BIOACTIVE GLASSES. Larry L. Hench and Orjan Andersson. Chapter INTRODUCTION 3.2. PROCESSING

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1 Chapter 3 BIOACTIVE GLASSES Larry L. Hench and Orjan Andersson 3.1. INTRODUCTION It was discovered by Hench and colleagues in 1969 that bone can bond chemically to certain glass compositions. 1 This group of glasses has become known as bioactive glasses, based upon the following definition: A bioactive material is one that elicits a specific biological response at the interface of the material which results in the formation of a bond between the tissues and the material. 2 5 Bioactive glasses have numerous applications in the repair and reconstruction of diseased and damaged tissue, especially hard tissue (bone). Clinical applications are discussed in reviews 3 5 and Chapters 6 12 and One aspect that makes bioactive glasses different from other bioactive ceramics and glass-ceramics is the possibility of controlling a range of chemical properties and rate of bonding to tissues. The most reactive glass compositions develop a stable, bonded interface with soft tissues, as shown by Wilson. 6,7 It is possible to design glasses with properties specific to a particular clinical application. This is also possible with some glass-ceramics, but their heterogeneous microstructure restricts their versatility PROCESSING Bioactive glasses are produced by conventional glass manufacturing methods (Chapter 1). Contamination of the glass must be avoided in order to retain the chemical reactivity of the material. Purity of raw materials must be assured. Analytical grade compounds are typically used for most components. Silica can be added in the form of high purity (flint quality) glass sand, since chemically prepared silicas are difficult to handle without adsorption of water and agglomeration. The choice of raw materials can affect the properties of the glass. Andersson has shown that the use of calcium phosphate compounds that contain crystal water result in glasses that crystallize more easily than if crystal water-free compounds are used. 8 This effect is due to the dissolution of OH ions in the glass structure and the associated decrease of viscosity. See Chapter 22 for a discussion 49

2 50 An Introduction to Bioceramics 2nd Edition of the effects of viscosity on glass formation and crystallization. Preferential vaporization of fluxes will also affect glass viscosity and tendency to crystallize or phase separate, as well as alter the final glass composition. Weighing, mixing, melting, homogenizing and forming of the glass must be done without introducing impurities or losing volatile constituents, such as Na 2 O or P 2. Melting is usually done in the range of C, depending on composition. The phase equilibrium diagram for the Na 2 O-CaO-SiO 2 system shows a ternary eutectic very near the 45S5 glass composition (Table 3.1), which was the original basis for selecting this composition for investigation. There is a very steep liquidus as the composition increases in SiO 2 content, which greatly affects the melting and homogenization behavior of the glass. Only platinum or platinum alloy crucibles or glass melter should be used to avoid contamination of the melt. Bulk specimens can be formed by casting or injection molding in graphite or steel molds. Annealing is crucial, C, because of the high coefficient of thermal expansion of the bioactive glass compositions. Each type of device must have its own annealing schedule established. Bioactive glasses are soft glasses and final shapes can be easily made by machining. Standard machine tools or dental handpieces can be used. Diamond-cutting tools are preferred with copious irrigation, although dry grinding is also possible. If a granulated or powdered material is required, the melt can be rapidly quenched in water or air before grinding and sieving into the desired particle sizes. The glass frit (see Chapter 37) should be rapidly dried to avoid corrosion while in contact with water. Other processing methods used for bioactive glass coatings are described in Chapter 22. Composites made with bioactive glasses are discussed in Chapter COMPOSITIONS The base components in most bioactive glasses are SiO 2, Na 2 O, CaO, and P 2 (Table 3.1). The first, and most well-studied composition, termed Bioglass 45S5 (Registered trademark University of Florida, Gainesville, FL), contains 45% SiO 2, 24.5% Na 2 O, 24.4% CaO and 6% P 2, all in weight percent. 1 The 45S5 composition in mole percent is given in Table 3.1, along with many other compositions investigated for surface reaction kinetics. Hench and coworkers have studied a series of glasses in this four-component system with a constant six weight percent P 2 content. This work is summarized in the ternary SiO 2 -Na 2 O-CaO diagram shown in Fig The figure establishes the bioactive-bonding-boundary

3 Bioactive Glasses 51 Table 3.1. Glass Compositions by Mole %. Designation SiO 2 Na 2 O CaOCaF 2 P 2 B 2 O 3 Al 2 O 3 45S5.4F S #1(S63.5P6) #9(S53P4) #10(S45P7) S S S SF SF SF SF SF SF S(gg) S(gg) S(gg) S(gg) S(gg) S(gg) S(gg) S(gg) (gg) = gel-glass 9 of compositions. 2 5 In region A the glasses are bioactive and bond to bone. In the middle of this area a smaller region is indicated (broken line), within which soft tissue bonding also occurs. Glasses in region B behave as nearly-inert materials and are encapsulated by non-adherent fibrous tissue when implanted. Compositions in

4 52 An Introduction to Bioceramics 2nd Edition A-WGC (high P 2 ) Figure 3.1. Compositional dependence (in weight percent) of bone bonding and softtissue bonding of bioactive glasses and glass-ceramics. All compositions in region A have a constant 6 weight percent of P 2, A/W glass-ceramic has higher P 2 content. Region E (soft-tissue bonding) is inside the dashed line, where I B > 8. ((*) 45S5 Bioglass, ( ) Ceravital, ( ) 55S4.3 Bioglass, and (---) soft-tissue bonding; I B =100/t 0.5bb.) (Reprinted from Hench, L.L. (1991). Bioceramics: From Concept to Clinic, J. Amer. Ceram. Soc., 74, pp , with permission.) region C are resorbed within 10 to 30 days in tissue. In region D the compositions are not technically practical and have not been implanted. The boundary between region A and C depends upon the ratio of surface area of the glass to the effective solution volume of the tissue, as well as the glass composition. Fine glass powders resorb more quickly than bulk implants. Partial substitution of CaO by CaF 2 does not significantly alter the bonebonding behavior. The fluoride additions, however, reduce the rate of dissolution and affect the location of the A C boundary in Fig Substitutions of MgO for CaO or K 2 O for Na 2 O also have little effect on bone bonding. B 2 O 3 and Al 2 O 3 have also been used in bioactive glasses to modify processing schedules and rates of surface reaction. Alumina is especially important in controlling glass surface durability and melting and forming characteristics. However, it is well established that Al 2 O 3, in contrast to B 2 O 3, can inhibit bone bonding. The amount of alumina that is tolerated depends on glass composition, but is generally in the order of weight percent. More alumina can be added to a glass with a high reactivity (high bioactivity) than to a glass which reacts more slowly. The dimensions of the bonebonding boundary (region A in Fig. 3.1) shrink as the percent of Al 2 O 3 increases. Gross and coworkers have shown that the same effect occurs for other multivalent

5 Bioactive Glasses 53 cations, such as Ta 2. Additions of more than 1.5 3% of multivalent ions usually make the glass inactive. 10 In a multi-component system like the SiO 2 Na 2 O CaO P 2 B 2 O 3 Al 2 O 3 system, it is not possible to find a simple relationship between composition and tissue bonding that can be expressed in a two-dimensional diagram, such as Fig Andersson et al. described the in vivo behavior of glasses in this complex system with a phenomenological model developed by regression analysis. 8 The method predicts the in vivo behavior of glasses within certain compositional ranges. The prediction is based upon empirically-determined factors and makes it possible to select glasses for specific applications without having to test them in animals. This method of compositional optimization is described in Chapter 22. It works because glass is an amorphous material and most properties are additive, within certain compositional limits. The role of phosphate in bioactive glasses is interesting. Early on it was assumed that P 2 was required for a glass to be bioactive. However, it is now known that phosphate-free glasses, as well as glass-ceramics in which the phosphate is bound in a siable, relatively insoluble apatite phase, are bioactive. Kokubo and coworkers have shown that the minimal melt-derived glass compositional system for bioactivity is CaO SiO 2, with a compositional limit of about 60 mole percent (Chapter 13). Li et al. have shown that gel-derived glasses in the Na 2 O CaO SiO 2 are bioactive even up to 85 mole percent SiO 2 (Table 3.1). 9 This very broad range of bioactive compositions makes it possible to tailor the reactivity of the glasses for various applications, as discussed in Chapter 32. The role of phosphate in the glass appears only to aid in the nucleation of the calcium phosphate phase on the surface but is not a critical constituent because the surface will adsorb both calcium and phosphate ions from solution PROPERTIES The primary advantage of bioactive glasses is their rapid rate of surface reaction, which leads to fast tissue bonding. Their primary disadvantage is mechanical weakness and low fracture toughness due to an amorphous twodimensional glass network. The tensile bending strength of most of the compositions in Table 3.1 is in the range of MPa, which make them unsuitable for load-bearing applications. Mechanical properties of bone are discussed in Chapter 1. For some applications low strength is offset by the glasses low modulus of elasticity of GPa. The importance of this value, which is close to that of cortical bone, is discussed in Chapter 1. The low strength does not influence

6 54 An Introduction to Bioceramics 2nd Edition the utility of bioactive glasses as a coating, where interfacial strength between metal and the coating is the limiting factor. Low strength also has no effect on use of bioactive glasses as buried implants, in low-loaded or compressively loaded devices, in the form of powders or as the bioactive phase in composites. A new generation of highly bioactive glass-ceramics that also have high strength is described in Chapter REACTION KINETICS The basis of the bone-bonding property of bioactive glasses is the chemical reactivity of the glass in body fluids. The surface chemical reactions result in the formation of a hydroxycarbonate apatite (HCA) layer to which bone can bond. Bonding occurs due to a sequence of reactions. See Hench et al. 5 for a detailed summary and extensive list of references. On immersion of a bioactive glass in an aqueous solution, three general processes occur: leaching, dissolution and precipitation. Leaching is characterized by release, usually by cation exchange with H + or H 3 O + ions, of alkali or alkaline earth elements. Ion exchange is easy because these cations are not part of the glass network; they only modify the network by forming non-bridging oxygen bonds (Chapter 1). The release of network-modifying ions is rapid for glasses in the bioactive compositional region (Region A in Fig. 3.1). This ion exchange process leads to an increase in interfacial ph, to values >7.4. Network dissolution occurs concurrently, by the breaking of S O Si O Si bonds through the action of hydroxyl (OH) ions. Breakdown of the network occurs locally and releases silica into solution in the form of silicic acid [Si(OH) 4 ]. The rate of dissolution of silica depends very much on glass composition. The dissolution rate decreases greatly for compositions of >60% SiO 2 because of the larger number of bridging oxygen bonds in the glass structure. The hydrated silica (SiOH) formed on the glass surface by these reactions undergoes rearrangement by polycondensation of neighboring silanols, resulting in a silicarich gel layer. In the precipitation reaction, calcium and phosphate ions released from the glass, together with those from the solution, form a calcia-phosphate-rich (CaP) layer on the surface. When formed in vitro, the CaP layer is mainly located on top of the silica gel, whereas in vivo it is formed within the gel layer. The calcium phosphate phase that accumulates in the gel surface is initially amorphous (a-cap). It later crystallizes to an HCA structure by incorporating carbonate anions from solution within the a-cap phase. The mechanism of nucleation and

7 Bioactive Glasses 55 Figure 3.2. Bilayer films formed on 45S5 Bioglass after 1 h in rat bone, in vivo (1 Å = 10 1 mm). (Reprinted from Hench, L.L. (1991). Bioceramics: From Concept to Clinic, J. Amer. Ceram. Soc., 74, , with permission.) growth of the HCA layer appears to be the same in vitro and in vivo, and is accelerated by the presence of hydrated silica. Figure 3.2 shows the CaP and silica-rich layers formed on a 45S5 bioactive glass within one hour of implantation in a rat bone. The implant was removed as the bonding sequence was beginning. The data were obtained using Auger electron spectroscopy (AES) combined with Ar ion milling. The analysis is of 50 Å slices of material, which combined together yields a compositional profile of the reaction interface. The silica-rich layer has formed to a thickness of more than 12,000 Å (>1 µm). The CaP layer is already 0.8 µm thick after one hour of reaction. Biological molecules are bonded within the bilayer to a depth of 0.1 µm, as indicated by the C and N signals in the outer layer of the surface. It is important to note that the mixed organic inorganic bonding occurs within a region that has Si as well as Ca and P present. Thus, the reactions on the implant side of the interface with a bioactive glass are: Stage 1: Leaching and formation of silanols (SiOH) Stage 2: Loss of soluble silica and formation of silanols Stage 3: Polycondensation of silanols to form a hydrated silica gel Stage 4: Formation of an amorphous calcium phosphate layer Stage 5: Crystallization of a hydroxycarbonate apatite layer.

8 56 An Introduction to Bioceramics 2nd Edition STAGE Table 3.2. Reaction Stages of a Bioactive Implant. 1 Rapid exchange of Na + or K + with H + or H 3 O + from solution: Si O Na + +H + + OH Si-OH + + Na + (solution) + OH This stage is usually controlled by diffusion and exhibits a t 1/2 dependence. 2 Loss of soluble silica in the form of Si(OH) 4 to the solution, resulting from breaking of Si O Si bonds and formation of Si OH ( silanols) at the glass solution interface: Si O Si + H 2 O Si OH + OH Si This stage is usually controlled by interfacial reaction and exhibits a t 1.0 dependence. 3 Condensation and repolymerization of a SiO 2 -rich layer on the surface depleted in alkalis and alkaline-earth cations: O O O O O Si OH + HO Si O O Si O Si O + H 2 O O O O O 4 Migration of Ca 2+ 3 and PO - 4 groups to the surface through the SiO 2 -rich layer forming a CaO P 2 rich film on top of the SiO 2 -rich layer, followed by growth of the amorphous CaO P 2 -rich film by incorporation of soluble calcium and phosphates from solution. 5 Crystallization of the amorphous CaO P 2 film by incorporation of OH 2, CO3 -, or F anions from solution to form a mixed hydroxyl, carbonate, fluorapatite layer. Table 3.2 summarizes these five reaction stages in more detail. 4,5 See the reference list for details of the experiments used to generate this reaction sequence. For a bond with tissue to occur a layer of biologically active HCA must form. This appears to be the only common characteristic of all the known bioactive implants. The rate of tissue bonding appears to depend on the rate of HCA formation. The kinetics of the reaction stages depend on the glass composition. Fourier transform infrared reflection (FTIR) spectroscopy can be used to determine the reaction rates and mechanism of all five stages of reaction. Figure 3.3 shows the FTIR spectra (using a diffuse reflection stage) of a 45S5 Bioglass implant after 0, 1 and 2 hours in tris-buffer solution at 37 C. The spectra are equivalent to those obtained using a simulated body fluid (see Chapter 37). The peak identifications are based upon previous assignments of IR spectra

9 Bioactive Glasses 57 Figure 3.3. FTIR spectra of a 45S5 Bioglass implant after 0, 1, and 2 h in TBS at 37 C. (Reprinted from Hench, L.L. (1991). Bioceramics: From Concept to Clinic, J. Amer. Ceram. Soc., 74, , with permission.) (see Chapter 17). The alkali-ion-hydronium ion exchange and network dissolution (Stages 1 and 2 in Table 3.2) rapidly reduces the intensity of the Si O Na and Si O Ca vibrational modes and replaces them with Si OH bonds that have only one nonbridging oxygen (NBO) ion. Alkali content is depleted to a depth >0.5 µm

10 58 An Introduction to Bioceramics 2nd Edition Figure 3.4. Time-dependent changes in IR vibrations of the surface of 45S5 Bioglass implant in a 37 C TBS. (Reprinted from Hench, L.L. (1991). Bioceramics: From Concept to Clinic, J. Amer. Ceram. Soc., 74, , with permission.) within a few minutes. Auger electron spectroscopy (AES) showed that by 2 minutes, alkali ion depletion occurred to depths >0.1 µm. 4,5 As the Stage 1 and 2 reactions continue, the single Si OH NBO modes are replaced by H O O Si OH O i.e., Si 2NBO stretching vibrations which are in the range of 930 cm 1, decreasing to 880 cm 1. By 20 minutes, the Si 2NBO vibrations are largely replaced by a new mode assigned to the Si O Si bond vibration between two adjacent SiO 4 tetrahedra (Fig. 3.4).

11 Bioactive Glasses 59 This new vibrational mode corresponds to the formation of the silica-gel layer by the Stage 3 (Table 3.2) polycondensation reaction between neighboring surface silanols. This mode decreases in frequency until it is hidden after one hour by the growing CaP-layer. As early as ten minutes, a P O bending vibration associated with formation of an amorphous CaP layer appears. This is due to precipitation from solution (Stage 4 in Table 3.2). Clark et al. showed (using AES) that by two minutes Ca and P enrichment occurred on the glass surface to a depth of approximately 20 nm. Ogino et al. showed (with AES) that by one hour the CaP layer grew to 200 nm in thickness. 11 Within 40 minutes (Fig. 3.4) the P O bonding vibration is strong and exhibits a continually decreasing frequency as the CaP-rich layer builds. At about 1.5 ± 0.2 h, the P O bending vibration associated with the amorphous calcium phosphate layer is replaced by two P O modes (Fig. 3.3) assigned to crystalline apatite. Concurrent with the onset of apatite crystallization (Stage 5 in Table 3.2) is the appearance of a C O vibrational mode associated with the incorporation of carbonate anions in the apatite crystal lattice, as discussed by LeGeros and LeGeros in Chapter 17. The crystals are nucleated and grow as HCA, the same phase as biological HCA formed in mineralizing tissues. The reason for the equivalence is the similarity of nucleating mechanisms and physiological growth conditions. The C O mode decreases in wave number as the HCA layer grows. By ten hours the HCA layer has grown to 4 µm in thickness, which is sufficient to dominate the FTIR spectra and mask most of the vibrational modes of the silica-gel layer or the bulk-glass substrate. By 100 hours the polycrystalline HCA layer is thick enough to yield X-ray diffraction (XRD) results, as discussed in Chapter 37. The primary 26 and 33 2Θ peaks of HCA are visible with considerable line broadening. By two weeks the FTIR spectra show three P-O vibrational modes and the XRD data are equivalent to biological HCA grown in vivo. Thus, the bioactive glass implant surface provides a substrate that is favorable for the rapid nucleation and growth of biologically-equivalent HCA (Stage 5 in Table 3.2). Differences in the in vivo behavior of various glass compositions are due to the differences in the rate of Stage 5, HCA formation. Table 3.3 summarizes a series of investigations of the effects of glass composition and solution composition on the kinetics of reaction Stages 1 5. The sequence of surface reactions is independent of solution composition. 4,5 However, the presence of Ca and P in a simulated body fluid (SBF) solution accelerates to a small extent the repolymerization of silica (Stage 3) and formation of the amorphous calcium-phosphate

12 Table 3.3. Time for Onset of Reaction Stages 1, 2, 3, 4 and 5 for Bioactive Glasses. Time (min) in tris buffer Composition S5.4F(2) S5 (old) (2) S5 (new) Composition #1 (S63.5P6) NO SPECTRAL CHANGES NOTED Composition #9 (S53P4) Composition #10 (2) (S45P7) SF SF SF SF SF SF S(Gel-Glass) 5 54S(Gel-Glass) 5 58S(Gel-Glass) 5 63S(Gel-Glass) S(Gel-Glass) S(Gel-Glass) S(Gel-Glass) 5 86S(Gel-Glass) 5 60 An Introduction to Bioceramics 2nd Edition

13 Bioactive Glasses 61 Figure 3.5. Effect of glass composition on time for onset of crystallization of hydroxycarbonate apatite (HCA) on the surface in TRIS buffer or simulated body fluid (K-9) solutions. (a-cap) layer (Stage 4). The major effect of solution composition is on the crystallization of HCA (Stage 5). Figure 3.5 shows that the process of HCA crystallization is the same for the various glasses and the tris-buffer or SBF solution K-9, the only difference is the rate of crystallization. The rate of crystallization increases to a small extent in Ca- and P- containing SBF solutions (in 90 minutes rather than 120 minutes). However, Mg ions in SBF slow down formation of the a-cap layer and greatly retard crystallization of HCA on the glass surface. Table 3.3 and Fig. 3.5 show that most of the effects of glass composition are on the time required for HCA crystallization. This finding makes it possible to summarize the relationship between surface reaction rates of bioactive glasses and their in vivo behavior. Figure 3.6 shows the critical compositional relationship between in vitro and in vivo kinetics. For glasses with up to about 53 mole percent Si, HCA crystallization occurs very rapidly on the glass surface, within two hours. These compositions develop a rapid bond with bone and also form an adherent, interdigitating collagen bond with the soft tissues. Glasses with Si content between 53 and 58 mole percent SiO 2 require two to three days to form both the a-cap layer and to crystallize HCA. Such glass compositions are bioactive, but they bond only to bone. When implanted in soft tissues, the fibrous capsule formed around them is parallel to the interface and is non-adherent. Compositions with >60% SiO 2 do not form a crystalline HCA layer even after four weeks in SBF. An amorphous calcium-phosphate layer forms but it does not

14 62 An Introduction to Bioceramics 2nd Edition Figure 3.6. Effect of bioactive glass composition on in vitro reaction kinetics and in vivo tissue response. crystallize to HCA. Such glasses are not bioactive and bond neither to bone nor soft tissues. The in vitro kinetics studies described above were conducted in well specified conditions with the ratio of glass surface area (SA) to solution volume (V) fixed at 0.1 cm 1. Changing the SA/V ratio changes the reaction kinetics. The SA/V ratio relevant to an implant is very difficult to estimate. It depends not only on the surface area of the implant and the size of the cavity into which it is placed but also on tightness of fit, blood flow, metabolic rate of the tissues, inflammation etc. Thus, the kinetics obtained from in vitro analyses are only an estimate of the reaction rates in vivo. However, the compositional effects observed in vitro appear to be equivalent in vivo.

15 Bioactive Glasses TISSUE BONDING The five reaction stages that occur on the material side of the interface do not depend on the presence of tissues. They occur in distilled water, tris-buffer solutions or simulated body fluids. Bonding to tissues requires an additional series of reactions. Chapters 4 and 5 present new findings that explain many of the biological reactions on the tissue side of the bioactive glass tissue interface. The sequence of events associated with formation of a bond with tissues is: Stage 6: Adsorption of biological moieties in the SiO 2 -HCA layer Stage 7: Action of macrophages Stage 8: Attachment of stem cells Stage 9: Differentiation of stem cells Stage 10: Generation of matrix Stage 11: Mineralization of matrix. Rapid growth of HCA agglomerates on a bioactive glass surface incorporates collagen fibrils in vitro (Fig. 3.7) without the presence of cells, enzymes or biological growth factors. The crystals appear to form around the collagen fibrils and form bonds with them on an ultrastructural level, similar to what has been observed by transmission electron microscopy of bone bonded to bioactive glass. 1 The same bonding process of collagen incorporation within the growing gel layer has been observed in soft tissues by Wilson (Figs 3.8 and 3.9). 6,7 Thus, Stage 6 appears to be well established with respect to collagen. Chapter 4 summarizes biological steps Within a week mineralizing bone appears at the interface of the more reactive bioactive glasses (Stage 11). By four weeks the interface is completely bonded to bone without any intervening fibrous tissues. The time sequence of bone formation on bioactive glasses is reviewed in detail elsewhere. 3 5 Figure 3.10 shows a scanning electron micrograph of a cross section of a bioactive glass (S46PO) after eight weeks in rabbit tibia. In the back-scatter mode it is easy to distinguish the characteristic silica-rich (dark) and calcium phosphate-rich (bright) layers. The composition of the glass is SiO 2, 46.0; Na 2 O, 26.0; CaO, 25.0; P 2, 0.0; B 2 O, 2.0 and Al 2 O 3, 1.0 weight percent. Since the glass does not originally contain phosphate it is clearly seen that phosphate is absorbed and that calcium phosphate forms mainly within the silica-rich layer. This emphasizes the role of the silica structure in inducing the HCA formation. The most important criterion for tissue bonding is the mechanical resistance of the tissue implant interface. There is no consensus on the best test

16 64 An Introduction to Bioceramics 2nd Edition Figure 3.7. (a) SEM micrograph of collagen fibrils incorporated within the HCA layer growing on a 45S5 Bioglass substrate in vitro. (b) Close-up (11,300X) of the HCA crystals bonding to a collagen fibril. (Photographs courtesy of C. Pantano.) method to determine interfacial adherence of bioactive implants. Studies by the Florida group of bone bonding to bioactive glasses and glass-ceramics, primarily 45S5 composition, showed very high interfacial strength values using a variety of mechanical test methods. 11 Most of the studies involved loaded prostheses, such as segmental bone replacements (Fig. 3.11) or femoral head prostheses. In most

17 Bioactive Glasses 65 Figure 3.8. Peeling of collagen, which remains adherent to the 4SSS bioactive-glass (G) decalcified section. (Original magnification 250 X.) Figure 3.9. Collagen fibers ( ) in cracks on the 45SS bioactive-glass surface (G) undecalcified section. (Original magnification 100 X.) studies the interface did not fail, fracture occurred either in the implant, such as shown in Fig. 3.11, or in the bone, distal to the implant. Various push-out or pull-out tests have also been reported for bioactive glasses and glass-ceramics. Chapters 13 and 14 discuss the method used by the

18 66 An Introduction to Bioceramics 2nd Edition Figure Scanning electron micrograph of a cross section of a phosphate-free bioactive glass (S46PO) after eight weeks in rabbit tibia. G = bulk glass; Si = Si-rich layer; CP = Ca, P-rich layer; B = bone. Figure Fracture of (BG) 45S5 Bioglass -ceramic segmental bone replace in monkey due to impact torsional loading. Note (B) bonded interface. (Photograph courtesy G. Piotrowski.) (Reprinted from Hench, L.L. (1991). Bioceramics: From Concept to Clinic, J. Amer. Ceram. Soc., 74, , with permission.)

19 Bioactive Glasses 67 Figure Schematic illustration of a specimen in the push-out test fixture. Figure Optical micrograph of a cross section of glass S46PO after push-out. B = bone; RL = reaction layer; G = bulk glass. Direction of loading is indicated by long arrow and fracture line by short arrows.

20 68 An Introduction to Bioceramics 2nd Edition Kyoto group. The push-out test method used by Andersson and colleagues in Finland is illustrated in Fig The conical implant is pushed out of the rabbit tibia at a cross-head speed of 0.5 mm/min. Prior to testing, the bone covering the base of the cone is removed by grinding. The base of the cone is also slightly ground in this process and a flat supporting surface is produced. The conical shape reduces the contribution from the surface roughness. It is difficult to measure the contact area since the bone grows along the surface of the implant and may be present only as a very thin layer. If the glass is sectioned along its axis after the test it is possible to estimate the thickness. With this method interfacial strength values of MPa have been obtained for a number of bioactive glasses. For an inert glass the same test gave a value of 0.5 MPa or less, for glasses just outside the bioactivity border values of 2 3 MPa, and for smooth titanium 2 MPa. Thus, non-bonding biocompatible glasses behave as titanium. When testing bioactive glasses, it is observed that the fracture line in some areas is within the silica-rich layer and in others within bone. Figure 3.13 shows a cross section of glass S46PO after the push out. It is clear that bone has adhered to the glass and fracture has occurred within the bone RATE OF BONDING The rate of development of the interfacial bond between an implant and bone can be referred to as the level of bioactivity. Hench introduced an index of bioactivity as a measure of this. 3 5 The index is given by I B = (100/t 0.5bb ), where t 0.5bb is the time for more than 50% of the surface to be bonded to bone. In the ternary diagram for the compositional dependence of the bioactivity, iso-i B contours have been indicated (Fig. 3.1). Thus, the closer to the bioactivity boundary a glass is, the slower the rate of bonding. When the constant 6 weight percent P 2 content is that in Fig. 3.1, I B goes toward zero as the SiO 2 content is raised close to 60%. Bioactive implants with intermediate I B values do not develop a stable bond with soft tissue. The broken line in Fig. 3.1 indicates the region within which the I B values are sufficiently high for soft tissue bonding. The thickness of the bonding zone is roughly proportional to the I B value and the failure strength of a bioactive bond appears to be inversely proportional to the thickness of the zone. Thus, a very high I B value gives a thick bonding zone and low shear strength. Depending on whether rapid bonding or high shear strength is preferred, different compositions are optimal.

21 Bioactive Glasses 69 REFERENCES 1. Hench, L.L., Splinter, R.J., Allen, W.C. et al. (1972). Bonding Mechanisms at the Interface of Ceramic Prosthetic Materials, J. Biomed. Maters. Res., 2, Hench, L.L. and Wilson, J.W. (1991). Surface-Active Biomaterials, Science, 226, Hench, L.L. (1991). Bioceramics: From Concept to Clinic, J. Amer. Ceram. Soc., 74, Hench, L.L. (1998). Bioceramics, J. Amer. Ceram. Soc., 81, Hench, L.L., Hench, J.W. and Greenspan, D.C. (2004). Bioglass: A Short History and Bibliography, J. Aust. Ceram. Soc., 40, Wilson, J., Pigott, G.H., Schoen, F.J. et al. (1981). Toxicology and Biocompatibility of Bioglass, J. Biomed. Mater. Res., 15, Wilson, J. and Nolletti, D. (1986). Bonding of Soft Tissues to Bioglass, in Christel, P., Meunier, A. and Lee, A.J.C. (eds), Biological and Biomechanical Performance of Biomaterials, Elsevier, Amsterdam, pp Andersson, O.H. (1990). The Bioactivity of Silicate Glass (PhD Thesis, Dept. of Chemical Engineering, Abo Akademi, Finland). 9. Li, R., Clark, A.E., Hench, L.L. (1991). An Investigation of Bioactive Glass Powders by Sol-Gel Processing, J. Appl. Biomaterials, 2, Gross, U. and Strunz, V. (1985). The Interface of Various Glasses and Glass- Ceramics with a Bony Implantation Bed, J. Biomed. Mater. Res., 19, Hench, L.L., Paschall, H.A., Allen, W.C. et al. (1975). Interfacial Behavior of Ceramic Implants, National Bureau of Standards Special Publication, 415, Andersson, O.H., Liu, G., Kangasniemi, K. et al. (1992). Evaluation of the Acceptance of Glass in Bone, J. Maters. Sci. Maters. in Med., 3,