FIG. 8. Phase equilibrium diagram pi calcium phosphates in a warer iitno-sphere. Shaded area is processing range to yield HA• containing implants. (After K. de Grot)!, Attn. New York AnJif. Sei. 12". l<iS8.

compressive strength and fatigue resistance depend on the total volume of porosity. Porosity can be in the form of micropores (<1 fira diameter, due to incomplete sintering) or macropores (>100 ju,m diameter, created to permit bone growth). The dependence of compressive strength (<xc) and total pore volume (Vp) is described in de Groot et al. (1990) by:

Tensile strength depends greatly on the volume fraction of microporosity (Vm):

The Weibull factor (n) of HA implants is low in physiological solutions (h = 12), which indicates low reliability under tensile loads. Consequently, in clinical practice, calcium phosphate bioceramics should be used as (1) powders; (2) small, unloaded implants such as in the middle ear; (3) with reinforcing metal posts, as in dental implants; (4) coatings (e.g., composites); or (5) low loaded porous implants where bone growth acts as a reinforcing phase.

The bonding mechanism of dense HA implants appears to be very different from that described above for bioactive glasses. The bonding process for HA implants is described by Jarcho (1981). A cellular bone matrix from differentiated osteoblasts appears at the surface, producing a narrow amorphous electron-dense band only 3 to 5 /im wide. Between this area and the cells, collagen bundles are seen. Bone mineral crystals have been identified in this amorphous area. As the site matures, the bonding zone shrinks to a depth of only 0.05 to 0.2 ¿¿m (Fig. 2). The result is normal bone attached through a thin epitaxial bonding layer to the bulk implant. TEM image analysis of dense HA bone interfaces show an almost perfect epitaxial alignment of some of the growing bone crystallites with the apatite crystals in the implant. A consequence of this ultrathin bonding zone is a very high gradient in elastic modulus at the bonding interface between HA and bone. This is one of the major differences between the bioactive apatites and the bioactive glasses and glass-ceramics. The implications of this difference for the implant interfacial response to Wolff's law is discussed in Hench and Ethridge (1982, Chap. 14).

resorbable calcium phosphates

Resorption or biodégradation of calcium phosphate ceramics is caused by three factors:

1. Physiochemical dissolution, which depends on the solubility product of the material and local pH of its environment. New surface phases may be formed, such as amorphous calcium phosphate, dicalcium phosphate di-hydrate, octacalcium phosphate, and anionic-substi-tuted HA.

2. Physical disintegration into small particles as a result of preferential chemical attack of grain boundaries.

3. Biological factors, such as phagocytosis, which causes a decrease in local pH concentrations (de Groot and Le Geros, 1988).

All calcium phosphate ceramics biodegrade to varying degrees in the following order:

The rate of biodégradation increases as:

1. Surface area increases (powders > porous solid > dense solid)

2. Crystallinity decreases

3. Crystal perfection decreases

4. Crystal and grain size decrease

5. There are ionic substitutions of CO2", Mg2+, and Sr2+ in HA

Factors that tend to decrease the rate of biodégradation include (1) F substitution in HA, (2) Mg2 ' substitution in /3-TCP, and (3) lower /3-TCP/HA ratios in biphasic calcium phosphates.


Christel, P., Meunier, A., Dorlot, J. M., Crolet, J. M., Witvolet, J., Sedel, L., and Boritin, P. (1988). Biomechanical comparability and design of ceramic implants for orthopedic surgery, in Bioceramics: Material Characteristics versus In-Vivo Behavior, P. Ducheyne and J. Lemons, eds. Ann. New York Acad. Sei., Vol. 523, p. 234.

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de Groot, K. (1983). Bioceramics of Calcium-Phosphate, CRC Press, Boca Raton, FL.

de Groot, K. (1988). Effect of porosity and physicochemical properties on the stability, resorption, and strength of calcium phosphate ceramics, in Bioceramics: Material Characteristics versus In-Vivo Behavior, Ann. New York Acad. Sei., Vol. 523, p. 227.

de Groot, K., and Le Geros, R. (1988). Position papers in Bioceramics: Materials Characteristics versus In-Vivo Behavior, P. Ducheyne and J. Lemons, eds. Ann. New York Acad. Sei., Vol. 523, pp. 227, 268, 272.

de Groot, K., Klein, C. P. A. T., Wolke, J. G. C., and de Blieck-Hogervorst, J. (1990). Chemistry of calcium phosphate bioceramics, in Handbook on Bioactive Ceramics, T. Yamamuro, L. L. Hench, and J. Wilson, eds. CRC Press, Boca Raton, FL, Vol. II, Ch. 1.

Gross, V., Kinne, R., Schmitz, H. J., and Strunz, V. (1988). The response of bone to surface active glass/giass-ceramics. CRC Crit. Rev. Biocompatibility 4: 2.

Gross, V., and Strunz, V. (1985). The interface of various glasses and glass-ceramics with a bony implantation bed. ]. Biomed. Mater. Res. 19: 251.

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Hench, L. L. (1988). Bioactive ceramics, in Bioceramics: Materials Characteristics versus In-Vivo Behavior, P. Ducheyne and J. Lemons, eds. Ann. New York Acad. Sei., Vol. 523, p. 54.

Hench, L. L. (1991). Bioceramics: From concept to clinic. J. Am. Ceram. Soc. 74: 1487-1510.

Hench, L. L. (1994). Bioactive ceramics: Theory and clinical applications in Bioceramics-7, O. H. Anderson and A. Yli-Urpo, eds. Butterworth-Heinemann, Oxford, England, pp. 3-14.

Hench, L. L., and Clark, D. E. (1978). Physical chemistry of glass surfaces. ]. Non-Cryst. Solids 28(1):83-105.

Hench, L. L., and Ethridge, E. C. (1982). Biomaterials: An Interfacial Approach. Academic Press, New York.

Hench, L. L., and Wilson, J. W. (1993). An Introduction to Bioce-ramics, World Scientific, Singapore.

Hench, L. L., and Wilson, J. W. (1996). Clinical Performance of Skeletal Prostheses, Chapman and Hall, London.

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Holand, W., and Vogel, V. (1993). Machineable and phosphate glass-ceramics, in An Introduction to Bioceramics, L. L. Hench and J. Wilson, eds. World Scientific, Singapore, pp. 125-138.

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Hulbert, S. F., Bokros, J. C, Hench, L. L., Wilson, J., and Heimke, G. (1987). Ceramics in clinical applications: Past, present, and future, in High Tech Ceramics, P. Vincenzini, ed. Elsevier, Amsterdam, pp. 189-213.

Jarcho, M. (1981). Calcium phosphate ceramics as hard tissue prosthetics. Cltn. Orthop. Relat. Res. 157: 259.

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200 porous hydroxyapatite. Dent. Clin. North Am. 30: 49-6?. Wilson, J. (1994). Clinical Applications of Bioglass Implants, inBiocer-amics-7, O. H. Andersson, ed. Butterworth-Heinemann, Oxford, England.

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2.7 Natural Materials loannis V. Yannas

Natural polymers offer the advantage of being very similar, often identical, to macromolecular substances which the biological environment is prepared to recognize and to deal with metabolically (Table 1). The problems of toxicity and stimulation of a chronic inflammatory reaction, which are frequently provoked by many synthetic polymers, may thereby be suppressed. Furthermore, the similarity to naturally occurring substances introduces the interesting capability of designing biomaterials which function biologically at the molecular, rather than the macroscopic, level. On the other hand, natural polymers are frequently quite immunogenic. Furthermore, because they are structurally much more complex than most synthetic polymers, their technological manipulation is much more elaborate. On balance, these opposing factors have conspired to lead to a substantial number of biomaterials applications in which naturally occurring polymers, or their chemically modified versions, have provided unprecedented solutions.

An intriguing characteristic of natural polymers is their ability to be degraded by naturally occurring enzymes, a virtual guarantee that the implant will be eventually metabolized by physiological mechanisms. This property may, at first glance, appear as a disadvantage since it detracts from the durability of the implant. However, it has been used to advantage in biomaterials applications in which it is desired to deliver a specific function for a temporary period of time, following which the implant is expected to degrade completely and to be disposed of by largely normal metabolic processes. Since, furthermore, it is possible to control the degradation rate of the implanted polymer by chemical cross-linking or other chemical modifications, the designer is offered the opportunity to control the lifetime of the implant.

A disadvantage of proteins on biomaterials is their frequently significant immunogenicity, which, of course, derives precisely from their similarity to naturally occurring substances. The immunological reaction of the host to the implant is directed against selected sites (antigenic determinants) in the protein molecule. This reaction can be mediated by molecules in solution in body fluids (immunoglobulins). A single such molecule (antibody) binds to single or multiple determinants on an antigen. The immunological reaction can also be mediated by molecules which are held tightly to the surface of immune cells (lymphocytes). The implant is eventually degraded. The reaction can be virtually eliminated provided that the antigenic determinants have been previously modified chemically. The immunogenicity of polysaccharides is typically

TABLE 1 General Properties of Certain Natural Polymers



Physiological function

A. Proteins

B. Polysaccharides

C. Polynucleotides


Keratin Collagen






Cellulose (cotton) Amylose Dextran Chitin


Deoxyribonucleic acids (DNA) Ribonucleic acids (RNA)

Synthesized by arthropods Hair

Connective tissues (tendon, skin, etc.) Partly amorphous collagen Blood

Neck ligament



Plants Plants

Synthesized by bacteria Insects, crustaceans Connective tissues

Cell nucleus Cell nucleus

Protective cocoon Thermal insulation Mechanical support

(Industrial product) Blood clotting Mechanical support Contraction, motility Contraction, motility

Mechanical support

Energy reservoir

Matrix for growth of organism

Provides shape and form

Contributes to mechanical support

Direct protein biosynthesis Direct protein biosynthesis far lower than that of proteins. The collagens are generally weak immunogens relative to the majority of proteins.

Another disadvantage of proteins as biomaterials derives from the fact that these polymers typically decompose or undergo pyrolytic modification at temperatures below their melting point, thereby precluding the convenience of high-temperature thermoplastic processing methods, such as melt extrusion, during the manufacturing of the implant. However, processes for extruding these temperature-sensitive polymers at room temperature have been developed. Another serious disadvantage is the natural variability in structure of macromolecular substances which are derived from animal sources. Each of these polymers appears as a chemically distinct entity not only from one species to another (species specificity) but also from one dssue to the next (tissue specificity). This testimonial to the elegance of the naturally evolved design of the mammalian body becomes a problem for the manufacturer of implants, which are typically required to adhere to rigid specifications from one batch to the next. Consequently, relatively stringent control methods must be used for the raw material.

Most of the natural polymers in use as biomaterials today are constituents of the extracellular matrix (ECM) of connective tissues such as tendons, ligaments, skin, blood vessels, and bone. These tissues are deformable, fiber-reinforced composite materials of superior architectural sophistication whose main function in the adult animal appears to be the maintenance of organ shape as well as of the organism itself. In the relatively crude description of these tissues as if they were man-made composites, collagen and elastin fibers mechanically reinforce a "matrix" that primarily consists of protein—polysaccharides (proteoglycans) highly swollen in water. Extensive chemical bonding connects these macromolecules to each other, rendering these tissues insoluble and, therefore, impossible to characterize with dilute solution methods unless the tissue is chemi cally and physically degraded. In the latter case, the solubilized components are subsequently extracted and characterized by biochemical and physicochemical method. Of the various components of extracellular materials which have been used to fashion biomaterials, collagen is the one most frequently used.

Almost inevitably, the physicochemical processes used to extract the individual polymer from tissues, as well as subsequent deliberate modifications, alter the native structure, sometimes significantly. The following description emphasizes the features of the naturally occurring, or native, macromolecular structures. Certain modified forms of these polymers are also described.

structure of native collagen

Structural order in collagen, as in other proteins, occurs at several discrete levels of the structural hierarchy. The collagen in the tissues of a vertebrate occurs in at least ten different forms, each of these being dominant in a specific tissue. Structurally, these collagens share the characteristic triple helix, and variations among them are restricted to the length of the nonhelical fraction, the length of the helix itself, and the number and nature of carbohydrate attachments on the triple helix. The collagen in skin, tendon, and bone is mostly type I collagen. Type II collagen is predominant in cartilage, while type III collagen is a major constituent of the blood vessel wall as well as being a minor contaminant of type I collagen in skin. In contrast to these collagens, all of which form fibrils with the distinct collagen periodicity, type IV collagen, a constituent of the basement membrane which separates epithelial tissues from mesodermal tissues is largely nonhelical and does not form fibrils. We follow here the nomenclature which was proposed by W. Kauzmann (1959) to describe in a general way the

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