GAIL K. SMITH
Properties Required of Fixation Materials
Specific Implant Materials
Titanium Based Alloys
Materials Performance: Materials-Associated
The late l9th century ushered in the concept of aseptic surgery, and with refinement of this principle came the possibility of implanting foreign materials into the body with an acceptably low risk of rejection due to infection. However, at this time there existed no suitable material to withstand the challenges posed by the biologic environment, particularly in regard to strength and corrosion resistance. Developments in metal refining and processing in the first half of the 20th century, stimulated largely by wartime needs, led to the production of improved materials that were rapidly, although empirically, adapted by surgeons for use in fracture fixahon. The materials currently popular are those that over the years have performed "acceptably" in clinical situations, and only recently (within the past 20 years) has the introduction of the field of biomaterials research had an influence on subsequent technologic advances in this regard.
The science of biomaterials, by liberal definition, encompasses all materials of biologic composition or for biologic application. For purposes of this chapter, however, the designation "orthopaedic biomaterials" will be synonymous with "materials used for internal fracture fixation." Excellent reviews tracing the history of biomaterials development and describing state-of-the-art technology have been published and should be consulted for more in-depth information on this subject,(2,5,10,11,13,15,17)
The ultimate aim of biomaterials as applied to fracture fixation is to restore the structural integrity of the damaged bone. The attainment of this objective is dependent upon a complex interplay of materials properties, device design, and physiologic requirements. Important are considerations such as the site and type of fracture, the possible operative approaches, the rapidity of bone healing, and the desired or feasible program of postoperative care. Specifically, the materials selection process must incorporate the chemical and mechanical demands of the biologic environment to achieve a functional outcome. In addition to in vivo performance, however, a material must also satisfy the practical manufacturing and marketing requirements including reasonable cost, ability to be fabricated, and ease of surgical application. The totality of these factors applied to materials selection and device design has resulted in an array of commercially available fracture-fixation instrumentation, which if applied and maintained according to established surgical principles will yield excellent functional results. It is the surgeon's responsibility to understand both the materials and structural limitations of these devices and the principles of application to minimize performance failure. The biomaterials scientist, on the other hand, is expected to develop implants of optimum design that take fullest advantage of the materials available.
For purposes of organization this chapter is divided into three sections: the general requirements of materials for internal fixation, the specific implant materials currently used, and materials performance including materials-associated complications.
PROPERTIES REQUIRED OF FIXATION MATERIALS
The requirements of materials for internal fixation fall into three categories: clinical, manufacturing, and economic, in descending order of importance. The obvious clinical requirements are that the material have suitable mechanical properties to fulfill its function of fixation and maintenance of fracture reduction. Additionally, the material must not degrade in response to the corrosive conditions of the biologic environment, thereby impairing mechanical properties or releasing potentially harmful degradation by-products locally and systemically.
Manufacturing requirements dictate that the material have properties that permit fabrication in the optimum design configuration, while cost requirements, although not as critical in human orthopaedic practice, must be considered in veterinary orthopaedics to compete with other, albeit less effective, forms of fracture fixation or treatment, not excluding euthanasia.
By convention the mechanical properties of a material are described in terms of the deformation or strain produced by an applied stress. Such behavior can be plotted on a stress-strain diagram (Fig. 13-1).(3) Normal stress (sigma) is defined as the applied force divided by the cross-sectional area over which it is acting, while the resultant strain is the change in length per unit original length. This designation facilitates the assessment of inherent materials properties and excludes structural aspects of the test specimen. Materials are commonly evaluated in tension; however, similar although more complicated formulas exist for materials testing in compression, bending, shear, and torsion. These modes of loading represent standard materials testing methodology and yield several parameters important to the quantitative characterization of the mechanical behavior of materials.
FIG. 13-1 Idealized stress-strain curve summarizes the static and low-cycle properties of materials Stress, the vertical coordinate, is the applied load divided by the cross-sectional area of the material. strain, the horizontal coordinate, is the resulting elongation per unit length Other parameters include sigma y, the yield stress, the greatest stress that can be applied without permanent deformation after removal of stress; sigma 0.2%. the 0.2% stress that causes a permanent strain of 0.2% after removal; sigma u, the ultimate stress that produces material tailure by fracture; epsilon u, the ultimate strain or strain to failure; M, the modulus, the ratio of stress divided by strain in the linear portion of the curve and a measure of the instrinsic stiffness of the material An additional parameter of interest, shown as the cross-hatched area, is the work to fracture per unit volume of the material, or the amount of energy that must be absorbed per unit volume before tracture takes place, a measure of toughness. (Black J Biomaterials for internal fixation In Heppenstall B(ed): Fracture Treatment and Healing, p 113 Philadelphia, WB Saunders, 1980)
Yield stress, cry (Fig. 13-1), is the maximum stress that can be applied without producing permanent deformation after removal of the stress. In this context it marks the end of the elastic (linear) portion of the stress- strain curve, sometimes termed the elastic limit.
The "proof stress" or 0.2% offset stress corresponds to the stress that produces a 0.2% permanent strain upon unloading the specimen and is typically used as a comparative parameter for materials that exhibit a poorly defined yield point.
The ultimate stress, sigma sigma u (Fig. 13-1), is the stress associated with complete mechanical failure of the test specimen. In brittle materials cru will occur within or just after the elastic portion of the stress-strain curve, whereas ductile materials exhibit a long phase of plastic deformation prior to ultimate failure. The ultimate strain is the fractional deformation at failure and is occasionally called the ductility.
The modulus of elasticity, M, is the ratio of applied stress to the resultant strain in the linear (elastic) portion of the stress-strain curve and in tensile testing is designated the "Young's modulus, 'E'". This parameter corresponds to the stiffness of a material and owing to the phenomenon of "stress protection" discussed in Chapter 12 is an important factor to consider in the design and application of rigid internal fixation devices.
Similarly the area under the stress-strain curve (cross-hatched area, Fig. 13-1) is a biologically important attribute of structural materials and reflects the toughness or the energy absorbed (work required) to produce failure. Brittle materials, although often exhibiting high ultimate stress and high modulus, typically have a smaller area under the curve owing to low breaking strain and therefore are not as tough, for example, as a "strong" ductile material having a moderate ultimate stress but large ultimate strain. For structural implant purposes the latter is preferred. The properties of toughness and ductility act to prevent catastrophic failure as a result of accidental overload or localized strain. In addition, these properties facilitate plate "contouring" for fracture fixation without severely jeopardizing plate strength.
The mechanical property of hardness refers to a material's resistance to indentation either by a ball (Brinell test) or by a pyramidal diamond (Vickers test). Depth and area of indentation are related to applied load, with high values reflecting hard materials and low values, soft materials. The measurement of hardness is important, particularly when combining various materials for bearing or wear conditions. Fracture-fixation devices, however, are specified to be manufactured of uniform alloy composition, phase, and surface preparation, and therefore large variations in hardness are not to be expected.
The required mechanical properties of implant materials discussed thus far are determined by static or low-cycle short-term in vitro materials-testing methodology. A fracture-fixation device, however, typically remains in service for a period of months or years and must endure repeated loading under corrosive conditions. The mechanical properties of "fatigue," "creep," and "stress relaxation" reflect time-dependent materials' behavior and are important mechanical parameters in the selection process.
Fatigue refers to a mode of failure that results from repeated stress at magnitudes lower than that required to cause failure in a single application (ultimate stress). Implant materials necessarily must have a high degree of fatigue resistance to perform over the long term; it is estimated that fixation hardware will experience 2 x 106 cycles per year. Fatigue behavior for a given material is best illustrated by an S-N curve, Fig. 13-2, which plots stress, S, as a function of cycles to failure, N.(3) At low cycles to failure the corresponding peak stress approaches the static ultimate stress, sigma u, of the material. As N increases, however, the peak stress to failure diminishes until for many materials a "fatigue" or "endurance" limit is reached. At stresses below this level fatigue failure will not occur regardless of the number of loading cycles. For some metals and alloys such as steels the observed endurance limit is approximately 0.4 times the ultimate tensile strength, and therefore to prolong implant life indefinitely, imposed loads should not exceed this figure. Materials not having an endurance limit must be structurally designed such that cyclic loading conditions over the expected life span of the device always fall under the S-N curve. For human structural implants the maximum criterion is that a device must be capable of withstanding a maximum cyclic stress that will not lead to fatigue failure in 5 x 107 cycles or approximately 25 years of norrnal use. Veterinary patients, because they have much shorter life spans and smaller mass, clearly do not require such rigorous low-stress, high-cycle criteria. A margin of safety, however, must be provided for high-stress, low- cycle conditions stemming from the inevitable uncertainty of controlling an animal's activity postoperatively.
Certain materials may undergo progressive deformation with time under constant stress. This mode of deformation and failure is called creep and relates to the flow or viscoelastic characteristics of a material. Non-crystalline materials such as polymers, consisting of long-chain molecules having weak Van der Waals bonds between chains, are particularly prone to this time-dependent form of deformation. The process in polymers is also more temperature-sensitive than in metals and deformations may occur rapidly on exposure to quite low stresses. Creep phenomena may be minimized in the implant design process by using metal alloys having a high melting point and face-centered cubic crystalline structure to inhibit dislocation movement. The presence of solutes or stable precipitates also will restrict the movement of dislocations and resist creep. Plastics can be made more resistant to creep by the addition of fillers or other reinforcing materials that raise the viscosity
Stress relaxation is a time-dependent, viscoelastic process similar to creep, but it refers to the decay of applied stress under conditions of constant strain (in contrast to creep, which is a time-dependent strain under conditions of constant stress). Interestingly, musculo-skeletal tissues including muscle, tendons, and bone all exhibit viscoelastic properties. An excellent biologic example of stress relaxation is the time-dependent reduction in applied interfragmentary compression at a fracture site that is fixed and compressed with a dynamic compression plate. (See Chapter 12.) Both creep and stress relaxation are undesirable properhes of orthopaedic bio-materials because they release frictional forces necessary to maintain rigid internal fixation.
FIG. 13-2 Idealized S-N curve for metals The higher the peak stress produced in a given cycle of loading and unloading, the fewer cycles can be sustained before failure The combination of stress (s) and number of cycles (N) that fall above this curve predicts failure of the material by fracture. The region to the left (high 8, low N) represents failure at stresses above the yield stress after a small number of cycles; that to the right (low S, high N) reflects repeated deformation, a result primarily of stress below the yield stress. Many materials display an endurance limit, a level of stress below which fatigue failure will not occur, regardless of how many cycles of loading and unloading are experienced. (Black J: Biomaterials for intemal fixation. In Heppenstall B (ed): Fracture Treatment and Healing, p 113. Philadelphia, WB Saunders, l980)
The most important nonmechanical requirement of an orthopaedic biomaterial is "inertness." Ideally an implant material should not degrade at all. In reality, however, such a state is unachievable, and therefore a relative degree of implant degradation is considered acceptable. The degradation process, however, must not impair significantly the mechanical strength of the device nor allow the release locally or systemically of by-products that might evoke a deleterious biologic response.
The corrosion of metals in biologic fluids is an electrochemical reaction that results in the release of metal ions into the surrounding aqueous electrolyte. This dissolution reaction is coupled with a corresponding reduction reaction of constituents in the aqueous environment to maintain charge neutrality. The alloys currently used as orthopaedic biomaterials are protected from accelerated corrosion rate by a passivating oxide layer that acts like an electrical resistor to retard the anodic dissolution of metal cations. It therefore follows that under optimum circumstances of complete passivation, all implant alloys have a finite, albeit slow, uniform corrosion rate in vivo but that damage to this passivating layer, such as by frethng or wear, may produce conditions conducive to accelerated focal corrosion and failure.
Factors that predispose to localized corrosive attack include metal transfer of materials from surgical tool to implant, differential ¡2 and H+ concentration of electrolytes over an implant, fretting corrosion or surface abrasion that causes the destruction of the protective oxide layer, mixing of various metals or various metallurgical states leading to galvanic corrosion, surface irregularities such as crevices or pits (e.g., the interface of screw and plate) that cause crevice corrosion, and other mechanochemical phenomena such as stress-corrosion cracking and corrosion fatigue. Although the design process has minimized the effects of these corrosive factors on implant performance, they cannot be eliminated completely. Thus metallic devices for fracture fixation should be considered temporary implants and removed whenever the clinical course, the patient's condition, and the owner's compliance make such removal reasonable in the surgeon's judgment.
Degradation of nonmetallic implant materials is more difflcult to assess than that of metals and hence the selection process for suitable nonmetallic materials have relied heavily on industrial experience. Of importance in the degradation process of polymers are the types of chemical bonds present, steric hindrance and electronegativity effects produced by atoms in close proximity with these chemical bonds, and supermolecular structure (e.g., a crystalline structure is generally more inert than an amorphous one). (1,11) Also important is a knowledge of the low-molecular-weight mobile moieties and their cytotoxicity, since such molecules constitute a diffusible fraction consisting of polymeric chain degradation by- products or stabilizing additives or plasticizers. The dearth of quantitative data available on the rate of dissolution of nonmetallic materials complicates the selection process further. Failures that have been observed in surgically implanted polymers appear to relate to the susceptibility of nonmetallic materials to various forms of mechanochemical deterioration, such as stress-solvent crazing, friction, and wear and fatigue.
Other requirements of metallic and nonmetallic orthopaedic biomaterials include manufacturing, fabricability, and cost. However, a discussion of these is beyond the scope of this chapter. For more detailed information, the interested reader is referred to the excellent texts by Mears" and Williams and Roaf.(17)
SPECIFIC IMPLANT MATERIALS
Historically, metals have been selected for structural implant purposes owing to their superior strength. With advances in refining technology, many metals and alloy systems have been developed with a high degree of corrosion resistance. The majority of metals and alloys, however, are not suitable for biologic implantation because of the poor tolerance of the body to even minute concentrations of dissolution products; that is few metals have suitable biologic compatibility as determined empirically and experimentally. Those currently considered acceptable are materials based on iron, cobalt, nickel, titanium, tantalum, zirconium, silver, gold, and the noble metals. Of these, tantalum and the noble metals (including gold and silver) do not possess adequate mechanical properties for structural implant purposes, while zirconium is too expensive. Of the remaining base metals, six alloys have been developed for use in orthopaedic surgery. Their chemical composition is given in Table 13-1, including references to relevant specifications and a short list of manufacturer's trade names.(3,9)
These alloy systems have been selected for their optimum combination of properties, including high elastic modulus and ultimate strength, ductile behavior for loads exceeding the yield stress (thus reducing the risk of catastrophic brittle failure), fabricability, and corrosion resistance. The six alloys shown in Table 13-1 may be grouped according to composition into three alloy systems: the austenitic stainless steels (iron, chromium, molybdenum) the cobalt-chromium alloys, and the titanium alloys.
The primary stainless steel alloy presently recommended for device manufacture is the American Iron and Steel Institute (AISI) type 316L (Table 13-1). The exact composition may vary slightly relative to the casting or forging variant; however, both forms are derived from the very common 18-8 stainless steel alloy (18% chromium, 8% nickel) used in tableware and other commercial applications. The composition differences between the 18-8 and 316L alloy are necessitated by the superior corrosion resistance required of implant devices. Very briefly, the addition of molybdenum (3%) to the 18-8 alloy and the reduction of carbon content (0.03% max) confers improved corrosion resistance particularly to pitting and intergranular attack, respectively. Such compositional changes, however, necessitate the addition of nickel (12%) to maintain the stability of the desired microstructure, austenite.
The cleanliness or purity of the refined implant alloy may influence greatly the corrosion resistance and mechanical properties. All steels contain impurities or non-metallic inclusions, which by design are minimized to obtain the desired combination of properties for implan- tation purposes.(3) Many manufacturers recommend fur- ther alloy refinement processes such as vacuum arc remelting and electro-slag refining to optimize implant performance. Although a cast stainless steel alloy is produced (Table 13-1), the vast majority of stainless steel devices used today are manufactured from the wrought alloy owing to the generally improved mechanical properties and reduced impurity content. It is estimated from retrieval studies that stainless steel alloys constitute approximately 60% of the implants used in the United States.(3)
TABLE 13-1 Composition of Surgical Implant Alloys
The mechanical properties of surgical-grade stainless steel are considered to be good relative to other implant alloy systems but certainly not outstanding in the general engineering field (Table 13-2). To maintain the specified austenitic microstructure, the normal hardening and tempering heat treatments of carbon-and low-alloy steels cannot be performed. Indeed, within the composition and phase specifications of SS316, hardening can be achieved only by a process known as cold-working. The effects of cold-working on mechanical properties are presented in Table 13-3, which shows that relative to a fully annealed alloy, cold-working can produce a twofold to threefold increase in yield strength, a 40% increase in ultimate strength, but a corresponding 80% decrease in ductility, thus making the material far more brittle. This is an important consideration when cold-working stainless steel fracture-fixation devices, as in the plate-contouring process. Althou gh such a contoured device may have improved ultimate properties, it is somewhat more prone to catastrophic brittle failure.
The forged stainless steel alloy is used most commonly for fracture-fixation devices because of the desirable effects the forging process has on the grain structure of the fixahon device. For example, the selective directional deformation in forging a plate from bar stock or in making wire results in an optimum "fiber texture" or grain structure in which grains are elongated into fibrous or spindle shapes parallel to the long axis of the device and the expected deforming forces. This microstructure provides greater resistance to crack propagation and breakage than fully-annealed or as-cast devices having equi-axed grain structure.
Of the alloys currently specified for device manufacture, stainless steel is the cheapest and most easily fabricated. To obtain high quality devices, however, careful attention must be given to the melting process, the carbon and impurity content, and the various thermal treatments necessary for shaping and developing desirable mechanical properties. Surface preparation is typically accomplished by mechanical polishing or electro-polishing to remove draw marks, pits, burrs, and surface contamination. The final step in processing stainless steel implants is that of surface passivation in nitric acid to remove surface iron particles and to artificially thicken the surface oxide layer.
Two main types of cobalt-based alloys are used for surgical implant purposes; a cast alloy and a wrought alloy, which vary significantly in composition (Table 13-1). Despite these differences, however, the trade designation of Vitallium (or in Britain, "Stellite") is often applied erroneously to both alloys. The cobalt-based alloys display a useful balance between mechanical properties and biocompatibility, both forms being somewhat superior to stainless steel in strength and corrosion resistance but more expensive to manufacture.
The casting variant is a cobalt-chromium-molybdenum alloy referred to as cast cobalt-chromium alloy, which because of its high rate of work hardening cannot be contoured at the time of surgery. Accordingly, this alloy is typically reserved for implantable devices having a fixed configuration (e.g. total hip prosthesis), and because of its high abrasion resistance is sometimes used for bearing applications including metal-on-metal devices.
The wrought cobalt-chromium alloy is composed primarily of cobalt, chromium, nickel, and tungsten and mechanically exhibits a lower rate of work hardening than the cast alloy. In the fully annealed state the wrought alloy displays a yield stress similar to the more brittle cast variant yet has a much improved ductility (60% strain at fracture) and an ultimate tensile strength approaching that of a heavily cold-worked stainless steel. Moreover, with appropriate working and annealing treatments, the wrought alloy can be made to yield a useful range of strengths and ductilities, giving it much versatility as an implant alloy. The wrought alloy, however, is somewhat less resistant to crevice corrosion than the cast cobalt-chromium-molybdenum alloy. Its use for fracture-fixation purposes is not as yet commonplace, probably as a result of its increased cost compared with that of stainless steel.
This third alloy system used in manufacturing structural implants includes two titanium-based alloys: commercially pure titanium and the Ti-6AI-4V, containing a nominal 6% aluminum and 4% vanadium (Table 13-1). Commerically pure titanium must adhere to specified impurity maxima, particularly with respect to oxygen, since mechanical properties vary markedly with impurity content; the higher the impurity content, the stronger but less ductile the metal. Mechanically, the disadvantages of pure titanium include its relatively low elastic modules (approximately one half that of stainless steel and cobalt-chromium), low shear stength, and its poor abrasion resistance causing it to gall or seize when it is in sliding contact with another metal surface. In addition, it is more difficult to fabricate than stainless steel and is more expensive. Beyond its light weight, its major advantage lies in its inherent corrosion resistance. This is attributable to the spontaneous formation of a strong passivating oxide layer particularly resistant to saline solutions.
The addition of 6% aluminum and 4% vanadium to commercially pure titanium results in an alloy having mechanical properties similar to cold-worked stainless steel (including superior fatigue resistance) yet retaining excellent corrosion resistance. In addition, the Ti-6AI-4V alloy is more easily weldable and machinable than the pure form. Its principal drawback remains its poor resistance to erosion, making it unacceptable for bearing surfaces.
The titanium-based alloys, because they have been more recently introduced and are more costly, have not found widespread application, particularly in veterinary surgery; however, they are the material of choice for patients having known hypersensitivity reactions to any of the constituents of stainless steel or cobalt-chromium alloys.
Many types of nonmetallic materials have been advanced for structural implantation, including amorphous glasses, crystalline ceramics, carbon composites, and polymeric materials. Most nonmetallic materials, however, owing to inherent deficiencies in mechanical properties and biocompatibility, have not found widespread use as structural implants, especially in veterinary practice. An exception, however, is PMMA, sometimes called "acrylic cement," which more and more is used in veterinary orthopaedics, particularly in conjunction with total hip arthroplasty, spinal fracture fixation (see Chapter 19), and as a permanent bone substitute in the treatment of pathologic fractures. Furthermore, very recently PMMA has been applied under internal fracture-fixation plates (so-called "luting") to more evenly distribute forces over the plate-bone interface, thereby combatting premature plate failure.
PMMA is an extremely versatile thermoplastic polymeric material having the capacity to be molded into fixed-design implants or to be polymerized at the time of surgery for tailor-made implant applications such as for "filler" material in total-hip insertion. In its completely amorphous form, PMMA is a clear solid suited to use as ophthalmic implants. The structure of PMMA incorporates methyl and methacrylate groups on alternate carbon atoms as shown.
PMMA manufactured for orthopaedic purposes may be prepared in situ by thoroughly mixing powdered, prepolymerized methylmethacrylate with liquid monomer. Manufacturers have added the appropriate "initiators" and "activators" to the components to begin the polymerization process, which normally lasts 6 or 7 minutes. The reaction is moderately exothermic with attainment of temperatures as high as 122¡C in the center of the material and 92¡C at the surface.(4) Several investigators have suggested that such heat production is not deleterious to vital bone provided blood perfusion is maintained. However, much debate still exists regarding this contention.(4) Clearly it is advisable to control heat production if possible, such as through minimization of the amount of PMMA or by adding fillers.
A further complication observed in the use of in situ polymerization of PMMA is the release of the monomer, MMA, into the circulation, with the potential to cause direct toxic effects such as a precipitous fall in blood pressure and occasional fatalities. There is much experimental evidence implicahng the monomer, MMA, in cardiovascular and respiratory complications as well as marked effects on the peripheral and central nervous system. (12) Although life-threatening complications during insertion of hip prostheses are rare, the most common clinical side-effects are hypotension and hypoxemia. In normal patients this produces no persisting consequences; however, caution should be exercised in applying PMMA in patients with preexistent cardiovascular compromise.
Beyond hemodynamic effects, it is thought that dissolved monomer in combination with traumatization of the oral mucosa from ill-fitting dentures may induce tumor growth in humans.(4) Additionally, the material has been reported to induce hypersensitivity reactions (contact-type) in some patients, as well as in orthopaedic surgeons. The monomer, because it is small and extremely lipophilic, diffuses through intact surgical rubber gloves.(4) Although no evidence exists to suggest that PMMA is a significant health hazard to surgeons or operating room personnel, it is advisable to avoid excessive contact and provide venting via suction when mixing the components in the operating room.
Mechanically, PMMA, when fully hardened, is strong but brittle. Compared with other common thermoplastics, it demonstrates excellent properties, including tensile modulus, tensile strength, flexural rigidity, and resistance to creep. Relative to the metal alloys previously described, however, it is mechanically inferior, having low ductility (<5% strain to failure) and a high susceptibility to stress solvent crazing. Its primary use in orthopaedic surgery, as introduced by Charnley in 1960, is as a "grouting agent."(7) Often the term "bone cement" is applied to PMMA; however, it should be recognized that the polymer has poor adhesive properties and does not function in this capacity.
When using PMMA for purposes other than a grouting agent as it is used in total hip implantation, the surgeon must be aware of its mechanical limitations and should safeguard against the deleterious effect of the exothermic polymerization process on the surrounding tissues. In general, implanted PMMA should be minimal in volume and have a structural configuration to facilitate heat dissipation yet sufficient to withstand imposed intrinsic and extrinsic forces.
Discussion thus far has focused on those materials that are either commonly used in veterinary orthopaedic practice or that have potential for application in small animals in the near future. However, a host of other materials have been introduced or are in the development process, some of which find current, although limited, application in human orthopaedic surgery. These include several new metal alloy systems having superior ductility along with improved corrosion resistance. One such system is MP35N, a multiphase alloy containing a nominal 35% nickel in addition to cobalt, chromium, and molybdenum. In contrast to currently used single-phase alloy systems, this material is a three-phase metal, which by appropriate work hardening and heat treatments can be made to display high yield strengths (300,000 psi) while retaining 10% elongation.
Another alloy system under investigation for implant purposes is a so-called TRIP steel (transformation induced plasticity), which has a composition of 9% chromium, 8% nickel, 4% molybdenum, 0.3% carbon, and the balance iron. This two-phase alloy system, compositionally similar to SS 316L, demonstrates desirable mechanical behavior in response to cold work, namely, excellent ductility and strength. Questions remain, however, regarding its fatigue life, corrosion resistance, and biocompatibility.
Ceramic materials have received intense investigation for application in structural implants owing to their typically high compressive strength and biocompatibility. However, they have not as yet been applied to fracture fixation largely because of their poor ductility, that is, susceptibility to brittle fracture. Commercially, an alumina (Al^2 0^3) total hip (alumina ball and socket with metal stem) has been introduced for human application.
Many other polymers have limited application for structural implant purposes; however, in general those that have sufficient biocompatibility suffer from unsatisfactory creep and stress-relaxation behavior. Ultra high molecular weight polyethylene (UHMWPE), because of its low coefficient of friction with metal, is used as a bearing surface in several multicomponent total joint devices. More recently this material has been reinforced with graphite to retard the inevitable in vivo creep and wear process. Early results from device retrieval studies, however, indicate little advantage of the graphite composite over the pure UHMWPE. Similarly, polytetrafluoroethylene (PTFE) was abandoned for use as an acetabular-bearing surface owing to poor wear and creep phenomena in vivo. Apparently such polymers, although satisfying in vitro wear-testing criteria, undergo a form of mechanochemical deterioration in vivo that is analogous to fretting corrosion in metals.
Polyvinylidine fluoride has been advanced by several veterinary surgeons for use as spinal fracture fixation plates (see section on spinal fracture). However, to my knowledge the material has not been compatibility tested according to ASTM (American Society for Testing and Materials) specifications. Nevertheless, clinical experience would suggest that it is sufficiently biocompatible not to elicit adverse local or systemic sequelae for the duration of implantation.
MATERIALS ASSOCIATED COMPLICATIONS
Clinically, implant failure may be defined as a failure of the implantation procedure to produce satisfactory results. In veterinary practice, failures of internal fixation devices occur with enough frequency to be of considerable concern. Causes of such failure can be grouped into four categories: surgical, material, idiosyncratic, or owner/ animal compliance. Surgical failures relate to errors in surgical judgment or application technique including surgically introduced complications such as infection. Material failures stem from either the chemistry, structural metallurgy, or engineering design deficiencies of the implant. Idiosyncratic failures refer to the selective rejection of an implant by certain patients, often associated with pain, hypersensitivity reactions, implant loosening, or sinus tract formation. Finally, an all too common failure mode in veterinary practice is poor owner/animal compliance with the prescribed postoperative management program. Typically this results in catastrophic failure of the fixation device. Upon retrieval of a failed orthopaedic device, any of these causes may be cited as a sole, major, or contributing factor in failure.
The veterinary surgeon can minimize the risk of surgically related implant failure by acquiring a solid understanding of both the mechanics of fracture fixation as well as the materials limitations of the devices implanted. Moreover, to reduce problems in owner/animal compliance, it is the veterinarian's responsibility to establish strict postoperative management protocols geared to the inherent stability of the fracture treatment. For example, even the best constructed internal fixation device is subject to mechanical failure in the absence of stable fracture reduction, rapid bone healing, or adequate owner/animal compliance.
Materials failures may result from deficiencies in engineering design, manufacturing processing, or handling in the operating room. Clinically, material failure modes fall into three categories: purely mechanical, purely environmental, and conjoint mechanical-environmental (26)
Purely mechanical failures are the result of direct overload including impact or fatigue (discussed previously). Environmental failures stem from the reaction of the physiological environment with the metal, resulting in corrosion (dissolution), which either weakens the device mechanically or elicits an adverse tissue response necessitating device removal. Conjoint mechanical-environmental failures are produced by the combined effects of applied stress in a corrosive environment. Conjoint failure modes include fretting corrosion and corrosion fatigue. Although it is commonly held that stress corrosion cracking is also a conjoint failure mechanism of structural implants, in actual fact the phenomenon has not been conclusively demonstrated in stainless steel in vivo, and only under very specific conditions above 75¡C can it be experimentally induced.
Very little is known about idiosyncratic failures; however, it is strongly suspected that such failures originate from corrosion product-induced hypersensitization phenomena resulting in implant rejection or loosening. Alternatively, selective accelerated corrosion producing a direct toxic effect on neighboring cells can be postulated. Presently there is no reliable means to test animals for preexistent or acquired hypersensitivity reactions to the constituents of implantable devices or to their active organometallic complexes. In humans, it is estimated that approximately 6% of the population has existing hypersensitivities to one or more of the constituents of stainless steel or cobalt-chromium alloys, suggesting a need for routine hypersensitivity screening prior to surgery.(8) Similar statistics are not available on the existence of metal allergy in domestic animals.
Because fracture-fixation devices in veterinary practice are, with few exceptions, manufactured exclusively from stainless steel 316L, it is wise to study the performance and failure mechanisms of this alloy system. Table 13-4 gives a simplified overview of the types of mechan- ical-chemical-environmental interactions to be expected in fracture treatment using internal fixation. It becomes apparent immediately that the success of internal fixation depends on a complex interplay of these factors. In a recent report of the retrieval and analysis of 1120 human orthopaedic implants, the vast majority of which were fracture-fixation devices made of stainless steel, a few important observations regarding performance were made: None of the retrieved implants exhibited obvious generalized forms of corrosion. Mechanical distortions set during the operation procedure were not preferred sites for localized corrosion (except at screw-plate junctions). Localized corrosion was not observed at bending areas. Localized attack in the form of fretting (mechanical) or fretting corrosion (mechanical-environmental) was commonplace at points of metal-on-metal contact in multicomponent implants. Only rarely was corrosion of the bone/plate interface observed.(12)
Such information would suggest that although plate contouring results in unavoidable surface marring, the process itself is mechanically and chemically sound. That is, if performed appropriately with smooth surfaced tools, plate bending does not affect the fatigue or corrosion resistance of the device adversely. Fretting and fretting corrosion, however, at the plate-screw interface cannot be eliminated with present design but if possible should be minimized to reduce the amount of local and systemic distribution of corrosion products. The surgeon's role is to optimize the stability of fracture reduction to minimize the relative motion between plate and screw.
The local tissue changes in response to the mechanical and chemical presence of a metallic implant have been extensively studied (11,17) Alloys currently popular are by design sufficiently inert to be considered biocompatible, at least over the short term. The biologic environment walls off the implanted alloy by interposing a relatively acellular tissue capsule. With accelerated implant degradation, however, inflammatory cells, macrophages, and occasionally foreign body giant cells may be found adjacent to the device. It is not clear whether the rare adverse tissue response to the presence of an implant stems from the toxic nature of the corrosion process or an individualized sensitivity to certain corrosion products or perhaps even a biologically accelerated corrosion rate in certain patients. Typically, from device retrieval and analysis studies in humans, a small number will present with evidence of local infection (pain, inflammation, edema, fluid accumulation, or draining sinuses) (6) months or more after the original surgery. Typically, culture and sensitivity testing reveals no growth of microorganisms, thus the designation of "sterile abscess." Removal of the device usually effects prompt relief (24-48 hours), and very commonly there is obvious tissue discoloration from corrosion. Again, whether this phenomenon reflects an acquired individual hypersensitivity or merely a toxic reaction to unexplained accelerated corrosion is uncertain. Clearly, there exist numerous metal-related and indeed metal-induced hypersensitivity reactions in the literature; however, the role of antigenicity in the implant rejection process is not well understood.(14)
TABLE 13-4 Examples of Interactions Between Implant and Body in Fracture Treatment
Beyond toxic or antigenic local effects, there presently exists concern regarding the potential for metallic implants to incite tumor formation in bone at the site of implantation. There are approximately 30 reported cases in the canine of osteosarcoma formation in the vicinity of fracture-fixation devices. The latency period for tumor induction is approximately 4 to 6 years, but clearly the incidence is low relative to the number of animals receiving metallic implants. The causal relationship, however, is strong and argues for routine removal of fracture-fixation devices following bony union and remodeling.
Another well-recognized local effect of fracture-fixation devices is the osteoporotic remodeling of bone immediately beneath the plate. Such a response is thought to stem from the load-sharing capacity of the plate, which acts to bypass forces around the underlying bone. Accordingly, Wolff's law of dynamic remodeling can in theory not operate to maintain the appropriate balance of osteoblastic versus osteoclastic activity, and osteoporosis ensues. In this regard much emphasis has been directed to reducing the rigidity (modulus) of the fracture fixation device and its attachment to bone. However, to date, no suitable low-modulus (or low-rigidity) system is available commercially.
The question of long-term systemic effects of metallic implantation is presently under active investigation, particularly in light of the inevitable in vivo corrosion process and a growing trend to implant metallic devices in younger human patient populations. Theories have been proposed and animal studies performed that implicate corrosion or wear products in remote allergenic response, depression of bacterial resistance, low-level aberrations in serum biochemistries, and remote carcinogenesis. However, except for the well-documented clinical incidence of metal-induced hypersensitivity reactions, the other effects are largely conjectural or of little potential clinical significance based on the present state of the art. Implantation, of high-surface-area orthopaedic devices, as has been suggested to promote bony ingrowth and secure fixation, however, may result in an unacceptably high rate of release of corrosion products, both locally and systemically. In general, until more information is known, it remains good practice to consider internal fixation devices as temporary implants and to suggest routine removal when practical.
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