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to a large extent resorbed by 38 weeks. The larger flame-sprayed implant in animal 3A14 (fig. 26) prevents total encapsulation by the callus (13 weeks) but also appears to form a contiguous structure with the remodeled callus (37 weeks). Animal 2B3 sustained complete paralysis of the right leg during the implantation, and shows a minimal amount of reparative activity at 16 weeks. At 36 weeks little evidence of fracture healing was seen (fig. 27). The animals were sacrificed at periods between 23 and 50 weeks. All major long bones of the lower extremities were removed, the Schneider nails withdrawn, and the bones tested mechanically to failure in a rapid loading torsion tester [34]. The bones were kept physiologically moist during sacrifice and test. After the torsional tests were completed, the bones were reassembled, embedded in a clear epoxy potting compound, and sectioned. The geometry of these sections was transmitted to the SCADS computer program [35, 36] which was used to determine the torsional stresses generated both at the fracture sites and at the implant surfaces.

TABLE 3. Mechanical properties of monkey femurs In all cases, the right femur contains the implant.

torques required to fracture the left and right bones respectively. Pairs of tibias were used as controls for this portion of the experiment. The group of tibias exhibits a mean difference of 0.010, indicating that there is no left-right bias, and a standard deviation of 0.07 [37]. The group of femurs containing a bulk bioglass-ceramic implant showed an average difference of 0.29; this difference is quite significant, in that it establishes that the structural integrity of the femurs which were operated on has not recovered fully by approximately 40 weeks. The femurs containing the flame-sprayed implants failed at extremely low torques, resulting in an average difference of 0.66. The torques indicate that the flame-sprayed implant system probably would not be able to sustain severe physiological torques. However, the monkeys used the legs with the flame-spray coated prostheses without difficulty or any noticeable physiological deficit throughout the post-operative period. Close examination of the failure surfaces, which were always near the implant bone interface, showed that the glass coating was still adhering to the bone but that it was torn off the metal substrate during the mechanical testing.

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Figure 28 shows the results of the stress analyses of the various bones involved. The upper histogram shows the distribution of fracture stresses of normal bones, i.e., left femurs and tibias in a bold outline. The average stress required to fracture a normal monkey bone is about 1560 kgf/cm2 with a standard deviation of 300 kgf/cm2. No difference was noted in the strength of bone from left femurs and from tibias. Stresses from three groups of bone implant systems are summarized in figure 28: bones with flame sprayed stainless steel implants, bones with bioglass-ceramic implants which broke during the mechanical testing, and bones with bioglassceramic implants which fracture through the bone rather than through the implant. The average stress at which the bones containing flame sprayed implants failed was 203 kgf/cm2. As noted before this is really the stress required to fracture the glass from the metal substrate (see fig. 29). Two

[graphic]

FIGURE 29. Fractured end of the flame spray coated segmental implant showing the exposed metal surface fracture site (grey areas).

bulk bioglass-ceramic implants failed through the implant rather than through the bone (fig. 30), and these two are shown as crosshatched blocks in figure 28. This low strength shows that at the present state of the art the bulk material is not of adequate strength to reliably support physiological loads to failure. Five femurs containing bulk bioglassceramic implants failed away from the bone implant interface. The lower histogram of figure 28 shows the distribution of the fracture stresses calculated in those bones, with the corresponding region in the upper histogram denoting the interfacial stresses

FIGURE 30. Failure of bone bioglass-ceramic implant test specimen (2B18) failing within the implant.

Note the implant bone interface withstands the fracture stress applied.

developed at the surface between the bioglassceramic and the newly healed bone. The average stresses in these two distributions are very close to each other, and are not significantly different from each other.

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13. Discussion

The structural strength of the monkey femurs was compromised considerably by the presence of the implant. During the torsion tests, fracture occurred at various sites, but appeared to be limited by either the strength of the implant material or the strength of the healing bone near the implant. The fact that the fracture stresses of the right femurs containing bioglass-ceramic implants was considerably lower than the average fracture stress for normal monkey bones was probably due to the fact that the Schneider nails used for fixation were not removed from the animal until just prior to test. This means that a substantial portion of the bending loads to which the bone-ceramic system was subjected was transmitted across the implant site by the nail itself. Thus the bone was protected from full physiologic stress and is therefore probably weaker than normal monkey bone. It should be kept in mind, however, that normal axial compression forces were able to act on the implant-bone interface. This axial compression is probably responsible for the in-vivo implant failures. The distribution of the interfacial stresses calculated for the bulk bioglass-ceramic implants indicates that the strength of the interface is nearly as great as that of the newly healing bone in the same vicinity. The data in figure 28 also indicates. that the strength of the interface is at least 75 percent of the strength of normal monkey bone. Due to the fact that failure did not occur preferentially at the interfaces in our study, it must be kept in mind that these figures represent only a lower limit on the strength of the bond developed between the bone and the bioglass-ceramic.

The authors gratefully acknowledge the able assistance of H. Parkhurst, A. W. Smith, and J. McVey of the University of Florida; S. W. Freiman and E. J. Onesto of IITRI, the financial support of U.S. Army Medical R & D Command Contract No. DADA 17-70-C-001 and cooperation of the Gainesville, Florida Veterans Administration Hospital.

14. References

[1] Hench, L. L., Ceramics, glasses and composites in medicine, Med. Instr. 1, 136–144 (1973).

[2] Boutin, P., Direct attachment of ceramic joints, Symposium on Materials for Reconstructive Surgery, (Clemson University, April 1974).

[3] Griss, P., Heimke, G., von Andrian-Werburg, H., Krempien, B., Reipa, S., Hartung, J., and Lauterbach, J., Morphological and biomedical aspects of Al2O3 ceramic joint replacement; experimental results and design considerations for human endoprostheses, Symposium on Materials for Reconstructive Surgery, (Clemson University, April 1974). [4] Griss, P., von Andrian-Werburg, H., Krempien, B., and Heimke, G., Biological activity and histocompatibility of dense Al2O3/MgO ceramic implants in rats, J. Biomed. Res. Symp., No. 4, pp. 453-462, (Interscience, New York, N.Y. 1973).

[5] Griss, P., Krempien, B., von Andrian-Werburg, H., Heimke, G., and Fleiner, R., Experimentelle Utersuchung zur Geivebsuertraglichkeit oxidkeramischer (Al2O3) Abriebteilchen, Arch. für orthopädische und Unfall-Chirurgie 76, 270-279 (1973).

[6] Englehardt, A., Salzer, M., Zeibig, A., and Locke, H., Experiences with Al2O3 implantation in humans to bridge resection defects, Symposium on Materials for Reconstructive Surgery, (Clemson University, April 1974). [7] Klawitter, J. J., and Hulbert, S. F., Applications of porous ceramics for the attachment of load bearing internal orthopedic applications, J. Biomed. Res. Symp., No. 2, Part 1, pp. 161–229 (Interscience, New York, N.Y. 1972). [8] Hulbert, S. F., Cooke, F. W., Klawitter, J. J., Leonard, R. B., Sauer, B. W., Moyle, D. D., and Skinner, H. B., Attachment of prostheses to the musculoskeletal system by tissue ingrowth and mechanical interlocking, Biomed. Res. Symp. No. 4, pp. 1-23, (Interscience, New York, N.Y., 1973).

[9] Hulbert, S. F., Young, F. A., Mathews, R. S., Klawitter, J. J., Talbert, C. P., and Stelling, F. H., Potential of ceramic materials as permanently implanted skeletal prostheses, J. Biomed. Mat. Res. 4, 433-455 (1970). [10] Hulbert, S. F., Klawitter, J. J., and Leonard, R. B., Compatibility of bioceramics with the physiological environment, Ceramics in Severe Environments, W. W. Kriegel and H. Palmour, III (Eds.), pp. 417-434 (Plenum Press, New York, N.Y., 1971).

[11] Piotrowski, G., Hench, L. L., Allen, W. C., and Miller, G., Mechanical studies of the bone-bioglass interfacial bond, Symposium on Materials for Reconstructive Surgery, (Clemson University, April 1974).

[12] Graves, G. A., The influence of compositional variations on bone ingrowth of implanted porous calcium aluminate ceramics, Symposium on Materials for Reconstructive Surgery, (Clemson University, April 1974).

[13] Graves, G. A., Hentrich, R. L., Stein, H. G., and Bajpai, P. K., Resorbable ceramic implants, J. Biomed, Res. Symp. No. 2, Part 1, pp. 91-115 (Interscience, New York, N.Y. 1972).

[14] Bhaskar, S. N., Brady, G. M., Getter, L., Grower, M. F.,

and Driskell, T. D., Biodegradable ceramic implants in

bone, Oral Surg. 32, 336 (1971).

[15] Bhaskar, S. N., Cutright, D. E., Knapp, M. J., Beasley, J. D., Perez, B., and Driskell, T. D. Tissue reaction to intrabony ceramic implants, Oral Surg. 31, 282 (1971).

[16] Schnittgrund, G. D., Kenner, G. H., and Brown, S. D., In vivo and in vitro changes in strength of orthopedic calcium aluminates, J. Biomed. Res. Symp. No. 4, pp. 435-452 (Interscience, New York, N.Y., 1973). [17] Kenner, G. H., Brown, S. D., Pasco, W. D., and Lovell, J. E., Bull. Amer. Ceramic Soc. 52, 432 (1973).

[18] Frakes, J. T., Brown, S. D., and Kenner, G. H., Bull, Amer. Ceramic Soc. in press, (1973).

[19] Jecman, R. M., Eggerding, C. L., Brown, S. D., and Schnittgrund, G. D., J. Biomed. Mat. Res. 7,369 (1973).

[20] Hench, L. L., Splinter, R. J., Allen, W. C., and Greenlee, T. K., Mechanisms of interfacial bonding between ceramics and bone, J. Biomed. Res. Symp. No. 2, Part 1, pp. 117-143 (Interscience, New York, N.Y., 1972).

[21] Hench, L. L., and Paschall, H. A., Direct chemical bond of bioactive glass-ceramic materials to bone and muscle, J. Biomed. Res. Symp. No. 4, pp. 25-42 (Interscience, New York, N.Y., 1973).

[22] Clark, A. E., Hench, L. L., and Paschall, H. A., The influence of surface chemistry on implant interface hisotlogy: A theoretical basis for implant materials selection, Symposium on Materials for Reconstructive Surgery, (Clemson University, April 1974).

[23] Hench, L. L., and Paschall, H. A., Histo-chemical responses at a biomaterials interface, J. Biomed. Res. Symp. on Prostheses and Tissue: The Interface Problem, (Clemson University, 1973).

[24] Beckham, C. A., Greenlee, T. K., Jr., and Crebo, A. R., Bone formation at a ceramic implant interface, Calc. Tis. Res. 8, No. 2, 165 (1971).

[25] Greenlee, T. K., Jr., Beckham, C. A., Crebo, A. R., and Malmborg, J. C., Tissue responses at the interface of a ceramic, J. Biomed. Mat. Res. 6, 244 (1972).

[26] Clark, A. E., and Hench, L. L., Effect of P5+, B3+, and Fon corrosion of invert soda-lime-silica glasses, Submitted to J. Amer. Ceramic Soc. (1974).

[27] Clark, A. E., Pantano, C. G., and Hench, L. L., Auger spectroscopic analysis of invert glass corrosion films, Submitted to J. Amer. Ceramic Soc. (1974).

[28]

[29]

Carlisle, E. M., Silicon- a possible factor in bone calcification, Science 167, 279-280 (1970).

Housefield, L. G., Mechanical property control of a bioglassceramic system, Masters Thesis, University of Florida (1972).

[30] Bliton, J. L., Ceramic coatings for cementitious and metallic surfaces, IITRI Final Report, Project No. B8009, (International Lead Zinc Research Organization, 1964).

[31] Oel, H. L., The relationship between free energy and kinetics in sintering processes, Kinetics of High-Temperature Processes, W. D. Kingery (Ed.), (John Wiley and Sons, Inc., New York, N.Y., 1959).

[32] Kingery, W. D., Introduction to Ceramics, pp. 628-647 (John Wiley and Sons, Inc., New York, N.Y., 1960).

[33] Hench, L. L., Onesto, E. J., and Freiman, S. W., Development of a bioglass flame spray coated hip prosthesis. U.S. Army Medical Research and Development Command Contract No. DADA 17-70-C-0001, Report No. 3, 180 pp. [34] Burstein, A. H., and Frankel, V. H., A standard test for laboratory animal bone, J. Biomech. 4, 155 (1971). [35] Piotrowski, G., and Wilcox, G. A., The STRESS program: a computer program for the analysis of stresses in long bones, J. Biomech. 4, 497 (1971).

[36] Piotrowski, G., and Kellman, G. I., A stress calculator for arbitrarily drawn sections - the S.C.A.D.S. computer program, U.S. Army Medical Research and Development Command Contract No. DADA 17-70-C-0001, Report No. 4, 126 pp. (1973).

[37] Miller, G., and Piotrowski, G., Variability of the torsional strength of paired bones, J. Biomech., 7, 247 (1974).

NATIONAL BUREAU OF STANDARDS SPECIAL PUBLICATION 415,

Biomaterials, Proceedings of a Symposium held in conjunction with the Ninth Annual Meeting of the
Association for the Advancement of Medical Instrumentation, New Orleans, La., April 19–20, 1974
(Issued May 1975).

Soft Tissue Response to a Series of Dense Ceramic
Materials and Two Clinically Used Biomaterials

W. C. Richardson, Jr.

Tulane University, New Orleans, La. 70112

S. F. Hulbert

Medical Center, University of Alabama, Birmingham, Ala. 35294

and

J. J. Klawitter, and B. W. Sauer

Clemson University, Clemson, S.C. 29631

Disc-shaped implants of spinel, alumina, mullite, zircon, a cast Co-Cr-Mo alloy, and ultra-high molecular weight polyethylene (UHMWPE) were implanted in the parasipinalis musle of twelve adult, male, White New Zealand rabbits. Prior to implantation the implants were characterized with respect to size and shape, weight and surface roughness. After periods of 1 month, 2 months, and 4 months the rabbits were sacrificed and the tissue specimens were retrieved with the implants still intact. Histological examination of the tissues surrounding the implants along with changes in the size and shape, weight, and surface roughness of the implants were used as criteria for evaluating these materials for implant purposes.

No surface degradation of any of the materials was detected using scanning electron microscopy. Fibrous tissue seems to adhere to the UHMWPE implants more than any other material used in this study. Large amounts of fibrous tissue were also found to adhere to the cast Co-Cr-Mo alloy implants. The histological results indicated that within the limits of this investigation, the biocompatibility of the ceramic materials used in this study compares favorably with the clinically used cast Co-Cr-Mo alloy implants and the UHMWPE implants.

Key words: Biocompatibility, ceramic implants, histological evaluation; implant characterization.

1. Introduction

Ceramic materials are beginning to find a place in the field of biomaterials as new demands are made for materials that can withstand the highly corrosive environment of the human body [1-26]. The physical and chemical properties of ceramics have recently led researchers to consider these materials for use in the replacement of bone [1, 3, 4, 6, 9, 10, 12-16, 19-26] and teeth [17-19].

Ceramic materials have certain advantages and disadvantages as compared to other materials for implantation. The major advantage of ceramics is that they are highly chemically inert in the physiological environment. They are, however, characteristically brittle, notch sensitive materials. Thus, a careful consideration of their mechanical properties will be essential to their applications as biomaterials.

Figures in brackets indicate the literature references at the end of the paper.

Fixation of orthopedic appliances to bone is a definite problem to the orthopedic surgeon. Some investigators have approached this problem by developing a mechanical interlocking between the implant and the tissue [3, 6, 10, 12-14, 17, 19, 20, 22, 24, 26]. The mechanical interlocking was achieved by tissue growth into the interconnecting pores of porous ceramic materials. Klawitter [10] reported that a minimum pore size of approximately 100 μm is necessary for bone ingrowth.

Hench et al. [21], reported achieving direct chemical bonding of a glass-ceramic material with bone and soft tissues. They accomplished this through the development of a series of surface-active glasses and glass-ceramics.

2. Compatibility Testing

One of the problems of conducting a biocompatibility experiment is that of methodology. Thus far there appears to be no universally accepted method for evaluating tissue compatibility [27–42].

There are two main techniques which have been used to evaluate candidate materials for implantation, in vitro and in vivo studies. In vitro studies incorporate the use of tissue culture or simulated body fluids to test the biocompatibility of candidate materials and they have advantages in that they are considerably less expensive. While in vitro studies serve as an excellent tool for screening materials, they do not provide conclusive information because the total dynamic physiological environment is not simulated [36, 39, 40].

In vivo biocompatibility studies, most generally, consist of implanting materials of interest into research animals for varying periods of time and subsequently performing a histological evaluation of the adjacent tissues.

There are some inherent problems associated with determining compatibility of biomaterials [44]. When implants are placed into the body, the material is subjected to a barage of rejection and healing mechanisms. As the body reacts to the foreign material, certain histological processes take place which are observable in the tissues surrounding the implant as well as in special organs that are far removed from the implant. The combined activity of the many cell types, tissues, and organ systems produce such a complex sequence of interactions that it is often difficult to explain certain histological observations that occur when an implant is placed in the body [44]. It is essential for the researchers in these types of studies to have a basic understanding of the nature and processes of foreign-body reactions and their characteristic signs. A paper by Hulbert et al. [44], provides a discussion of phagocytosis, the immune response, abscesses, neoplasms, poisons, inflammation and normal wound healing and relates these phenomena to their importance in evaluating the biocompatibility of implant materials.

Many investigators attempt to monitor the biocompatibility of an implant by designing a device of the type desired and to test it in the exact location of the intended application [7, 17, 18, 24, 25, 37, 45]. In an attempt to reduce the number of experimental variables, other researchers have implanted sample materials in various shapes and forms [5, 6, 9–12, 15, 27-30, 33, 34]. Wood et al. [29], reported that an increased tissue response which typically occurs near the ends of rod-shaped implants does not occur around disc-shaped implants.

Autian [38] mentioned that a material implanted in the body may initiate a response in two general ways: (1) the effects of the body on the material and (2) the effects of the material on the body. In view of this fact, it seems that a properly designed compatibility experiment would seek to not only investigate the histological reactions, but would also attempt to monitor changes which may occur

to the material. Indeed, material characterization is an extremely important but often neglected portion of biocompatibility studies.

3. Materials and Methods

3.1. Scope of the Research

Disc-shaped implants of spinel 2 (MgO Al2O3), alumina 2 (Al2O3), mullite 2 (3Al2O3 2SiO2), zircon2 (ZrSiO4), a cast Co-Cr-Mo alloy, and ultra-high molecular weight polyethylene, were implanted in the paraspinalis muscle of twelve adult, male, White New Zealand rabbits. Prior to implantation the implants were characterized with respect to size and shape, weight and surface roughness. After periods of 1 month, 2 months, and 4 months the rabbits were sacrificed and the tissue specimens were retrieved with the implants still intact. Histological examination of the tissues surrounding the implants along with changes in the size and shape, weight and surface roughness of the implants were used as criteria for evaluating these materials for implant purposes.

3.2. Materials Selection

Spinel2 (MgO Al2O3), alumina 2 (Al2O3), mullite 2 (3Al2O3 2SiO2), and zircon2 (ZrSiO4), were chosen for this experiment because they represent a group of ceramic materials which are often selected for use in severe environments and which were available from commercial suppliers.5

an

In order to facilitate quality comparisons of these ceramic materials with clinically used biomaterials, this experiment included the implantation of ultra-high molecular molecular weight polyethylene and (UHMWPE) 4 a cast cobalt-chromiummolybdenum alloy (meeting the ASTM Standard Specification for Surgical Implants, designation number F75-67) as experimental controls.

3.3. Materials Characterization

Each implant was characterized with respect to size and shape, surface roughness, and weight before and after implantation. Scanning electron microscopy was also used before and after implantation. Chemical analysis, x-ray diffraction, and reflected light microscopy were used to characterize the materials before implantation.

2 Supplied by American Lava Corporation, Chattanooga, Tenn.

3 Zimaloy, Zimmer Manufacturing Company, Warsaw, Ind.

RCH 1000, Hoechst, Supplied by Zimmer Manufacturing Co., Warsaw, Ind. Certain commercial materials and instruments may be identified in this publication in order to adequately specify the experimental procedure. In no case does such identification imply recommendation or endorsement by the National Bureau of Standards, nor does it imply that the equipment or instruments identified are necessarily the best available for the purpose.

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