano-Controlled Molecular Interaction at Adhesive Interfaces for Hard Tissue Reconstruction1


GLOSSARY OF TERMSBPO: Benzoyl peroxideCf: Cavity Configuration FactorFEA: Finite element analysisGIC: Glass ionomer cementGRF: Gelatin-resorcinol-formaldehydeHA: hydroxyapatiteH-bonds: Hydrogen bondsHEMA: Hydroxy ethyl methacrylateIPN: Inter-penetrating networkMEP: 10-methacryloyloxydecyl dihydrogen phosphate4-MET: 4 methacrloyloxyethyl trimellitic acidMMA: Methyl methacrylateMMP: Matrix metalloproteinasesPEG: Polyethylene glycolPMMA: PolymethylmethacrylateRMGIC: Resin Modified Glass Ionomer CementTCP: Tricalcium phosphateTEGDMA: Triethylene glycol dimethacrylateTHR: Total hip replacement.

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