Orthopedic Applications

Chapter ii.5.6  Orthopedic ­Applications Malchesky, P. S., Koo, A. P., Skibinski, C. I., Hadsell, A. T., & Rybicki, L. A. (2009). Apheresis technologies and clinical applications: The 2007 International Apheresis Registry. Therapeutic Apheresis and Dialysis, published online: Aug 11 2009. Mulholland, J. W. (2008). Cardiopulmonary Bypass. Surgery (Oxford), 26(12), 486–488. Murphy, G. J., & Bryan, A. J. (2004). Cardiopulmonary Bypass. Surgery (Oxford), 22(6), 126–128. Nishinaka, T., Nishida, H., Endo, M., Miyagishima, M., Ohtsuka, G., et al. (1996). Less blood damage in the impeller centrifugal pump: A comparative study with the roller pump in open heart surgery. Artificial Organs, 20, 707. Nogawa, A. (2002). The future of artificial lungs: An industry perspective. Journal of Artificial Organs, 5, 211. Organisation for Economic Co-operation and Development. (2005). Health at a Glance – OECD Indicators 2005: OECD Publishing. Sargent, J. A., & Gotch, F. A. (1996). Principles and Biophysics of Dialysis. In C. Jacobs, C. M. Kjellstrand, K. M. Koch, & J. F. Winchester (Eds.), Replacement of Renal Function by Dialysis (4th ed.). Netherlands: Springer.

CHAPTER II.5.6  ORTHOPEDIC ­APPLICATIONS Nadim James Hallab and Joshua James Jacobs Department of Orthopedic Surgery, Rush University Medical Center, Chicago, IL, USA

INTRODUCTION Orthopedic biomaterials are enormously successful in restoring mobility and quality of life to millions of individuals each year. Orthopedic implants include reconstructive implants, fracture management products, spinal products, rehabilitation products, arthroscopy products, electrical stimulation products, and casting products. These products are generally used for fracture fixation enhancement, joint replacement or dynamic stabilization. More specific orthopedic applications within these categories are listed here. Fracture fixation devices: • Spinal fixation devices • Fracture plates • Wires, pins, and screws • Intramedullary devices • Artificial ligaments Joint replacement (Figure II.5.6.1): • Hip arthroplasty • Knee arthroplasty • Spine arthoplasty • Ankle arthroplasty • Shoulder arthroplasty • Elbow arthroplasty • Wrist arthroplasty • Finger arthroplasty Dynamic stabilization devices (new): • Spine stabilization devices.

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Tayama, E., Hayashida, N., Oda, T., Tomoeda, H., Akasu, K., et al. (1999). Recovery from lymphocytopenia following extracoporeal circulation: Simple indicator to assess surgical stress. Artificial Organs, 23, 736. United States Renal Data Service (USRDS). (2009). USRDS 2009 Annual Data Report: Atlas of End-Stage Renal Disease in the United States. Bethesda, MD: National Institutes of Health, National Institute of Diabetes and Digestive and Kidney Diseases. Wing, A. J., & Jones, E. (2000). Epidemiology of End-Stage Failure: A Global Perspective. In A. M. El Nahas (Ed.), Mechanisms and Clinical Management of Chronic Renal Failure (2nd ed.). Oxford, UK: Oxford University Press. Zydney, A. L. (2006). Therapeutic Apheresis and Blood Fractionation. In J. D. Bronzino (Ed.), The Biomedical Engineering Handbook: Tissue Engineering and Artificial Organs (3rd ed.). Boca Raton, FL: CRC Press.

Orthopedic Biomaterials Market The overwhelming success of orthopedic biomaterials is exemplified by their worldwide market, dominating biomaterial sales at approximately $24 billion in 2007, with an expected growth rate of 7–9% annually. Global sales of trauma fracture management products only totaled approximately $3.7 billion in 2007, whereas $10 billion was spent on knee and hip joint replacements (Figure II.5.6.1). Global sales of knee implant products equaled approximately $5.8 billion in 2007, representing approximately 1.5 million knee replacement surgeries which include first-time joint replacement procedures and revision procedures for replacement, repair or enhancement of an implant product or component from a previous procedure. Revision procedures are growing at an accelerated rate of approximately 60% in the United States.

Orthopedic Biomaterials Orthopedic biomaterials are generally limited to those materials that withstand cyclic loadbearing applications. While metals, polymers, and ceramics are used in orthopedics, it remains metals, which have over the years uniquely provided appropriate material properties such as high strength, ductility, fracture toughness, hardness, corrosion resistance, formability, and biocompatibility necessary for most loadbearing roles required in fracture fixation and total joint arthroplasty (TJA). The use of orthopedic biomaterials generally falls into one of three surgical specialty categories: upper extremity; spine; or lower extremity; and each specialty is typically divided into three general categories: pediatric; trauma; and reconstruction. Despite these numerous specialties and the hundreds of orthopedic applications, there are only a few orthopedic

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SECTION II.5  Applications of Biomaterials TABLE II.5.6.1     Most Common Orthopedic Biomaterials Material Metals Ti alloy (Ti-6%Al-4%V) Co–Cr–Mo alloy Stainless Steel Polymers Polymethylmethacrylate (PMMA) Ultra-high molecular weight polyethylene (UHMWPE) Ceramics Alumina (Al2O3) Zirconia (ZrO2)

Primary Use(s) Plates, screws, TJA components (non-bearing surface) TJA components TJA components, screws, plates, cabling Bone cement Low friction inserts for bearing surfaces in TJA

Bearing surface TJA components Bearing surface TJA components

STRUCTURE AND PROPERTIES OF CALCIFIED TISSUES

FIGURE II.5.6.1  Total joint arthroplasties (TJA) are currently used

to replace hip, knee, shoulder, etc. (Courtesy of BioEngineering Solutions Inc.)

metals, ceramics, and polymers which dominate all implants. Knowing the general properties, uses, and limitations of the “primary” orthopedic biomaterials is requisite to understanding what is required to improve the performance of current implant materials, and why only a few dominate the industry. A summary of seven more prevalent orthopedic biomaterials and their primary use(s) are listed in Table II.5.6.1.

Orthopedic Biomaterials Design New biomaterials for orthopedic purposes face the same concerns present in current implants: (1) the material must not adversely affect its biological environment; (2) in return the material must not be adversely affected by the surrounding host tissues and fluids; and (3) new materials must exceed the performance of present materials. Thus, understanding the interrelationship between the structure and properties of the natural tissues that are being replaced is important. An appreciation of the “form–function” relationship in calcified tissues will help provide insight into critical factors determining implant design, as well as deciding which materials best meet a specific orthopedic need.

There are several different calcified tissues in the human body, and several different ways of categorizing them. All calcified tissues have one thing in common: in addition to the principal protein component, collagen, and small amounts of other organic phases, they all have an inorganic component hydroxyapatite (abbreviated OHAp, HA or Ca10(PO4)6(OH)2). In the case of long bones, such as the tibia or femur, an understanding of the organization of these two principal components is central to characterization. It has been convenient to treat the structure of compact cortical bone (e.g., the dense bone tissue found in the shafts of long bones) using four levels of organization. The first, or molecular, level of organization is the collagen triple helical structure (tropocollagen) and OHAp crystallography. It forms a hexagonal unit cell with space group symmetry P63/m and lattice constants a = 9.880 Å and c = 6.418 Å, containing two molecular units, Ca5(PO4)3OH, per unit cell. How cells produce this mineral phase, and whether it is the first calcium phosphate laid down, are subjects of considerable research at present. Because of its small crystallite size in bone (approximately 2 × 20 × 40 nm), the X-ray diffraction pattern of bone exhibits considerable line broadening, compounding the difficulty of identifying additional phases. A Cabearing inorganic compound in one of the components of calcified tissues has led to the development of a whole class of ceramic and glass–ceramic materials that are osteophilic within the body (i.e., they present surfaces that bone chemically attaches to). As yet, we do not know fully how the two components, collagen and OHAp, are arranged and held together at this molecular level. Whatever the arrangement, when it is interfered with (as is apparently the case in certain bone pathologies in which

Chapter ii.5.6  Orthopedic ­Applications the collagen structure is altered during formation), the result is a bone that is formed or remodeled with seriously compromised physical properties. The second or ultrastructural level may be loosely defined as the structural level observed with transmission electron microscopy (TEM) or high magnification scanning electron microscopy (SEM). Here, too, we have not yet achieved a full understanding of the collagen–OHAp organization. It appears that the OHAp can be found both inter- and intrafibrillarly within the collagen. At this level, we can model the elastic properties of this essentially two-component system by resorting to some sort of linear superposition of the elastic moduli of each component, weighted by the percent volume concentration of each. The third or microstructural level of organization is where these fibrillar composites form larger structures, fibers, and fiber bundles, which then pack into lamellartype units that can be observed with both SEM and optical microscopy. The straight lamellar units forming the plexiform (lamellar) bone are found generally in young quadruped animals, the size of cats and larger. This is the structural level that is described when the term “bone tissue” is used or when histology is generally being discussed. At this level, composite analysis can also be used to model the elastic properties of the tissue, thus providing an understanding of the macroscopic properties of bone. Unfortunately, this modeling is very complex and a complete description lies beyond the scope of this chapter (and the authors). Interested readers are referred to some of the original sources (Katz, 1980a,b). The fourth level is that of each of the macroscopic levels of each bone sample or large section of bone. Since a significant portion of bone is composed of collagen, it is not surprising to find that, in addition to being anisotropic and inhomogeneous, bone is also viscoelastic like all other biological tissues. Duplicating such properties with long-lasting synthetic biomaterials remains an unrealized goal of orthopedic biomaterials, where the history of implant development has been characterized by the elimination of available candidate materials based on their poor performance, rather than production of biocompatible synthetic bone-mimetic materials. Perhaps the best example of how orthopedic biomaterials have undergone implant design improvements over the past 100 years resulting in widespread success is total hip replacement (THR) or total hip arthroplasty (THA). Newer types of total joint arthroplasties, such as those currently used for disc replacements (total disc arthroplasty, TDA), have benefitted from the arduous history of the total joint arthroplasty. Many, if not all, of the biomaterial-related issues (both ­mechanical and biological) that impact the performance of the THR’s are applicable to other orthopedic implants. Therefore, this chapter will detail where we have come from, using the total hip arthroplasty, and where we are now, using the example of current total disc arthroplasty designs, general current orthopedic materials technology, and future developments

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as a proxy of clinical concerns of orthopedic biomaterial development and current technology. The history of total hip arthroplasty is particularly pertinent to biomaterials science, because it is one of the best illustrations of how an implant first used over a century ago has evolved to the highly successful status it has, primarily because of advances in biomaterials.

BIOMATERIALS DEVELOPMENT: A HISTORY OF TOTAL HIP ARTHROPLASTY The earliest attempts to restore mobility to painful and deformed hip joints took place in the 1820s (White in 1822 and Barton in 1827), and centered on simply removing the affected femoral and acetabular bone involved. This evolved in the 1830–1880s into ghastly attempts to restore mobility using interpositional membranes between the femoral head and acetabulum, where such materials as wooden blocks and animal (e.g., pig) soft tissue were tried. The first prosthetic hip replacement is dated to 1890, when Gluck published a description of a carved ivory femoral head replacement using bone cement-like materials such as pumice and plaster of paris to secure the implants in place (Walker, 1978; Stillwell, 1987). The interpositional membrane strategy continued from the 19th into the 20th century, where the use of new implant materials in the early 1900s (1900–1920) included organic materials (e.g., pig bladders and peri-implant soft tissues), and inorganic materials such as gold foil. The use of the individual’s own soft tissues was the most popular method of interpositional membrane hip surgery. The limited success of this procedure prevented widespread use, and thus the treatment (rather than surgery) of painful, disfigured, and “frozen” (ankylosed) hip joints remained commonplace into the 1920s.

Mold Arthroplasty It was not until 1923 when Marius Smith-Peterson was credited with ushering in the modern era of total joint replacement with his development of the “mold” arthroplasty (Figure II.5.6.2), made of glass, inspired by a shard of glass found in a patient’s back with a benign synovial-like membrane around it. This mold or cup arthroplasty was designed as a cup that fitted in between the femoral head and the acetabular cup, and articulated on both surfaces prompting a “tissue-engineered” synovial/cartilage-like layer. This was the first widespread attempt to develop a better interpositional membrane, a technique that had been in practice for the previous 100 years. The efforts of Smith-Peterson and his colleagues over the years from 1923–1938 were spent improving the fracture resistance of the glass mold arthroplasty cup design, using materials such as early polymers (e.g., celluloid or phenol-formeldehyde Bakelite® or Formica®) and improved glass, e.g., Pyrex®. But it was not until

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SECTION II.5  Applications of Biomaterials

FIGURE II.5.6.2  The history of total hip arthroplasty is particularly pertinent to biomaterials science because it is one of the best illustrations of how an implant first used over a century ago has evolved to the highly successful status it has, primarily because of advances in biomaterials.

1939 when the first metal implant, a cobalt alloy termed Vitallium®, became available and, as used by Venable, Stuck, and Beach, showed that the corrosion resistance of the hip arthroplasty provided sufficient biocompatibility and performance to be incorporated into future popular implants such as the Judet, Moore and Thompson hip arthroplasties. In 1937 Venable, Stuck, and Beach published a landmark article that was the first to analyze in a systematic fashion the electrolytic effects of various metals and alloys on bone and tissues (e.g., aluminum, copper, iron, nickel,

lead, gold, magnesium, silver, stainless steel, and other alloys). They arrived at the conclusion that Vitallium® (a cobalt–chromium alloy) was superior to the other metals in corrosion resistance and in the mechanical properties required for implants (Venable et al., 1937; Charnley, 1979). By observing the effects of corrosion, and proposing guidelines for performance Venable, Stuck, and Beach set the standard by which future metallic alloys were selected for use in hip and other types of implants. The superior material properties of the Vitallium® alloy facilitated further design modifications of the mold

Chapter ii.5.6  Orthopedic ­Applications arthroplasty by Otto E. Aufranc (Figure II.5.6.2), where the rim of the Smith-Peterson mold was removed (which was often the cause of adhesions and cup “freezing,” with subsequent pain and immobility), and matching curves on the inner and outer surface were machined to meet at the rounded outer edge. Despite the high short-term success rates (<4 years) reported by Aufranc (>82%), the overall failure rate remained high (>50%). Another design modification of the mold arthroplasty in the 1940–1950s was the fixation of the mold to the acetabulum rim with screws by such physicians as AlbeePearson and Gaenslen. Although used in only four cases, Gaeslen reported using a cobalt alloy mold fixed to the acetabulum and another fixed to the femoral head, creating a metal-on-metal total hip replacement. The popularity of mold arthroplasties endured into the 1970s, when they remained touted as the treatment of choice for traumatic arthritis of the hip by leading orthopedic surgeons (Harris, 1969). However, back in the 1930s the natural progress in THA development was the progression from mold arthroplasty to short-stem prosthesis.

Femoral Head Prostheses/Short-Stem Prostheses Femoral head prosthetics were first made of such materials as ivory (Gluck in 1890) and rubber (Delbet in 1919), and were cemented (using a plaster-like cement) for stability (Walker, 1978; Stillwell, 1987). At about the same time these replacement heads were first fitted with a short-stem by Earnest Hey Groves, who used an ivory nail to replace the articular surface of the femur. These types of implants were rare and remained unpopular compared to mold arthroplasties until 1937, when Harold Bohlman, using the work of Venable and Stuck, designed a corrosion-­resistant cobalt–chrome alloy femoral head replacement with a short-stem. This design was popularized by the Judet brothers in Paris in 1946; they used polymethylmethacrylate (PMMA), which was presumed biologically inert in vivo, to manufacture short-stemmed prostheses (Figure II.5.6.2). Initial good results were soon replaced with problems of implant fracture and excessive wear debris, and by the early 1950s these implants were losing favor and being removed by surgeons. Vitallium® (cobalt–chrome alloy) eventually replaced acrylic in several other short-stem designs. However, there were sound short-stem designs as early as 1938, when Wiles introduced the cobalt alloy femoral shell attached to the femur with a central nail. This design was later popularized by Peterson in 1950, where he used a similar Vitallium® shell design with a central nail and a plate attached to the nail for added stability. Others adopted and adapted the Judet brothers design using Vitallium®, such as J. Thompson (1951) and Rossignal (1950). Rossignal designed large threads onto the stem to aid in fixation. These short-stem designs were subject to what was deemed high shear stress, and

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resulted in early loosening and failure in some patients. Short-stem designs were gradually replaced by longerstem designs that provided less stress concentration.

Long-Stem Prostheses Long-stem prostheses continued the trend established by short-stemmed prostheses, that is, more and more weightbearing forces were transferred to the femur though an intramedullary stem. The pattern for a long-stem prosthesis was established in 1940 by Bohlman in collaboration with Austin T. Moore, in which they implanted a 12 inch Vitallium® prosthesis that replaced the femoral head and had long supports that were screwed into the outside of the femoral shaft (Moore, 1943). And while there were innovations in long-stem design in the 1940s, such as the door knob design of Earl McBride where a threaded stem was screwed into the intramedullary canal of the femur for fixation and load transferal, these designs were not popular. It was not until 1950 with the designs of Frederick R. Thompson and Austin T. Moore that longstemmed prostheses became popular (Figure II.5.6.2). These designs were cast in Vitallium® (cobalt–chrome alloy), and required the removal of the femoral head but only part of the neck. The design of Moore differed from that of Thompson in that it had fenestrations through the implant to allow bone growth, and it had a rear vane to enhance rotational stability. Initially, these implants were used without bone cement. Evidence for the successful designs of the Thompson and Moore prostheses is proved by their continued use today, with only slight variations from the original. Despite the excellent design of these early long-stemmed prostheses they were primarily successful when used in place of diseased femoral heads, and did not work well when acetabular reaming was required. Therefore, this inadequacy prompted the development of the total hip replacement arthroplasty.

Total Hip Replacement Arthroplasty Philip Wiles is credited with first total hip arthroplasty in 1938, when he used a stainless steel ball secured to the femur with a bolt and a stainless acetabular liner secured with screws (Wiles, 1953). The results of this design were disappointing, because of the poor corrosion resistance of early stainless steel in  vivo, and the high stress concentrations of short-stemmed prostheses. An adaptation of this design that proved successful was developed by G. K. McKee and J. Watson-Farrar in 1951. They used a stain-less steel cup and long-stemmed prosthesis (­Thompson stem) which failed rapidly due to the poor corrosion resistance of the stainless steel, and was then replaced by cobalt–chrome alloy with greater success. The McKee–Farrar prosthesis evolved quickly to incorporate a true spherical femoral head that was undercut at the neck to reduce the impingement of the head on the rim of the acetabular prosthesis to provide a greater

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SECTION II.5  Applications of Biomaterials

range of mobility (Figure II.5.6.2) (McKee and WatsonFarrar, 1943). The next milestone in the evolution of modern total hip arthroplasty was the advent/popularization of acrylic dental bone cement, first used by Sven Kiar in 1950 to attach a plastic prosthesis to bone (Charnley, 1964). Later that year the Hospital for Joint Diseases in New York used polymethylmethacrylate (acrylic) bone cement as a means of fixation in total hip arthroplasties (Wilson and Scales, 1970). The development of acrylic bone cement dramatically reduced the rates of loosening associated with metal–metal total hip arthroplasty. The Stanmore metal–metal design which used a horseshoeshaped cup was popular, but it led to excessive wear and was replaced by a complete cup. McKee and WatsonFarrar adapted their design to facilitate bone cement with a land-mine-like studded acetabular cup intended to maximize mechanical fixation. The 1950s marked the introduction and popularization of the total hip arthroplasty where it became simple and reliable enough to be practiced on a wide scale by the average orthopedic surgeon. However, the squeaking reported to occur in Judet and some later metal prostheses was identified by Charnley to be a result of the relatively high frictional forces in the joint. These high torque and frictional forces resulted in the generation of significant metallic debris, which purportedly resulted in early loosening. In 1960 Charnley developed a “low friction arthroplasty” device using shells of polytetrafluoroethylene, PTFE (commonly called Teflon® in related publications) on the femoral and acetabular sides, which resulted in early/immediate failures because of massive debonding and wear debris. This was quickly followed by a thick-walled PTFE acetabular component articulating on a small head designed to reduce the shearing forces and torque. However, this design also ­generated excessive wear debris, which produced immediate and severe inflammation and failure of the ­prosthesis. ­Charnley then replaced the PTFE with high density polyethylene which was not as friction-free as Teflon®, but was 1000-times more wear resistant. This prototype of total hip arthroplasty developed in 1962 was the basis for future designs which remain the most popular form of total hip arthroplasty performed today (Figure II.5.6.2). The basic design of Charnley was modified by Muller with variable neck sizes and larger heads. At the same period metal-on-metal designs by Smith, Ring, and others (Ring, 1968) were unsuccessful challengers to the basic Charnley metal-on-polymer design. Other currently adopted design modifications were developed by Ling, Aufranc, Turner, Amstutz, Harris, and Galante, which include such innovations as femoral prosthesis geometrical modification for increases in stability and mobility, modular components for increased customization, porous coatings, surface texturing/coating to increase fixation and bone ingrowth, etc. Charnley is often deified in orthopedic literature as the metaphorical spark

that lit the flames of innovation in prosthetic design. This is a typical surgeon-centered over-glorification. For one thing, other implant designs which predate Charnley, such as the all metal McKee-Farrar THA implant, have enjoyed similar success rates to those reported by Charnley. More importantly, total hip arthroplasty is, perhaps, the best example of how orthopedic biomaterials and implant success have evolved over the last century through the innovation and hard work of many scientists and physicians, and advances in areas of materials technology, biomechanics, biochemistry, immunology, infectious diseases, thrombosis, and pharmacology, to name a few.

NEW DEVELOPMENTS: TOTAL DISC ARTHROPLASTY In contrast to THA, total disc arthroplasty (TDA) is a relatively recent development which has yet to become a mainstream option for treating disc degeneration when compared with fusion. Spinal fixation device usage is steadily increasing, where the number of cervical and lumber fusions increased 111% from the years 1993 to 2003, to roughly 105 fusions per 100,000 people in the US which is about 305,000 fusions per year (Cowan et  al., 2006). The ultimate goal of intervertebral disc replacement technology is to replace spine fusion, eliminate pain, and restore structure and mobility. The first disc arthroplasties were cobalt alloy spheres, implanted as early as 1957 in between vertebrae without any method of fixation (Harmon, 1963). A decade later, stainless steel metal spheres termed Fernström balls were used in 103 patients (Figure II.5.6.3) starting in 1969 (Fernstrom, 1966; McKenzie, 1995). Other clinicians around the same time tried using polymer balls made of polymethylmethacrylate, but the results were disastrous (Hamby Wallace, 1959). The era of modern disc arthroplasty began in 1982 with the first functional artificial intervertebral disc, the SB Charité™, at the Charité Hospital in Berlin (Figure II.5.6.4) (Büttner-Janz, 1992). This design used the low-friction arthroplasty principle of John Charnley, which by that time had been successfully used in total hip replacement for over 20 years. The Charité TDA consisted of a UHMWPE sliding core, which articulated between two metal endplates with multiple teeth-like projections for fixation to the vertebral endplates (Figure II.5.6.4). The following sections detail the different kinds of material selection used in different modern total disc replacement (TDR) designs in the bourgeoning world of both lumbar and cervical total disc replacements. In addition to what is termed total disc arthroplasty, there are several disc sparing and motion preservation devices that, while not technically TDAs, are a new type of implant called “dynamic stabilization” implants. However, none of these dynamic stabilization implants are approved for general use. The following

Chapter ii.5.6  Orthopedic ­Applications

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FIGURE II.5.6.3  Fernström Ball implants (stainless steel metal

spheres) were the first disc arthroplasties. (Courtesy of BioEngineering Solutions Inc.)

list summarizes three central types of motion preserving spine implants, their articulation couple, and primary material constituents: 1. Lumbar Posterior Motion Sparing Technology ­(Figure II.5.6.5) Stabilimax NZ®, Applied Spine Technologies, Inc. (metal-on-metal: articulation: cobalt alloy-oncobalt alloy, titanium) TOPS®, Impliant, Inc. (metal-polymer-metal with elastic core: cobalt alloy-polyurethane-cobalt alloy) Dynesis®, Zimmer, Inc. (metal-polymer-metal with elastic core-like structures: cobalt alloy-polyurethane-cobalt alloy) DIAM™, Medtronic (all polymer, polyester, silicone) 2. Cervical Disc Arthroplasty (Figure II.5.6.6) Bryan® Cervical Disc, Medtronic (metal-polymer-metal: elastic core: cobalt alloy-polyurethane-cobalt alloy) PCM®, Cervitech (metal-on-polymer: articulation: cobalt alloy-on-UWMWPE) PRESTIGE® Cervical Disc, Medtronic (metal-onmetal: articulation: stainless steel-on-stainless steel)

FIGURE II.5.6.4  In 1982, Schellnack and Büttner-Janz developed

the SB Charité™ artificial disc, which consisted of a UHMWPE sliding core articulating unconstrained between two highly polished stainless steel metal endplates. (Courtesy of Depuy Spine, Inc.)

PRODISC-C®, Synthes, Inc. (metal-on-polymer: articulation: cobalt alloy-on-UWMWPE) Secure®-C, Globus Medical (metal-on-polymer: articulation: cobalt alloy-on-UWMWPE) 3. Lumbar Disc Arthroplasty (Figure II.5.6.7) NUBAC™, Pioneeer Surgical, Inc. (polymer-on-­ polymer: articulation: PEEK on PEEK) SB Charité®, DePuy Spine, Inc. (metal-on-polymer: articulation: cobalt alloy-on-UWMWPE) Prodisc™ II, Synthes, Inc. (metal-on-polymer: articulation: cobalt alloy-on-UWMWPE) Maverick®, Medtronic (metal-on-metal: articulation: cobalt alloy-on-cobalt alloy) eDisc™, Theken (metal-polymer-metal: elastic core: cobalt alloy-polyurethane-cobalt alloy) Freedom™ Lumbar Disc, Axiomed, (metal-polymermetal: elastic core: cobalt alloy-polyurethanecobalt alloy) Activ L, Asculap (metal-on-polymer: articulation: cobalt alloy-on-UWMWPE)

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SECTION II.5  Applications of Biomaterials

(b)

(a)

(c)

FIGURE II.5.6.5  Dynamic Spine Stabilization Implants: A new class of motion preservation implants are under investigational use in the

lumbar spine: (a) Stabilimax NZ® (Applied Spine Technologies, Inc.); (b) TOPS® (Impliant, Inc.); (c) Dynesis® (Zimmer, Inc.); all of these use a combination of articulation and spring or elastomeric interior components to provide both articulation and resistance force back a neutral position. (Courtesy of BioEngineering Solutions, Inc.)

(a)

(b)

(c)

(d)

FIGURE II.5.6.6  Cervical IDRs (Intervertebral Disc Replacement): There are a number of cervical total disc replacements in use which employ

the three primary types of disc joint articulation. 1. Metal-on-Polymer Articulation: (a) The PRODISC-C® (Synthes, Inc.); and (b) PCM® (Cervitech); use Co-alloy endplates that articulate on a polymeric (UHMWPE) core that is mechanically fixed to one of the endplates and articulates in a ball-and-socket type manner. 2. Metal-on-Metal Articulation: (c) The PRESTIGE® Disc (Medtronic) has metal-on-metal articulation where the end plates are constructed of stainless steel. 3. Elastic Core Articulation: (d) The Bryan® Cervical Disc System (Medtronic) is axially symmetric and incorporates cobalt-chrome alloy clamshell-shaped endplates, which flex upon a loadbearing polymeric (polyurethane-based) nucleus core. A novel feature in the design of this component is the polyurethane flexible membrane that surrounds the entire articulation and forms a sealed space containing a saline lubricant to reduce friction and prevent migration of any wear and corrosion debris. (Pictures provided courtesy of DePuy Spine, Inc., Medtronic, Spine Solutions, Inc., and BioEngineering Solutions, Inc.)

Chapter ii.5.6  Orthopedic ­Applications

(a)

(b)

(c)

(d)

(e)

(g)

(f)

(h)

FIGURE II.5.6.7  Lumbar total disc replacements include metal-

on-polymer, metal-on-metal, polymer-on-polymer, and flexible core technologies. (a) Metal-on-Polymer Articulation: The LINK® SB Charité III (DePuy Spine, Inc.) cobalt–chrome alloy endplates articulate on a mobile bearing ultra-high molecular weight polyethylene core. The endplates are covered with an osteoconductive surface of titanium/calcium-phosphate double coating under the trade name “TiCaP®.” (b) Metal-on-Polymer Articulation: The Prodisc (Synthes, Inc.) lumbar TDR is composed of cobalt–chrome–molybdenum alloy and covered with a porous titanium alloy and articulates on a central core of UHMWPE. (c) Metal-on-Polymer Articulation: ActivL (Aesculap) uses a polymeric center core intended to allow both translation and rotation and to more closely approximate physiological motion. (d) Metal-on-Polymer Articulation: Dynardi (Zimmer, Inc.) is a disc replacement implant with two opposing cobalt alloy (Co–Cr–Mo) endplates coated with porous pure titanium for bone ingrowth, that articulate on a semi-constrained UHMWPE core. (e) Polymer-on-Polymer Articulation: NUBAC™ (Pioneer Surgical, Inc.) is a polymer-on-polymer disc arthroplasty device and the first polyetheretherketone (PEEK)-on-PEEK articulated disc arthroplasty device. (f) Metal-on-Metal articulation: The Maverick Disc (Medtronic) uses metal-on-metal articulation where the end plates are constructed of Co-alloy. (g) Elastic Core Articulation: The Theken eDisc™ (Theken Disc) represents another step in the evolution orthopedic implant devices in that as well as containing an elastic polyurethane based core it provides measured in  vivo load information to the surgeon and patient via electronic sensors and transmitters. (h) Elastic Core Articulation: The Freedom™ Lumbar Disc (Axiomed) uses a viscoelastic polymer (like polyurethane) to replicate the native function of a natural disc. The elastic core in combination with the implant design provides a three-dimensional motion that functions within the natural biomechanics of the spine. (Pictures provided courtesy of DePuy Spine, Inc., Medtronic, Spine Solutions, Inc., and courtesy of BioEngineering Solutions, Inc.)

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In general, current TDA designs have a THA-like primary articulation (or motion) using polymer-on-metal, polymer-on-polymer, metal-on-metal or some form of an all elastic core technology. However, because the emerging area of motion preservation in spine orthopedics facilitates the use of a wide array of implant designs and materials, the following sections will detail some of these designs and the materials they are comprised of. Metal-on-Polymer Articulation.  Examples of metalon-polymer articulating TDA implants are as follows: SB Charité III Artificial Disc (Depuy Spine, Inc.). The Charité III includes thick cobalt chrome endplates (Figure II.5.6.7). The metal endplates are coated with an osteoconductive surface of porous titanium beads and calcium-phosphate double coating under the trade name “TiCaP®.” This coating is also used in non-cemented total joint arthroplasty such as femoral stems, acetabular cups, ankle joint prostheses, and dental implants (Liefeith et al., 2003). Prodisc (Synthes, Inc.). The lumbar TDA Prodisc reflects how hip and knee designs have shaped TDA designs (Figure II.5.6.7). The two cobalt-chrome-molybdenum alloy endplates are covered with a porous titanium alloy, and articulate on a UHMWPE core. The locking polyethylene core provides ball and socket movement which is different from other “mobile” bearing designs, e.g., the Charité. The cobalt-chrome alloy endplates are also coated with pure titanium ­(Plasmapore), in similar fashion to the Charité. Metal-on-Metal Articulation.  Examples of metal-onmetal articulating TDA are currently used in both lumbar and cervical disc replacement applications, and dynamic stabilization devices of the lumbar spine. Dynamic Stabilization: Stabilimax NZ® (Applied Spine Technologies, Inc.). The Stabilimax NZ® dynamic spine system is an implant that provides stabilization of the lumbar spine for patients with spine surgery that results in less superstructure for stability (e.g., bone removal for spinal stenosis; ­Figure II.5.6.5). It is composed of two ball-and-socket joints with cobalt alloy on cobalt alloy (ASTM F-75) articulation, two titanium alloy (Ti-6Al-4V) screws, and an interposed spring of cobalt alloy in between the two ball-andsocket joints (Figure II.5.6.5a). This type of device attempts to offer a less invasive surgical procedure than fusion or disc replacement. Total Disc Arthroplasty: Prestige Disc. The Prestige Disc (Medtronic) uses metal-on-metal articulation where the end plates are currently constructed of stainless steel (and in the near future cobalt alloy) (Figure II.5.6.6). The articulation mimics that of a ball-andsocket construct with a more constrained center of rotation, i.e., not the joint kinematics of mobile bearing. Because the wear resistance of stainless steel

850

SECTION II.5  Applications of Biomaterials

exceeds that of polyethylene, the amount of total particulate debris generated by metal-on-metal couples is an order of magnitude less than that produced by metal-on-­polymer couples. The location of this implant next to vital tissues raises concerns associated with the release of metallic debris. However, the importance of metal debris will be resolved primarily through careful clinical follow-up and peer reviewed analysis. Polymer-on-Polymer Articulation.  The lower loads of the spine when compared to hip and knee replacements has enabled the use of polymer-on-polymer articulation designs of total disc arthroplasty devices, which if successful will be a new phenomena in the world of total joint arthroplasty. NUBAC™ (Pioneer Surgical, Inc.) is a polymer-on-­polymer disc arthroplasty device made of polyetheretherketone (PEEK). This implant aims to maintain or restore the disc height and mechanical function by replacing only the nucleus of a spine disc, theoretically using a less invasive procedure than other total disc arthroplasty implants, i.e., it is inserted into a partially resected disc, and has partial structural support of part of the remaining disc (the annulus around the outside of the implant). It does not restrict any physiological rotational motions, leaving constraint and stability to the retained surrounding annulus of the disc and ligaments. The risk of implant extrusion is low, because biomechanics of motion tend to push it in the opposite direction (Figure II.5.6.7). Elastic Intervertebral Replacement Devices.  The advent of elastic core total disc replacement implants represents another pioneering effort in the world of total joint implants. Elastic core total joint replacement (TJR) technology is feasible because there are relatively small motions associated with disc arthroplasty implants when compared to a hip or knee. There are generally two different kinds of total disc arthroplasty designs with elastic “core “components: (1) elastic cores interposed between two endplates; and (2) all elastic devices for nucleus replacement. Total Disc Arthroplasty: Bryan® Cervical Disc Prosthesis (Medtronic). The Bryan® Cervical Disc System was approved by the United States FDA for distribution on May 12, 2009, and incorporates cobalt-chrome alloy endplates, which sandwich and flex on a loadbearing polymeric (polyurethane-based) nucleus core. The endplates have a porous coating of 250 micron titanium beads sintered to the cobalt alloy endplates. There is also a polyurethane flexible membrane that surrounds the entire articulation, which forms a sealed space to prevent migration of any wear and corrosion debris. This sheath also aims to prevent the intrusion of connective tissue or the creation of a pseudo-­capsule over time. The

inclusion of a sheath component to address the clinical concern of wear and corrosion debris (discussed later) is a significant development in the design of orthopedic joint arthroplasty components. All these spinal implants seek to replace fusion as the operative solution for disc degeneration or injury. Many of these designs are new to orthopedics in terms of materials and/or design, and have not withstood the test of time; it is likely that only a few of these designs will dominate and become more popular over time. However, these disc replacement implants represent how imaginative new designs and materials are being used to address age-old orthopedic problems that remain unresolved in this new 21st century.

CURRENT BIOMATERIALS IN TOTAL ARTHROPLASTY Today, the archetype of the total hip implant remains much as it was in the 1970s, albeit with a wider variety of implant materials and geometries (Figure II.5.6.8). Current THA is typically constructed of a titanium or cobalt–chromium alloy femoral stem (cemented with polymethylmethacrylate, PMMA, or press fit into place; Figure II.5.6.9), connected to a “modular” cobalt–chrome alloy or ceramic head that articulates on a ultra-high molecular weight polyethylene (UHMWPE) or ceramic acetabular cup fitted into a titanium or cobalt–chromium cup liner which is cemented, screwed or press fitted into place (Figure II.5.6.9). Despite this simple archetype of the total hip replacement, there are hundreds of variations on this theme offered to today’s orthopedic surgeons, with little in terms of absolute guidelines as to which type of implant is (or which of the over 10 major manufacturers has) the best for well-defined orthopedic disease states. However, there are some general guidelines. Typically, implants in older individuals (>80 years of age) are cemented into place with PMMA bone cement, because the chance for revision is minimal when compared to younger individuals (<60 years), and removing bone cement is both technically challenging and may compromise the availability of bone stock. Generally, there are choices of surface roughness, coatings, geometry, material composition, etc., and each manufacturer claims that its product is superior to the rest. This, in combination with little or no publicly available information tracking the performance of each type of implant in patients, precludes accurate scientific analysis of which implant materials and designs perform best. Additionally, competition between manufacturers, and the requisite attention to marketing required to compete in the marketplace, has resulted in a dizzying array of new implants released each year claiming to be an improvement over last year’s model. These claims are suspect, because the typical total hip replacement enjoys a success rate of over 90% at seven years, therefore in most cases a minimum of seven to ten years must pass

Chapter ii.5.6  Orthopedic ­Applications

851

FIGURE II.5.6.8  Examples of typical current total hip arthroplasty (THA) components available from a single manufacturer: titanium alloy

stem, with a cobalt-base alloy (ASTM F-75) modular head bearing on an ultra-high molecular weight polyethylene (UHMWPE) liner within a titanium alloy cup. Also shown are a ceramic head and three acetabular sockets with various surfaces for both cemented and cementless fixation. From left to right the stems are all components of the VerSys® Hip System (Zimmer, Inc, Warsaw, IN, USA), and from left to right are designated Beaded Fullcoat, Beaded Fullcoat with distal flutes, Cemented, Fiber Metal Taper, and Fiber Metal Midcoat. (Photographs courtesy of Zimmer, Inc.)

before such claims can be substantiated, and even then proof of superior performance is compromised by a myriad of external factors, such as the surgeon, region of the country, average activity of patient populations, etc. This conflict between science, marketing, and market share may (in the opinion of this author) represent the single biggest obstacle to the scientific determination of superior implant design and progress. The unenviable responsibility rests with the FDA to prevent the zeal of

economic pressure from undermining implant design in a regressive fashion. Today, optimal implant selection is primarily based on which material couple may best suit the individual (Figure II.5.6.10).

Polymers Polymers are most commonly used in orthopedics as articulating bearing surfaces of joint replacements

852

SECTION II.5  Applications of Biomaterials

FIGURE II.5.6.9  Examples of cemented and non-cemented stems

showing the plug and centralizer used with a cemented stem. (Picture courtesy of Zimmer, Inc.)

FIGURE II.5.6.10  Examples of the three types of bearing cou-

ples used in modern TJA. From top to bottom: metal-on-polymer, ceramic-on-ceramic, and metal-on-metal (Lineage™ line from Wright Medical Technology, Inc. Arlington, TN, USA).

(­Figure II.5.6.8), and as an interpositional cementing material between the implant surface and bone (­Figure II.5.6.8). Polymers used as articulating surfaces must have a low coefficient of friction, and low wear rates when in articulating contact with the opposing surface, which is usually made of metal. Initially, John Charnley used polytetrafluoroethylene (PTFE) for the acetabular component of his total hip arthroplasty (Figure II.5.6.2). However, its accelerated creep and poor stress degradation (for the material he used) caused it to fail in vivo, requiring replacement with his ultimate choice, ultrahigh ­molecular weight polyethylene (UHMWPE). Polymers used for fixation as a structural interface between the implant component and bone tissue require the appropriate mechanical properties of a polymer, which can be molded into shape and cured in vivo. The first type to be used, polymethylmethacrylate (PMMA), was again popularized by Charnley, who borrowed it from the field of dentistry. He adapted dental PMMA as a “grouting” material to fix both the stem of the femoral component and the acetabular component in place, and thus distributed the loads more uniformly from the implant to the bone. Since high interfacial stresses result from the accommodation of a high modulus prosthesis within the much lower modulus bone, the use of a lower modulus interpositional material has been a goal of those seeking to improve upon PMMA fixation. Thus, polymers such as polysulfone have been tried as porous coatings on the implant’s metallic core to permit mechanical interlocking through bone and/or soft tissue ingrowth into the pores. However, to date PMMA remains the substance of choice for orthopedic surgeons. This requires polymers that have surfaces that resist creep under the stresses found in clinical situations, and that have high enough yield strengths to minimize plastic deformation. As indicated earlier, the important mechanical properties of orthopedic polymers are yield stress, creep resistance, and wear rate. These factors are controlled by such parameters as molecular chain structure, molecular weight, and degree of branching or (conversely) of chain linearity. One of the more prevalent polymerics used in orthopedics today is a highly cross-linked ultra-high molecular weight polyethylene (UHMWPE), which is typically used in total joint arthroplasty as a loadbearing articulating surface, designed to provide low friction loadbearing articulation. Polyethylene is available commercially in three different grades: low density; high density; and UHMWPE. The better packing of linear chains within UHMWPE results in increased crystallinity, and provides improved mechanical properties required for orthopedic use even though there is a decrease in both ductility and fracture toughness. In total hip arthroplasty applications, an acetabular cup of UHMWPE typically articulates against a femoral ball of cobalt–chromium alloy. The predominant problem presented by these metal– polymer articulating surfaces is the production of wear

Chapter ii.5.6  Orthopedic ­Applications particles, i.e., polymer debris. The resultant wear of the polyethylene bearing purportedly produces billions of sub-micron sized wear particles annually, in the <1–10 micron range. Producing greater cross-linking of polyethylene, using chemical and radiation techniques, has only recently improved its wear resistance in orthopedic applications. Wear tests have shown that the wear resistance of ­UHMWPE is improved by cross-linking with gamma irradiation at 2.5–5.0 Mrad and below, as evidenced by simulator studies; however, this can negatively affect such physical properties as tensile strength (McKellop et al., 2000). Therefore, care must be taken to minimize any negative oxidative effects, while preserving high wear resistance characteristics. Although newer more highly cross-linked polyethylene has generally been accepted as superior to previous implant UHMWPE, there remains incomplete data regarding its ultimate long-term performance. In order to maximize the performance characteristics of polyethylene, it is cross-linked prior to fabrication into its final form, e.g., an acetabular cup. Typically, an extruded bar of polyethylene is crosslinked using conventional gamma irradiation, and then heat treated to reduce residual free radicals.

Ceramics In recent years, ceramics and glass ceramics have played an increasingly important role in implants. Although used in Europe for over a quarter century, the FDA has only recently (3 February, 2003) approved the first ceramic-on-ceramic bearing hip implant to be used in total hip replacement procedures (Figure II.5.6.10). The primary reason for the introduction of this alternative bearing surface is the superior wear resistance of ceramics when compared to metal–metal or metal–polymer bearing surfaces. This, and other improved properties such as resistance to further oxidation (implying inertness within the body), high stiffness, and low friction require the use of full-density, controlled, small, uniform grain size (­usually less than 5 μm) ceramic materials. The small grain size and full density are important, since these are the two principal bulk parameters controlling the ceramic’s mechanical properties. Any voids within the ceramic’s body will increase stress, degrading the mechanical properties. Grain size controls the magnitude of the internal stresses produced by thermal contractions during cooling. In ceramics, such thermal contraction stresses are critical, because they cannot be dissipated as they can in ductile materials via plastic deformation. Alumina (Al2O3) and zirconia (ZrO2) ceramics have been used in orthopedic THA for the past 30 years. The first ceramic couple (alumina–alumina) was implanted in 1970 by Pierre Boutin. Since the outset, the theoretical advantage of hard-on-hard articulating surfaces was low wear. Ceramics, because of their ionic bonds and chemical stability, are also relatively biocompatible. Initial concerns about fracture toughness and wear have

853

been addressed by reducing grain size, increasing purity, lowering porosity, and improving manufacturing techniques (e.g., hot isostatic pressing, HIP). Early failures of these couples were plagued with both material-related and surgical errors. The very low wear rates combined with steadily decreasing rates of fracture (now estimated to occur 1 in 2000 over 10 years) have resulted in the growing popularity of all ceramic bearings. Zirconia was introduced in 1985 as a material alternative to Al2O3 for ceramic femoral heads, and was gaining market share because of its demonstrably enhanced mechanical properties in the laboratory when compared to alumina. Femoral heads of zirconia can typically withstand 250 kN (or 25 tons), a value generally exceeding that possible with alumina or metal femoral heads. However, mechanical integrity of all ceramic components are extremely dependent on manufacturing quality controls, as evidenced in the recall of thousands of zirconia ceramic femoral heads by their manufacturer, St. Gobain ­Desmarquest, in 2001. This was because of in  vivo ­fracture of some components due to a slight unintended variation in the manufacturing sintering process, caused when the company bought a newer high-throughput assembly line type oven. In general, ceramic particulate debris is chemically stable and biocompatible, and causes no untoward biologic responses at high concentrations. There have been recent attempts to take advantage of the osteophilic surface of certain ceramics and glass ceramics. These materials provide an interface of such biological compatibility with osteoblasts (bone-forming cells) that these cells lay down bone in direct apposition to the material in some form of direct chemicophysical bond. Special compositions of glass ceramics, termed bioglasses, have been used for implant applications in orthopedics. The model proposed for the “chemical” bond formed between glass and bone is that the former undergoes a controlled surface degradation, producing an SiO-rich layer, and a Ca, P-rich layer at the interface. Originally amorphous, the Ca, P-rich layer eventually crystallizes as a mixed hydroxycarbonate apatite structurally integrated with collagen, which permits subsequent bonding by newly-formed mineralized tissues. There is still an entirely different series of inorganic compounds that have also been shown to be osteophilic. These include OHAp, which is the form of the naturally occurring inorganic component of calcified tissues, and calcite, CaCO, and its Mg analog, dolomite, among others being studied. The most extensive applications in both orthopedics and dentistry have involved OHAp. This has been used as a cladding for metal prostheses for the former, and in dense, particulate form for the latter. The elastic properties (modulus) of OHAp and related compounds are compared with those of bone, dentin, and enamel in Table II.5.6.2. The use of both OHAp and glass ceramics as cladding on the metallic stems of hip prostheses is still another method of providing fixation instead of using PMMA.

854

TA B L E I I . 5 . 6 . 2    Mechanical Properties of Dominant Orthopedic Biomaterials

ASTM Designation

Trade Name and Company

Yield Strength (Elastic limit)

Ultimate Strength

Fatigue Strength (Endurance Limit)

Hardness

Elongation at Fracture

(Examples)

(GPa)

(MPa)

(MPa)

(MPa)

HVN

(%)

Cortical Bone$ Low strain High strain

15.2 40.8

114t –

150c/90t 400c–270t

30–45 –

– –

– –

Polymers UHMWPE PMMA

0.5–1.3 1.8–3.3

20–30 35–70

30–40t 38–80t

13–20 19–39

60–90 (Mpa) 100–200 (Mpa)

130–500 2.5–6

Ceramics Al2O3 ZrO2

366 201

– –

3790c/310t 7500c/420t

– –

20–30 (Gpa) 12 (Gpa)

– –

Metals Stainless steels

ASTM F138

Protusul S30, Sulzer

190

792

930t

241–820

130–180

43–45

ASTM F75

Alivium, Biomet CoCrMo, Biomet Endocast SlL, Krupp Francobal, Benoist Girard Orthochrome, DePuy Protosul 2, Sultzer Vinertia, Deloro Vitallium C, Howmedica VitalliumFHS, Howmedica Zimaloy, Zimmer Zimalloy, Micrograin Vitallium W, Howdmedica HS25l, Haynes Stellite MP35N, Std Pressed Steel Corp. TJA 1537, Allvac Metasul, Sulzer

210–253

448–841

655–1277t

207–950

300–400

4–14

210 200–230

448–1606 300–2000

1896t 800–2068t

586–1220 340–520

300–400 8–50 (RC)

10–22 10–40

200–300

960

1300t

200–300

41 (RC)

20

CSTi, Sulzer Isotan, Aesculap Werke Protosul 64WF, Sulzer Tilastan, Waldemar Link Tivaloy 12, Biomet Tivanium, Zimmer

110 116

485 897–1034

760t 965–1103t

300 620–689

120–200 310

14–18 8

Co–Cr Alloys

ASTM F90 ASTM F562 ASTM 1537 Ti Alloys CPTi Ti-6Al-4V

ASTM F67 ASTM 136

ASTM: American Society for Testing and Materials (ASTM International). $: Cortical bone is both anisotropic and viscoelastic thus properties listed are generalized. c: Compression. t: Tension. RC: Rockwell Hardness Scale.

SECTION II.5  Applications of Biomaterials

Orthopedic Biomaterial

Elastic Modulus (Young’s Modulus)

Chapter ii.5.6  Orthopedic ­Applications In these cases, the fixation is via the direct bonding of bone to the ­cladding surface.

Metals Since the principal function of the long bones of the lower body is to act as loadbearing members, it was reasonable that the initial materials introduced to replace joints, such as artificial hips, were metals. Both stainless steel, such as 316L, and cobalt–chromium alloys became the early materials of choice, because of their relatively good corrosion resistance and reasonable fatigue life within the human body. Of course, their stiffness, rigidity, and strength exceeded those of bone considerably. However, in certain applications, owing to size restrictions and design limitations (e.g., in rods used to straighten the spine in scoliosis), fatigue failures did occur. Metals remain the central material component of state-of-the-art total hip arthroplasties. Metals provide appropriate material properties, such as high strength, ductility, fracture toughness, hardness, corrosion resistance, formability, and biocompatibility necessary for use in loadbearing roles required in fracture fixation and total joint arthroplasty (TJA). Implant alloys were originally developed for maritime and aviation uses, where mechanical properties such as high strength and corrosion resistance are paramount. There are three principal metal alloys used in orthopedics and particularly in total joint replacement: (1) titanium-based alloys; (2) cobaltbased alloys; and (3) iron-(stainless steel) based alloys. The alloy’s specific differences in strength, ductility, and hardness generally determine which of these three alloys is used for a particular application or implant component. However, it is the high corrosion resistance of all three alloys, more than anything, which has led to their widespread use as loadbearing implant materials. These material properties of metals (Table II.5.6.2) are due to the miraculous nature of the metallic bond, atomic microstructure, and elemental composition of alloys.

Stainless Steel Alloys Stainless steels were the first metals to be used in orthopedics in 1926. However, it was not until 1943, when ASTM 304 was recommended as a standard implant alloy material, that steels were reliable as an implant alloy. All steels are comprised of iron and carbon, and may typically contain chromium, nickel, and molybdenum. Trace elements such as manganese, phosphorous, sulfur, and silicon are also present. Carbon and the other alloy elements affect the mechanical properties of steel through alteration of its microstructure. The form of stainless steel most commonly used in orthopedic practice is designated 316LV (American ­Society for Testing and Materials F138, ASTM F138; others include F139, F899, F1586, F621, etc). “316” classifies the material as austenitic, the “L” denotes the

855

low carbon content, and “V” defines the vacuum under which it is formed. The ­carbon content must be kept at a low level to prevent carbide (chromium–carbon) accumulation at the grain boundaries. Although the mechanical properties of stainless steels are generally less desirable than those of the other implant alloys (lower strength and corrosion resistance), stainless steels do possess greater ductility, indicated quantitatively by a three-fold greater “percentage of elongation at fracture” when compared to other implant metals (Table II.5.6.2). This aspect of stainless steel has allowed it to remain popular as a material for cable fixation components in total knee arthroplasty, and a low cost alternative to titanium and cobalt alloys.

Cobalt–Chromium Alloys Of the many Co–Cr alloys available, there are currently only two predominantly used as implant alloys (Table II.5.6.3). These two are: (1) cobalt–chromium–molybdenum (CoCrMo), which is designated ASTM F75 and F76; and (2) cobalt–nickel–chromium–molybdenum (­CoNiCrMo) designated as ASTM F562. Other cobalt alloys approved for implant use include one which incorporates tungsten (W) (CoCrNiW, ASTM F90), and another with iron (CoNiCrMoWFe, ASTM F563). Co–Ni–Cr–Mo alloys which contain large percentages of nickel (25–37%) promise increased corrosion resistance, yet raise concerns of possible toxicity and/or immunogenic reactivity (discussed later) from released nickel. The biologic reactivity of released nickel from Co–Ni–Cr alloys is a cause for concern under static conditions, and due to their poor frictional (wear) properties Co–Ni–Cr alloys are also inappropriate for use in articulating components. Therefore the dominant implant alloy used for total joint components is CoCrMo (ASTM F75). Cobalt alloys are generally cast into their final shape, because they are susceptible to work-hardening at room temperatures. The improvements in strength and hardness gained by cold-working are not worth the loss in fracture toughness. Thus, Co–Cr–Mo alloy hip implant components are predominantly manufactured using lost wax (investment) casting methods. Although Co–Cr–Mo alloys are the strongest, hardest, and most fatigue resistant of the alloys used for joint replacement components, care must be taken to maintain these properties, because the use of finishing treatments can also reduce these same properties (Table II.5.6.2). For example, sintering of porous coatings onto femoral or tibal TJA Co–Cr–Mo stems can decrease the fatigue strength of the alloy from 200–250 MPa to 150 MPa after heating (annealing) the implant at 1225°C.

Titanium Alloys Titanium alloys were developed in the mid-1940s for the aviation industry, and were first used in orthopedics

856

SECTION II.5  Applications of Biomaterials

TAB L E I I . 5 . 6 . 3    Approximate Weight Percent of Different Metals Within Popular Orthopaedic Alloys Alloy

Ni

Stainless steel (ASTM F138) Co–Cr–Mo alloys (ASTM F75) (ASTM F90) (ASTM F562) Ti alloys CPTi (ASTM F67) Ti-6Al-4V (ASTM F136) 45TiNi Zr Alloy (95% Zr, 5% Nb)

N

Co

Cr

10–15.5 <0.5

*

<2.0

*

9–11 33–37

Ti

Mo

Al

Fe

Mn

Cu

W

C

Si

V

17–19 *

2–4

*

61–68

*

<0.5

<2.0

<0.06

<1.0

*

61–66

27–30 *

4.5–7.0 *

<1.5

<1.0

*

*

<0.35

<1.0

*

* *

46–51 35

19–20 * 19–21 <1

* 9.0–11

* *

<3.0 <1

<2.5 * <0.15 *

14–16 *

<0.15 *

<1.0 * <0.15 *

*

*

*

*

99

*

*

0.2–0.5 *

*

*

<0.1

*

*

*

*

*

*

89–91 *

5.5–6.5

*

*

*

*

<0.08

*

3.5–4.5

55 *

* *

* *

* *

45 *

* *

* *

* *

* *

* *

* *

* *

* *

* *

* Indicates less than 0.05%. Note: Alloy compositions are standardized by the American Society for Testing and Materials (ASTM vol. 13.01).

TAB L E I I . 5 . 6 . 4    Electrochemical Properties of Implant Metals (Corrosion Resistance) in 0.1 M NaCl at pH=7 Alloy Stainless steel

ASTM Designation

Density

Corrosion Potential (vs Calomel)

Passive Current Density

Breakdown Potential

(g/cm3)

(mVolts)

(mAmps/cm2)

(mVolts)

ASTM F138

8.0

−400

0.56

200–770

ASTM F75

8.3

−390

1.36

420

ASTM F67 ASTM 136 ** **

4.5 4.43 4.45 6.4–6.5

−90 to −630 −180 to −510 −530 −430

0.72–9.0 0.9–2.0 0.68 0.44

>2000 >1500 >1500 890

Co–Cr–Mo Alloys Ti Alloys CPTi Ti-6Al-4V Ti5Al2.5Fe Ni45Ti

** No current ASTM standard. The Corrosion Potential represents the open circuit potential (OCP) between the metal and a calomel electrode. The more negative the OCP, the more chemically reactive and thus the less corrosion resistance. Generally low current density indicates greater corrosion resistance. The higher the breakdown potential the better, (i.e. the more elevated the breakdown potential, the more stable the protective layer).

around the same time. Two post-World War II alloys, commercially pure titanium (CPTi) and Ti-6Al-4V, remain the two dominant titanium alloys used in implants. Commercially pure titanium (CPTi, ASTM F67) is 98–99.6% pure titanium. While CPTi is most commonly used in dental applications, the stability of the oxide layer formed on CPTi, and consequently its high corrosion resistance (Table II.5.6.4), and its relatively higher ductility (i.e., the ability to be cold-worked) compared to Ti-6Al-4V, has led to the use of CPTi in porous coatings (e.g., fiber metal) of TJA components. Generally, joint replacement components (i.e., TJA stems) are made of Ti-6Al-4V (ASTM F136) Nb rather than CPTi, because of its superior mechanical properties (Table II.5.6.2). Titanium alloys are particularly good for THA components because of their high corrosion resistance compared with stainless steel and Co–Cr–Mo alloys. A passive oxide film (primarily of TiO2) protects both

Ti-6Al-4V and CPTi alloys. Generally, Ti-6Al-4V has mechanical properties that exceed those of stainless steel, with a flexural rigidity less than stainless steel and Co– Cr–Mo alloys. The torsional and axial stiffness (moduli) of Ti alloys are closer to those of bone, and theoretically provide less stress shielding than cobalt alloys and stainless steel. However, titanium alloys are particularly sensitive to geometrical factors, in particular notch sensitivity. This reduces the effective strength of a component by increasing the material’s susceptibility to crack initiation and propagation through the component. Therefore, care is taken both in the design geometry and in the fabrication of titanium alloy components. Perhaps the greatest drawback of titanium alloys is their relative softness compared with Co–Cr–Mo alloys (Table II.5.6.2), and their relatively poor wear and frictional properties. Ti-6Al-4V is >15% softer than Co–Cr–Mo alloy, and also results in significantly more wear than Co–Cr–Mo when used

Chapter ii.5.6  Orthopedic ­Applications in applications requiring articulation, e.g., total knee arthroplasty (TKA) or THA femoral heads. Thus, titanium alloys are seldom used as materials where hardness or resistance to wear is the primary concern.

NEW ALLOYS AND SURFACE COATINGS The quest for new THA metal alloys with improved biocompatibility and mechanical properties remains an ongoing one. The use of Ti alloys, Co–Cr–Mo alloys or stainless steels in a specific application generally involves trade-offs of one desirable property for another. Some examples of this are the sacrifice of chemical inertness for hardness (wear resistance), as is the case with titanium alloy for Co–Cr–Mo in TJA bearing surfaces, and the compromise of strength for ductility when using stainless steel instead of titanium and Co–Cr–Mo alloys for bone fixation cables. Although new alloys claim to be just that, “new,” they are often merely variations of the three categories of implant metals previously described (which are already approved for use). These improved alloys usually contain only the minor addition of new elements to protect assertions of substantial equivalence to existing ASTM and FDA approved alloys, therefore easing the burden of regulatory approval. These new alloys generally fall under one of four categories, i.e.: (1) titanium alloys; (2) cobalt alloys; (3) stainless steels; and less approved (4) refractory group metals and alloys.

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New Zirconium and Tantalum Alloys Zirconium (Zr) and tantalum (Ta) are characterized as refractory metals (others include molybdenum and ­tungsten) because of their relative chemical stability (­passive oxide layer) and high melting points. Because of its high strength, chemical stability, and resistance to wear, alloys such as zirconium (e.g., Oxinium™) are likely to gain popularity as orthopedic biomaterials. Because of their surface oxide layer’s stability, zirconium and tantalum are highly corrosion resistant. Corrosion resistance generally correlates with biocompatibility (although not always), because more stable metal oxides tend to be less chemically active and/or biologically available and are thus less participatory in biologic processes. This enhanced biocompatibility is produced by the relatively thick surface oxide layer (approximately 5 micrometers), and the ability to extend ceramic-like material properties (i.e., hardness) into the material through techniques such as oxygen enrichment. This has resulted in the production of new implant components using these alloys (see Figure II.5.6.11) (e.g., oxidized zirconium TKA femoral components, Smith and Nephew, Inc.). Although, new zirconium alloys such as Oxinium™ generally possess high levels of hardness (12 GPa) and wear resistance (approximately 10-fold that of cobalt and titanium alloys, using abrasion testing), which makes them wellsuited for bearing surface applications, they are costly

FIGURE II.5.6.11  Examples of new THA and TKA oxidized zirconium components currently gaining popularity because of enhanced mechanical and biocompatibility properties (Oxinium™, Smith and Nephew, Inc., Memphis, TN, USA). (Photographs courtesy of Smith and Nephew, Inc.)

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to manufacture and currently are sought after in special circumstances where issues such as a metal allergy (or more accurately metal hypersensitivity) require particular attention to biocompatibility. As difficulties associated with the cost of forming and machining these metals are overcome, the use of these materials is expected to grow (Black, 1992).

New Titanium Alloys One new group of titanium alloys put forward for orthopedic component uses molybdenum at concentrations greater than 10%. The addition of molybdenum acts to stabilize the BCC (Body Centered Cubic) (beta) phase at room temperature; thus these alloys are referred to as beta titanium alloys. These beta titanium alloys promise 20% lower moduli, which are closer to bone and thus provide better formability with maintenance of other mechanical properties typical of Ti-6-4. Other attempts at improving traditional Ti-6Al-4V alloys seek to improve biocompatibility and mechanical properties by the substitution of vanadium (a relatively toxic metal) with other less toxic metals. Two such titanium alloys include Ti5Al2.5Fe and Ti6Al7Nb, which substitute iron and niobium for vanadium, respectively. These alloys have similar properties to traditional Ti-6-4, yet they claim higher fatigue strength and a lower modulus, thus enhancing bone to implant load transfer.

New Cobalt Alloys Some “newer” cobalt alloys are identical in composition to traditional alloys, but use novel processing techniques to manipulate the microstructure of the implant materials to improve their mechanical properties. One such example recently patented, TJA-1537, although compositionally identical to ASTM F75, claims enhanced wear resistance and fatigue strength through elimination of carbide, nitride, and second phase particles (Allegheny Technologies). These particles normally form at the grain boundaries within a standard F75 CoCrMo alloy, and act to decrease wear and fatigue resistance. Other new cobalt alloys under development for use in orthopedics seek to improve biocompatibility by eliminating nickel, and improve mechanical properties by reducing the carbon content, thus avoiding carbide precipitation at grain boundaries.

New Stainless Steels Because of the desirable cost, machinability, and ductility of stainless steels, there are still efforts to improve its mechanical properties to compete with cobalt and titanium alloys. The relatively poor corrosion resistance and biocompatibility of stainless steels when compared to titanium and Co–Cr–Mo alloys provide

FIGURE II.5.6.12  Examples of currently used surface coatings on stems of THA to enhance both short- and long-term fixation.

incentives for development of improved stainless steels. New alloys such as BioDur® 108 (Carpenter Technology Corp.) attempt to solve the problem of corrosion with an essentially nickel-free austenitic stainless alloy. This steel contains a high nitrogen content to maintain its austenitic structure, and boasts improved levels of tensile yield strength, fatigue strength, and improved resistance to pitting corrosion and crevice corrosion, as compared to nickel-containing alloys such as Type 316L (ASTM F138).

Surfaces and Coatings A variety of surface coatings are currently used to enhance the short- and long-term performance of implants by encouraging bone ingrowth and providing enhanced fixation. These different surfaces include roughened titanium, porous coatings made of cobalt chromium or titanium beads, titanium wire mesh (fiber mesh), plasmasprayed titanium, and bioactive non-metallic materials such as hydroxyapatite or other calcium phosphate compositions (Figure II.5.6.12). Currently osteoconductive and osteoinductive growth factors such as transforming growth factor beta (TGF beta) are being developed for use as osteogenic surface coating treatments to enhance orthopedic implant fixation.

ORTHOPEDIC BIOMATERIALS: CLINICAL CONCERNS Implant biocompatibility/performance is dependent on the type and amount of degradation produced by wear and electrochemical corrosion. Host response to orthopedic implant debris is central to clinical performance (Willert and Semlitsch, 1977). Implant loosening due

Chapter ii.5.6  Orthopedic ­Applications to aseptic osteolysis accounts for over 75% of TJA implant failure, and is the predominant factor limiting the longevity of current total joint arthroplasties; other reasons include infection (7%), recurrent dislocation (6%), periprosthetic fracture (5%), and surgical error (3%) (Holt et  al., 2007). Properly positioned implants tend to wear at predictable rates. However, there are variable amounts of debris-induced bone loss around implants in patients with similar rates of implant wear (i.e., debris generation). It is commonly noted that some individuals with severely worn components can demonstrate little periprosthetic bone loss, while others with modest amounts of wear can demonstrate extensive osteolysis and implant loosening (Willert et  al., 1990; Huo et al., 1992; Jacobs et al., 1992; Jasty et al., 1994; Huk et al., 1994; Schmalzried et al., 1994; Harris, 1995; Thompson and Puleo, 1995; Yao et al., 1995; Goodman et al., 1998; Granchi et al., 1998; Jones et al., 1999; von Knoch et  al., 2000). Debris-induced immune reactivity, aseptic inflammation, and subsequent early failure can be as high as 4–5% at 6–7 years post-operatively in current generation metal-on-metal total hip arthroplasties (Jacobs and ­Hallab, 2006; Korovessis et  al., 2006; Milosev et al., 2006). The benefits provided to patients by orthopedic implants in terms of pain, mobility, and quality of life, are immeasurable. Therefore, the following sections which focus on the problems associated with implants, seek to provide the student of biomaterials with a foundation for understanding the relevant issues for orthopedic research and are not intended to serve as an indictment of orthopedic materials. Despite their overwhelming success over the long term (>7 years), orthopedic biomaterials have been associated with adverse local and remote tissue responses. It is generally the degradation products of orthopedic biomaterials (generated by wear and electrochemical corrosion) which mediate these adverse effects. This debris may be present as particulate wear, colloidal nanometer size complexes (specifically or non-specifically bound by protein), free metallic ions, inorganic metal salts/oxides or in an organic storage form such as hemosiderin. Clinical aspects of biocompatibility regarding polymer and metal release from orthopedic prosthetic devices have taken on an increasing sense of urgency, due to the escalating rates of people receiving implants and the recognition of extensive implant debris within local and remote tissues. Particulate debris has enormous specific surface areas available for interaction with the surroundings and chronic elevations in serum metal content. Clinical issues associated with biomaterial degradation can be broken down into four basic questions: (l) how much material is released from the implant; (2) where is the material transported to and in what quantity; (3) what is the chemical form of the released degradation products (e.g., inorganic precipitate versus soluble organometallic complex); and (4) what are the pathophysiological interactions and consequences

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of such degradation? The answers to these questions, over the long term, remain largely unknown. There is a growing body of literature addressing the issues associated with the first two questions. However, little is currently known with regard to the latter two questions. The remainder of this chapter will focus on that which is known (and of orthopedic clinical concern) regarding biomaterial degradation (through wear and electrochemical corrosion), dissemination of debris, and consequent local/systemic effects.

Orthopedic Biomaterial Wear The generation of wear debris, and the subsequent tissue reaction to such debris, is central to the longevity of total joint replacements. In fact, particulate debris is currently extolled as the primary factor affecting the long-term performance of joint replacement prostheses, and the primary source of orthopedic biomaterial degradation (based on overall implant mass or volume lost). Particulate debris generated by wear, fretting or fragmentation induces the formation of an inflammatory reaction, which at a certain point promotes a foreign– body granulation tissue response that has the ability to invade the bone–implant interface. This commonly results in progressive local bone loss that threatens the fixation of both cemented and cementless devices (Jacobs et al., 1994b; Jacobs, 1995; Jacobs et al., 2001).

Mechanisms of Wear Debris Generation Wear involves the loss of material in particulate form as a consequence of relative motion between two surfaces. Two materials placed together under load will only contact over a small area of the higher peaks or asperities. Electro-repulsive and atomic binding interactions occur at the individual contacts and, when the two surfaces slide relative to one another, these interactions are disrupted; this results in the release of material in the form of particles (wear debris). The particles may be lost from the system, transferred to the counterface or remain between the sliding surfaces. There are primarily three processes which can cause wear: (1) abrasion – by which a harder surface “plows” grooves in the softer material; (2) adhesion – by which a softer material is smeared onto a harder counter surface forming a transfer film; and (3) fatigue – by which alternating episodes of loading and unloading result in the formation of subsurface cracks which propagate to form particles that are shed from the surface. Wear Rates.  During an initial “wearing in” period, the relative motion of surfaces causes a large number of asperities to break, resulting in a high wear rate. After this initial period, the actual contact area increases and the two surfaces can be said to have adapted to one other. Over time, the wear rates decrease and eventually become linearly dependent on the contact force and

SECTION II.5  Applications of Biomaterials

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sliding distance represented by the steady-state wear equation:

year). The most common wear couple for hip and knee arthroplasty currently in use in the US is a cobalt-based alloy head (most commonly a Co–Cr–Mo alloy ASTM F75) bearing on an ultra-high molecular weight polyethylene (UHMWPE) cup or liner. The linear wear rates of this couple are generally in the order of 0.1 mm/year, with particulate generation as high as 1 × 106 particles per step or per cycle. Clinically, implant wear rates have been found to increase with the following: (1) physical activity; (2) weight of the patient; (3) size of the femoral head (32 versus 28 millimeters); (4) roughness of the metallic counterface; and (5) oxidation of the polyethylene (Jacobs et al., 1994a; Jacobs, 1995; Jacobs et al., 2001). Although well established in the hip and knee, newer spinal implants are now undergoing the same types of analyses to try to assess which type of implant works best. The following two basic types of articulating bearing designs are available in spine, knee, and hip arthroplasties.

V = KFx

where V is volumetric wear (mm3/year), K is a material constant of the material couple, F is the contact force (N) and x is the distance of relative travel (mm). Different types of orthopedic materials and couples produce different amounts and kinds of wear debris. Hard-on-hard material couples such as metal-on-metal articulations generally produce less wear (weight loss) than metal-on-polymer (see Figure II.5.6.13). There is a great deal of variability associated with in  vivo wear rates of orthopedic biomaterials, which are generally measured by radiographic follow-up studies. Radiographic wear measurements are expressed as linear wear rates, whereas in vitro studies generally report volumetric wear. Volumetric wear can be directly related to the number of wear particles released into periprosthetic fluids (typically in the order of 1 × 109 of particles per

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FIGURE II.5.6.13  A comparison of the amount of wear debris generated from different types of total joint arthroplasties. There is relatively

less (10×) polymeric debris generated by a total disc arthroplasty with a metal-on-polymer articulation. Note: Figure References: Metal-Poly (Holt et al., 2007); Ceramic-Poly (Jacobs et al., 1994c); Metal-X-linked Poly (Martinon et al., 2006); Metal-X-linked Poly (Jacobs et al., 1994c); Metal-Xlinked Poly (Hallab, 2009); Metal-X-linked Poly (Hallab, 2009); Metal-Metal (Huk et al., 1994); Ceramic-Ceramic (Hallab, 2009); Metal-UHMWPE TDA (Hallab, 2009); Metal-Metal TDA (Hallab, 2009).

Chapter ii.5.6  Orthopedic ­Applications Metal-on-polymer spine arthroplasty wear. In  vitro analysis of wear has demonstrated wear rates of metalon-polymer bearing lumbar total disc arthroplasty devices range from 2 to 20.8 mm3 per million cycles (Pare et al., 2007; Popoola et al., 2007), where the size of the wear debris generally ranges from 0.1 µm to 100 microns in diameter UHMWPE (Anderson et al., 2003; Popoola et al., 2007; van Ooij et al., 2007; Hallab et al., 2008). This amount of debris is 10-fold less wear than THAs and TKAs that are composed of metal on highly cross-linked polyethylene (x-UHMWPE) bearing surfaces (Greenwald and Garino, 2001; Kurtz et al., 2005). Metal-on-metal spinal arthroplasty wear. In general the wear of metal-on-metal TJA is well below that of metal-on-polymer (Figure II.5.6.13) (Wroblewski et al., 1996; Saikko et  al., 2002; Tipper et  al., 2002; Catelas et al., 2004; Heisel et al., 2004; Kurtz et al., 2005; Minoda et  al., 2005; Callaghan et  al., 2007). The few published reports on the volumetric wear rates of metalon-metal disc arthroplasty prostheses indicate a wear rate of 0.93–1.26 mm3 per million cycles (cobalt alloy) (Firkins et  al., 2001; Pare et  al., 2007). Another study of an all titanium-6%Al-4%V alloy disc arthroplasty found wear rates to be as high as 3 mm3 per million cycles (Hellier et  al., 1992). These values are similar to those reported for metal-on-metal hip replacements, which have been shown to range from approximately 0.05 to 6 mm3 per million cycles (cobalt alloy) (McKellop et al., 1996; Clarke et al., 2000; Catelas et al., 2004). Long-term follow-up of patients undergoing total disc arthroplasty is required to assess how intimately wear will correlate with inflammation and poor implant performance.

Orthopedic Biomaterial Corrosion Electrochemical corrosion occurs to some extent on all metallic surfaces including implants. This is undesirable for two primary reasons: (1) the degradative process may reduce structural integrity of the implant; and (2) the release of products of degradation is potentially toxic to the host. Metallic biomaterial degradation may result from electrochemical dissolution phenomena or wear, but most commonly occurs through a synergistic combination of the two. Electrochemical processes include generalized corrosion uniformly affecting an entire surface, and localized corrosion affecting either areas of a device relatively shielded from the environment (crevice corrosion) or seemingly random sites on the surface (pitting corrosion). Additionally, these electrochemical and other mechanical processes interact, potentially causing premature structural failure and/or accelerated metal release (e.g., stress corrosion cracking, corrosion fatigue, and fretting corrosion) (Brown and Merritt, 1981; Cook et  al., 1983; Bundy et  al., 1991; Brown et  al., 1992; Collier et  al., 1992b; Gilbert and Jacobs, 1997).

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Corrosion Mechanisms Corrosion of orthopedic biomaterials is a multifactorial phenomenon, and is dependent on five primary factors: (1) geometric variables (e.g., taper crevices in modular component hip prostheses); (2) metallurgical variables (e.g., surface microstructure, oxide structure and composition); (3) mechanical variables (e.g., stress and/or relative motion); (4) solution variables (e.g., pH, solution proteins, enzymes); and (5) the mechanical loading environment (e.g., degree of movement, contact forces, etc.). Current investigational efforts to minimize the corrosion of orthopedic biomaterials deal directly with the complex interactions of these factors. There are two essential features associated with how and why a metal corrodes. The first has to do with thermodynamic driving forces, which cause corrosion (oxidation/reduction) reactions. In general, whether or not corrosion will take place under the conditions of interest depends on the chemical driving force (ΔG), and the charge separation. This separation contributes to what is known as the electrical double layer (Figure II.5.6.14) which creates an electrical potential across the metal– solution interface (much like a capacitor): ΔG = −nFΔE

where n is the valence of the ion, F is the Faraday constant (95,000 coulombs/mole electrons), and E is the voltage across the metal solution interface. This potential is a measure of the reactivity of the metals or the driving force for metal oxidation. Therefore, the more negative the potential of a metal in solution, the more reactive it will tend to be (i.e., the greater is ΔG for reduction). The second factor governing the corrosion process of metallic biomaterials is the kinetic barrier to corrosion (e.g., surface oxide layer). Kinetic barriers prevent corrosion not by energetic considerations, but by physically limiting the rate at which oxidation or reduction processes can take place. The well-known process of passivation or the formation of a metal–oxide passive film on a metal surface is one example of a kinetic limitation to corrosion. In general, kinetic barriers to corrosion prevent either the migration of metallic ions from the metal to the solution, the migration of anions from solution to metal, or the migration of electrons across the metal–solution interface. Passive oxide films are the most well-known forms of kinetic barriers in corrosion, but other kinetic barriers exist, including polymeric coatings (Gilbert and Jacobs, 1997; Jacobs et al., 1998a).

Passivating Oxide Films Most alloys used in orthopedic appliances rely on the formation of passive films to prevent significant oxidation from taking place. These films consist of metal oxides, which form spontaneously on the surface of the

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SECTION II.5  Applications of Biomaterials

FIGURE II.5.6.14  Schematic of the interface of a passivating alloy surface in contact with a biological environment. metal in such a way that they prevent further transport of metallic ions and/or electrons across the film. Passive films must have certain characteristics to be able to limit further oxidation. The films must be compact and fully cover the metal surface, they must have an atomic structure which limits the migration of ions and/or electrons across the metal–­oxide–solution interface, and they must be able to remain on the surface of these alloys even with the mechanical stress and abrasion which can be expected with orthopedic devices. Passivating oxide films, which spontaneously grow on the surface of many metals and alloys, have five primary structural and physical characteristics, which are particularly relevant to implant degradation processes: 1. First, these oxide films are very thin, typically in the order of 5–70 Å which depends on the potential across the interface as well as solution variables

(e.g., pH). Furthermore, the oxide structure may be amorphous or crystalline. Since the potentials across the metal solution interface for these reactive metals are typically 1–2 volts, the electric field across the oxide is very high, in the order of 106–107 V/cm. One of the more widely accepted models is based on the theory of Mott and Cabrera, which states that oxide film growth depends on the electric field across the oxide. If the potential across the metal–oxide– solution interface is decreased (i.e., made closer to the electrochemical series potential), then the film thickness will decrease by reductive dissolution processes at the oxide. If the interfacial potential is made sufficiently negative or the pH of the solution is made low enough, then these oxide films will no longer be thermodynamically stable, and will undergo reductive dissolution without which corrosion will increase (Gilbert and Jacobs, 1997; Jacobs et al., 1998a).

Chapter ii.5.6  Orthopedic ­Applications 2. Second, oxide films have the characteristics of semiconductors with an atomic defect structure, which determines the ability for ionic and electronic transport across films. Metal cations and oxygen anions require the presence of cationic or anionic vacancies (respectively) in the oxide for transport of these species across the film. If there is a deficit of metal ions in the oxide film (i.e., there are cationic vacancies), for example, then metal ion transport is possible and these oxides are known as p-type semiconductors. Chromium oxide (Cr2O3) is such a metal-deficit oxide. On the other hand, if there is an excess of metal ions in the oxide (or a deficit of anions) then cation transport is limited, but anion transport can occur. These oxides will also have excess electrons and are known as n-type semiconductors. TiO2 spontaneously formed on titanium alloy implant (Ti-6Al-4V) surfaces is one such n-type semiconductor. The greater the number of defects (vacancies or other valence species), the less able is the oxide film to prevent migration of ionic species, and the lower is the kinetic barrier to corrosion. TiO2 is very close to being stoichiometric (chemically homogeneous) and hence does not have many ionic defects, resulting in an increased resistance to ionic transport. Other defects may be present in these passive oxide films which may alter their ability to limit corrosion. For instance, the addition of other metal ions with valence states which are different from the native metal ions can alter both the electronic and ionic transport of charge across the interface. These additions may enhance or degrade the ability of the oxide to prevent corrosion, depending on the nature of the oxide. One example of improved corrosion resistance from mixed oxides comes from what is known as a spinel. Spinels are typically mixed oxides of the form (A2O3)BO, where A and B are +3 and +2 valence metal ions. In Co–Cr alloys, for instance, a spinel of (Cr2O3)CoO can form on the surface. Spinels are typically known to have higher strengths and better resistance to diffusion of ions compared to single metal ion oxides. Therefore, a high concentration of spinels in the oxide layer will act to resist dissolution of a metal implant (Gilbert and Jacobs, 1997; Jacobs et al., 1998a). 3. Third, the ratio of the “oxide specific volume” to metal alloy specific volume (i.e., Pilling Bedworth ratio) will determine if the oxide will adhere to the metal or not. If there is too great a mismatch between the metal and oxide lattice parameters, then consequential stresses will be generated between the two. The magnitude of the internal stress will vary with the thickness of the oxide. Too great an oxide thickness will thus result in spontaneous fracture or spalling of the oxide, lowering the kinetic barrier effect of the oxide to corrosion (Gilbert and Jacobs, 1997; Jacobs et al., 1998a).

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4. Fourth, the morphology of these oxide films is not one of a smooth, flat, continuous sheet of adherent oxide covering the metal. Transmission electron microscopy (TEM) and atomic force microscopy (AFM) techniques have shown that oxides of titanium, for instance, consist of needle or dome shapes. The size and shape of these oxide domes change with applied potential when immersed in oxalic and other acids (Gilbert and Jacobs, 1997; Jacobs et al., 1998a). 5. Finally, mechanical factors such as fretting, micromotion or applied stresses may abrade or fracture oxide films. When an oxide film is ruptured from the metal substrate, fresh unoxidized metal is exposed to solution. When these films reform or repassivate, the magnitude of the repassivation currents which are subsequently generated may be large. This is because large driving forces exist for oxidation, and when the kinetic barrier is removed these large driving forces can operate to cause oxidation. However, the extent and duration of the oxidation currents will depend on the repassivation kinetics for oxide film formation. Hence, the mechanical stability of the oxide films, as well as the nature of their repassivation process, are central to the performance of oxide films in orthopedic applications (Gilbert and Jacobs, 1997; Jacobs et al., 1998a).

Corrosion at Modular Interfaces of Joint Replacements One issue associated with orthopedic alloys is the corrosion observed in the taper connections of retrieved modular joint replacement components (Figure II.5.6.15). With the growing number of total joint designs which use metal-on-metal conical tapers as modular connectors between components, the effects of crevices, stress, and motion take on increasing importance. Severe corrosion attack can take place in the crevices formed by these tapers in vivo. Gilbert et al. (1993) have shown that, of 148 retrieved implants, approximately 35% showed signs of moderate to severe corrosion attack in the head–neck taper connections of total hip prostheses. This attack was observed in components which consisted of Ti-6Al-4V alloy stems and Co–Cr heads, as well as Co–Cr stems on Co–Cr heads. This corrosion process is the result of a combination of stress and motion at the taper connection, and the crevice geometry of the taper; the stresses resulting from use cause fracturing and abrasion of the oxide film covering these passive metal surfaces. This, in turn, causes significant changes in the metal surface potential (makes it more negative) and in the crevice solution chemistry as the oxides continuously fracture and repassivate. These changes may result in deaeration (loss of O2) of the crevice solution, and a lowering of the pH in the crevice as is expected in crevice corrosion attack. The ultimate result of this process is a loss of the oxide film and its kinetic barrier effect, and an increase in

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SECTION II.5  Applications of Biomaterials

FIGURE II.5.6.15  Modular junction taper connection of a total hip arthroplasty showing corrosion of the taper connections. Macrograph of deposits of CrPO4 corrosion particles on the rim of a modular cobalt–chrome femoral head.

the rate of corrosive attack in the taper region. The corrosion processes in the Co–Cr alloys have been observed to consist of intergranular corrosion, etching, selective dissolution of cobalt, and the formation of Cr-rich particles. In isolated cases, this occurs to such an extent that intergranular corrosion caused fatigue failure in the neck of a Co–Cr stem. Corrosion attack of titanium alloy stems has also been observed in some cases. Very little is known about the mechanical stability of passive oxide films and the electrochemical reactions (e.g., ion and particle release) which occur when the oxide film is fractured. What is known is that when the oxide films of these orthopedic alloys are abraded or removed from the surface by rubbing, the open circuit potential can decrease to as low as −500 mV (versus Standard Calomel ­Electrode). These voltage potentials may be significant and prolonged enough to cause changes in the oxide structure and stability by bringing the interface potential into the active corrosion range of the alloy, thereby dramatically accelerating the corrosion rate and decreasing implant performance (Brown et al., 1992; Collier et al., 1992a; ­Gilbert et al., 1993; Bobyn et al., 1994; Brown et al., 1995).

Implant Debris Types: Particles and Ions The degradation products of all orthopedic implants are one of two basic types: particles or soluble debris (metal ions). While there is a large distinction between the two (ions versus particles), the difference between them blurs as the size of particles decreases into the nanometer range and becomes essentially “in solution.” Typically, particulate wear debris (metal, ceramic or polymer) exists from the submicron size up to thousands of microns (mm), while so-called “soluble debris” is limited to metal ions (or nanoparticles that are too small to be distinguished from ions) that are bound to plasma proteins. Particulate Debris.  Different types of joint arthroplasty designs not only produce different amounts and kinds of wear debris, but also different sizes and shapes of debris that are specific to the type of implant materials used for the bearing interfaces. For instance, hard-on-hard material couples such as metal-on-metal articulations generally produce smaller-sized (submicron), fairly round

debris, whereas traditional metal-on-polymer bearings produce larger (micron-sized) debris that is more elongated in shape (see Figure II.5.6.16). As is evident in Figure II.5.6.16, hard-on-hard material couples (e.g., metal-on-metal) produce smaller debris than do hard-on-soft material couples (e.g., metal on polyethylene). The particles produced from articulating bearing in any metal-on-polymer bearing implants are dominated by polymer particles, with little metallic debris unless there are other sources of metal release, such as corrosion at metal–metal connections. Polymeric particles produced from implants generally fall into the range from 0.23 to 1 μm (Figure II.5.6.16). Past investigations, primarily of UHMWPE wear debris in periimplant tissues, have shown that 70–90% of recovered particulates were submicron, with the mean size being approximately 0.5 μm (Maloney et  al., 1993; Jacobs et al., 1994a; Campbell et al., 1995). Metal and ceramic particles have generally been characterized as an order of magnitude smaller than polymer particles (at approximately 0.05 μm in diameter).

Histological Identification of Particles In Vivo The tissues surrounding modern implants may include areas of osseointegration, fibrous encapsulation, and a variable presence of the foreign-body response to polyethylene and cement debris in joint replacement devices. Absent is any specific histologic evidence of the slow release of metallic species that is known to occur with all metallic implants. However, accelerated corrosion and a tissue response that can be directly related to identifiable corrosion products can be demonstrated in the tissues surrounding multi-part devices (Urban et al., 2000). 1. Stainless steels: Histological sections of the tissues surrounding stainless steel internal fixation devices generally show encapsulation by a fibrous membrane, with little or no inflammation over most of the device. At screw plate junctions, however, the membranes often contain macrophages, foreign-body giant cells, and a variable number of lymphocytes in association

Chapter ii.5.6  Orthopedic ­Applications

865

FIGURE II.5.6.16  Implant debris from four types of materials are shown where the metal (cobalt alloy and titanium) and ceramic (alumina) debris are more rounded versus the polymeric (UHMWPE) debris which is more elongated in shape. Note: Bar = 5 µm.

with two types of corrosion products. The first consists of iron-containing granules. The second, termed microplates, consists of relatively larger particles of a chromium compound. Microplates are found within the tissues as closely packed, plate-like particle aggregates ranging in size from 0.5 mm to 5.0 mm. Hemosiderin–like granules often surround the collections of microplates, but the granules are also found alone. The granules are yellow–brown, mainly spherical, and 0.1–3 or more micrometers in diameter. They are predominantly intracellular, most often in macrophages, but may also be found in other periprosthetic cells (e.g., fibroblasts). X–ray diffraction has indicated that the granules consist of a mixture of two or more of the iron oxides, αFe2O3 and σFe2O3, and the hydrated iron oxides, αFe2O3.H2O and σFe2O3.H2O. 2. Cobalt–based alloys: The nature of corrosion at modular connections are similar, whether modular heads are mated with cobalt–chromium alloy or Ti–6Al–4V alloy femoral stems. The principal corrosion product identified by electron microprobe energy-­dispersive X–ray analysis and Fourier transform infrared microprobe spectroscopy is a chromium-phosphate

(Cr(PO4)4H2O) hydrate-rich material termed “orthophosphate.” This corrosion product can be found at the modular head–neck junction and as particles within the joint capsules, at the bone–implant interfacial membranes, and at sites of femoral osteolytic lesions. Particles of the orthophosphate material have been found at the bearing surface of the ­UHMWPE acetabular liners, suggesting their participation in three-body wear and an increased production of polyethylene debris. Particles of the chromium orthophosphate hydrate-rich corrosion product found in the tissues ranged in size from submicron to aggregates of particles up to 500 micrometers. 3. Titanium-based alloys: The degradation products observed in histologic sections of tissues adjacent to titanium-based alloys are of a different nature than the precipitates associated with stainless steel and cobalt-based alloys. Despite the remarkable corrosion resistance of titanium-based alloys, there have been persistent reports of tissue discoloration due to metallic debris in the periprosthetic tissues. These particulates observed in local tissues surrounding titanium alloy implants have the same elemental composition

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SECTION II.5  Applications of Biomaterials

FIGURE II.5.6.17  Transmission Electron Photomicrographs: (a) Macrophage containing phagocytized titanium particles. (b) Endothelial cell lining with embedded titanium debris. These specimens were obtained from a tissue sample overlying the posterolateral fusion mass (sixteenweek autograft + titanium) (TEM magnification = 20,000 ×).

as the parent alloy, as opposed to precipitated corrosion products which occur with stainless steel and cobalt–chromium alloys (Figure II.5.6.17). However, wear debris presents an enormous surface area for electrochemical dissolution, which, in all likelihood, is a major factor contributing to observed systemic elevations in titanium of patients with titanium implants (Urban et al., 1996a, 1997).

Particle Characterization Traditionally particle characterization uses methods such as Scanning Electron Microscopy (SEM) or Transmission Electron Microscopy (TEM), both of which are

number-based counting methods. These methods have indicated that the majority of the wear (mass loss) from an implant is comprised of particles in the nanometer-tosubmicron range. This understanding stems from the relatively low numbers of particles, (e.g., 100s–1000s) that are counted using image-based analysis techniques such as SEM. Newer analytical techniques, such as low angle laser diffraction (LALLS) have the capability of sampling millions to billions of particles, counted as they pass in front of and scatter a laser light beam proportionally to their size. Thus, as millions of particles flow by, LALLS analysis can detect the one-in-a-million large particle that comprises a significant portion of the total mass loss (i.e., total debris). This brings up the confusing concept that particle size of

Chapter ii.5.6  Orthopedic ­Applications Number Distributions %PASS 100.0

Debris Sample A

Volume Distributions

%PASS 100.0

%CHAN 10.0

90.0

90.0

9.0

80.0

80.0

8.0

70.0

70.0

7.0

60.0

60.0

6.0

50.0

5.0

40.0

4.0

30.0

30.0

3.0

20.0

20.0

2.0

10.0

10.0

1.0

0.0 0.100

1.000

10.00 -Size (microns)-

0.0 0.100

1.000

Debris Sample B

%PASS 100.0

10.00

100.0

-Size (microns)- average size

0.0 1000

mv=88um

%PASS 100.0

%CHAN 20.0

90.0

90.0

18.0

80.0

80.0

16.0

70.0

70.0

14.0

60.0

60.0

12.0

50.0

10.0

40.0

8.0

30.0

30.0

6.0

20.0

20.0

4.0

10.0

10.0

2.0

0.0 0.100

867

1.000 average size mn=1.2um

10.00 -Size (microns)-

0.0 0.100

(a)

1.000

10.00 -Size (microns)average size (b) mv=7um

100.0

0.0 1000

FIGURE II.5.6.18  Analyses of: (a) volume; and (b) number distributions of two debris samples demonstrate that similar number distributions

can result from very different actual size distributions. Note: The x-axis is particle diameter and the y-axis is (i) percentage of total number of particles in each size range; and (ii) the percentage of total mass in each size range. (Courtesy of BioEngineering Solutions, Inc.)

any given distribution depends on the method of evaluation. There is no one particle size. For example, the average size of 500 marbles and 5 basketballs is approximately the size of the marbles on a ­number basis, and approximately the size of the basketball on a volume basis. Thus, the question “what is the average size of particles that comprises approximately 50% of the total volume of particles?” is another way of asking what is the mean size (or average size) on a volume basis. Asking “what is the average size of the particles that comprises 50% of the total number of particles?” is the average size based on number basis. The stark differences between a volume and number-based analysis of implant debris are shown in Figure II.5.6.18. When comparing the volume and number distributions in this figure, the dominant contribution of larger particles to the total debris mass (volume) is evident, yet it is insignificant compared to the total ­number of ­particles (Figure II.5.6.18). Thus, as can be seen in both the simple example of marbles and basketballs and in the LALLS analysis of implant debris particles in Figure II.5.6.18, volume-based analysis tends to represent the largest (most massive) particles, and number analysis tends to depict the size characteristics of the most numerous. The ability to comprehensively characterize implant debris is important to the new designs and bearing surfaces used in new spinal implants. This multi-analysis approach is necessary because a given amount of wear debris (weight loss from the implant) after a year of use could

be attributed to the loss of a relatively few large particles or hundreds of millions of small particles (e.g., approx 0.2 mm3 volume loss after a million cycles of use could be from approx 400 particles of 100 microns diameter or 400 million particles only 1 micron in diameter). The bias of techniques limited to number analysis is that very similar number-based distributions can look very different from a volume perspective. This phenomenon is illustrated in Figure II.5.6.18 when comparing samples A and B, where different samples of particles look like very different volume-based distributions, however they look like very similar number-based distributions. This shows how important it is to have both number- and volume-based distributions of the same particles to fully characterize the types of particles in the mix, if there is enough particulate mass (>0.05 mg of particles) for a LALLS-type analysis. Unfortunately, in implant debris analysis there is usually less than the 0.05 mg of debris required for obtaining an accurate volume distribution, and thus historically SEM and TEM analyses have been used to characterize debris.

Particulate Debris Reactivity Characterization Macrophages are immune cells that are involved in the phagocytosis of implant debris and the resulting inflammatory responses. Once debris is ingested by macrophages a host of biologic reactions can occur, such

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SECTION II.5  Applications of Biomaterials

as activation of T-cells through antigen presentation (­Hallab et al., 2001a), release of pro-inflammatory mediators (Glant and Jacobs, 1994; Shanbhag et  al., 1994; ­Ingham et al., 2000; Matthews et al., 2000; Cunningham et  al., 2002), cytotoxicity (Hallab et  al., 2005), DNA damage (Nagaya et  al., 1989; Savarino et  al., 2000; ­Hallab et al., 2005), and oxidative stress (Soloviev et al., 2005). Macrophage reactions to debris are responsible for mediating debris-induced inflammation that is the leading cause of implant loosening over time. Despite the large number of studies on metallic, polymeric, and ceramic particles’ effects on peri-implant cells (e.g., macrophages, fibroblasts, osteoblasts, and osteoclasts), there are surprisingly few guidelines on what type of debris is most deleterious or bioreactive. However, there are a few general particle characteristics on which local inflammation has been shown to depend: (1) particle load (which depends on both particle size and total volume); (2) aspect ratio (the shape of the particles); and (3) chemical reactivity (the chemical composition). Thus, theoretically, a bioreactivity index of particulate would take the mathematical form: Particle Pro − inflammatory Index ≈ KLoad (particle load) × KShape (aspect ratio) × KMaterial (material type)

(1)

where K’s are constants that depend on the testing environment, and particle load is a function of both particle size and total debris volume. 1. Greater particle load: (size and volume) increases inflammation. An inflammatory response in  vivo is proportional to the particle load (the concentration of phagocytosable particles per tissue volume, which is characterized by both the size and total volume) (González et al., 1996; Matthews et al., 2000). While this seems obvious, the ramifications and the conditions under which this remains true are important and not so obvious. If a given amount of debris (mass loss from an implant) is comprised of small diameter particles, there will be far greater numbers than if that same mass of debris was composed of larger diameter particles (Matthews et al., 2000). The degree to which equal numbers (doses) of large particles (e.g., 10 micron diameter) and smaller particles (1 micron diameter) induce a proinflammatory response has not been thoroughly investigated, and remains unknown. 2. Aspect ratio: elongated (fibers) particles are more proinflammatory than round particles (Laquerriere et al., 2003; Sieving et al., 2003). That fibers are more reactive than round debris was well-established over 30 years ago, with studies of asbestos fibers (Bruch, 1974). However, it remains unknown at what aspect ratio (aspect ratio: length/width) in the transition from round particles to fibers that elevated inflammation is generally initiated, and thus to date there is no “guideline” aspect ratio for implant debris particles to remain below.

3. More chemically reactive particles are more ­proinflammatory. There is a growing consensus that metal particles are more proinflammatory when compared to other materials such as polymers (­Ramachandran et al., 2006). However, this is not a unanimous opinion; others have concluded polymers are more proinflammatory than metals (von Knoch et  al., 2000), and other reports have shown no differences between similar metal and polymer particles (Baumann et  al., 2004, 2005, 2006). Despite these reports to the contrary, there is a growing consensus that metallic particles are more proinflammatory because they are capable of corroding and releasing ions that have been associated with hypersensitivity responses, cytotoxicity, and DNA damage (Hallab et al., 2001a, 2005; Caicedo et al., 2007).

Controversial Particle Characteristics Does particle size matter? Absolutely. In vitro inflammatory responses generally require particles <10 microns in diameter to be phagocytosed. Thus, to produce an in vitro inflammatory response, particles need to be less than 10 μm, that is, within a phagocytosable range. Purportedly, particles with a mean size of 0.2–10 μm are the most proinflammatory. Within this range there is no consensus as to which specific size(s) and/or dose of particles (particles/cell or particles/tissue volume) are maximally inflammatory (Shanbhag et al., 1995; Rader et al., 1999; Matthews et al., 2000; Ingram et al., 2002). The effects of bacteria products (or endotoxin) on implant debris particles is presumably important because it has been found in periprosthetic tissue of failed implants, even in the absence of clinical signs of infection (Nalepka et  al., 2006). Furthermore, the bacteria that are present at below clinically detectable levels have been shown to affect implant performance, because antibiotic-eluting bone cement and systemically administered antibiotics reportedly reduce the frequency of long-term failure, i.e., aseptic loosening (Espehaug et al., 1997).

Metal Ions (Soluble Debris) There is continuing concern regarding the release of chemically active metal ions which bind to proteins and remain in solution from which they can then disseminate into the surrounding tissues, bloodstream, and remote organs. Particulate metallic wear debris presents an enormous surface area for electrochemical dissolution, which, in all likelihood, is a major factor contributing to observed systemic elevations in metals of patients with titanium implants (Urban et  al., 1996b, 1997, 1998; Jacobs et al., 1998b; Urban et al., 2000). Normal human serum levels of prominent implant metals are approximately 1–10 ng/ml Al, 0.15 ng/ml Cr, <0.01 ng/ml V, 0.1–0.2 ng/ml Co, and <4.1 ng/ml Ti. Following total joint arthroplasty, levels of circulating metal have been

Chapter ii.5.6  Orthopedic ­Applications

869

TABLE II.5.6.5    Approximate Average Concentrations of Metal in Human Body Fluids With and Without Total Joint Replacements (Michel et al., 1984; Stulberg et al., 1994; Jacobs et al., 1998a,b, 1999,) (ng/ml or ppb) Fluid

Ti

Al

V

Co

Cr

Mo

Ni

Serum

Normal THA THA-F THA-F TKA TKA-F

2.7 4.4 8.1 8.1 3.2 135.6

2.2 2.4 2.2 2.2 1.9 3.7

<0.8 1.7 1.3 1.3 <0.8 0.9

0.18 0.2-0.6 * * * *

0.05–0.15 0.3 0.2 0.2 * *

* * * * * *

0.4–3.6 <9.1 * * * *

Urine

Normal TJA

<1.9 3.55

6.4 6.53

0.5 <0.4

* *

0.06 0.45

* *

* *

Synovial fluid Normal TJA

13 556

109 654

5 62

5 588

3 385

21 58

5 32

Joint capsule

Normal TJA TJA-F

723 1540 19173

951 2053 1277

122 288 1514

25 1203 821

133 651 3329

17 109 447

3996 2317 5789

Whole blood

Normal TJA

17 67

13 218

6 23

0.1–0.1.2 20

2.0–4.0 110

0.5–1.8 10

2.9–7.0 29

Normal: Subjects without any metallic prosthesis (not including dental). THA: Subjects with well-functioning total hip arthroplasty. THA-F: Subjects with a poorly-functioning total hip arthroplasty (needing surgical revision). TKA: Subjects with well-functioning total knee arthroplasty. TKA-F: Subjects with a poorly-functioning total knee arthroplasty (needing surgical revision). TJA: Subjects with well-functioning total joint arthroplasty. TJA-F: Subjects with a poorly-functioning total joint arthroplasty (needing surgical revision). *Not tested.

shown to increase (Table II.5.6.5). The values in this table show that following successful primary total joint replacement there are measurable elevations in serum and urine cobalt, chromium, and titanium. The clinical ramifications of metal ion release are discussed in the next few sections.

Local Tissue Effects of Wear and Corrosion Implant debris limits the long-term performance of total joint replacement by causing a local inflammatory response that leads to bone erosion and implant loosening. Normal bone maintenance relies in the balance of bone formation and bone resorption, which mainly involves the coordinated function of osteoblasts and osteoclasts. Thus, either a decrease in osteoblastic bone formation or an increase in osteoclastic bone resorption can result in net bone loss and osteolysis. Bone loss (i.e., osteolysis) around an implant is the primary concern associated with the local effects of orthopedic implant degradation. This osteolysis causing implant debris occurs through both wear and corrosion mechanisms. Osteolysis is observed either as diffuse cortical thinning or as a focal cyst-like lesion. It was initially thought that reaction to particulate polymethylmethacrylate (PMMA) bone cement produced osteolytic lesions based on histological studies demonstrating cement debris associated with macrophages, giant cells, and a vascular

granulation tissue. Recently, however, osteolysis has been recognized in association with loose and well-fixed uncemented implants, demonstrating that the absence of acrylic cement does not preclude the occurrence of osteolysis (Jacobs et al., 2001; Vermes et al., 2001a,b). Implant debris causes low grade inflammation that ultimately leads to implant failure. Exactly how this happens remains unclear. Over the past 40 years implant debris-induced inflammation has been characterized ad nauseam, where debris-induced localized inflammation is caused in large part by macrophages which upregulate NFκβ and secrete inflammatory cytokines like IL-1β, TNFα, IL-6, and IL-8 (Jacobs et  al., 2001) (­Figure II.5.6.19). Other anti-inflammatory cytokines such as IL-10 modulate the inflammatory process. Other factors involved with bone resorption include the enzymes responsible for catabolism of the organic component of bone. These include matrix metalloproteinases collagenase and stromelysin. Prostaglandins, in particular PGE2, are also known to be important intercellular messengers in the osteolytic cascade produced by implant debris. More recently, several mediators known to be involved in stimulation or inhibition of osteoclast differentiation and maturation, such as RANKL (also referred to as osteoclast differentiation factor) and osteoprotegerin, respectively, have been suggested as key factors in the development and progression of bone loss (osteolytic lesions) produced from implant debris. Over the past

870

SECTION II.5  Applications of Biomaterials

Macrophage Reactivity to Implant

Implant Debris Metal particles /ions Mo5+ 2+ Mo5+ Ni2+ Ni 2+ 3+ Cr Ni 5+ Mo Mo5+ Co2+

2+ Ni2+ Cr3+ + Ni2+Co + + + Ni2+ Mo5+Ni2+ + +

NADPH / ROS

Debris ?

?

Nalp3

INFLAMMASOME

ASC

Implant Debris

Caspase-1 IL-1β

Pro-IL-1β

NFκβ

TLR Others

LPS ?

IL-1β IL-1β, IL-6, IL-8, IL-10, IL-I RECEPTOR

IL-12, IL15, IL-18

Osteoclast

FIGURE II.5.6.19  This schematic shows the numerous pro-­ inflammatory mediators produced by peri-implant tissue and immune cells reacting to implant debris, which can negatively affect bone turnover. The proinflammatory cytokines IL-1, IL-6, and TNF-α are thought to be some of the most potent cytokines in this cascade of signaling. These cytokines produced by cells react to implant debris acting through a variety of pathways to negatively affect bone turnover.

30 years we have understood that these mediators act to promote inflammation that decreases bone remodeling and is associated with the pathogenesis of osteolysis. However, we are only beginning to understand how implant debris could actually induce this immune system response at the cellular level. Implant debris is typically sterile, relatively inert, and does not “look” like a pathogen in any molecularly recognizable way. How then can implant debris provoke an inflammatory response? That is, how do intracellular mechanisms sense and respond to sterile nonbiological challenge agents such as implant debris? This question has remained unanswered for the past 40 years, but recently progress has implicated the “inflammasome,” a danger signaling pathway (Figure II.5.6.20) (Caicedo et al., 2009). In 1996 the discovery of specific pattern recognition receptors (PRRs) in the membrane and cytosol of human immune cells, such as macrophages, identified toll-like receptors (Taguchi et al., 1996) and their role in recognizing specific bacterial glycoproteins, now called

Osteoblasts

IL-8, IL-6, MCP-1 RANKL

© N.J. Hallab

BONE 1. Decreased Bone Deposition 2. Increased Bone Resorption

FIGURE II.5.6.20  The inflammasome pathway within cells such

as macrophages has recently been reported to be central to implant debris-mediated proinflammatory reactivity. (Picture courtesy of ­BioEngineering Solutions, Inc.)

“pathogen-associated molecular patterns” or PAMPs. We now understand that these receptors to highly conserved pathogen-associated molecular patterns (­Mariathasan and Monack, 2007) include toll-like receptors (TLRs), mannose receptors (MR), and NOD-like receptors (NLRs) (Mariathasan and Monack, 2007). Upon pathogen/cell contact these PRRs initiate a downstream cascade of events that activate the cell and induce the secretion of proinflammatory cytokines, leading to a broader inflammatory response. In 2005 danger signal pathways were discovered where nonpathogenic-derived stimuli were found to activate immune cells, similarly to PAMPs. Key components in this pathway were named the “inflammasome,” and the activating stimuli were termed “danger-associated molecular patterns,” or DAMPs (Martinon et  al., 2006). The paradigm for immune system activation now includes reliance on specific receptors that recognize

Chapter ii.5.6  Orthopedic ­Applications both pathogen-associated molecular patterns (PAMPs) and danger-associated molecular patterns (DAMPs) (Medzhitov, 2008; Ting et al., 2008). The inflammasome complex of proteins were the first pathway to explain how cells transduce sterile, nonpathogen-derived stimuli (e.g., cell stress and cell necrosis), into an inflammatory response (Mariathasan et  al., 2004; Mariathasan and Monack, 2007). Nonpathogen-derived danger includes such nonbiological stimuli as UV light, particulate adjuvants present in modern vaccines (Dostert et al., 2008; Hornung et  al., 2008), and, as it turns out, implant debris (Caicedo et al., 2010). When the inflammasome pathway is activated it causes the release of IL-1β, IL-18, IL-33 and other cytokines. How this happens is as follows: Debris → Phagocytosis → Lysosome damage → ROS(reactive oxygen species) → Inflammasome(NALP3/ASC) → Caspase1 → IL-1ß (and other IL-1-family) cytokines (Figure II.5.6.20).

Once ingested by immune cells, DAMPs, such as asbestos and implant debris, etc., (i.e., macrophages) induce some degree of lysosomal destabilization. This causes an increase in NADPH (nicotinamide adenine dinucleotide phosphate oxidase), and an increase in reactive oxygen species (ROS). This is not surprising, given the protease and acid rich extreme environment inside lysosomes used to digest and breakdown ingested particles/ bacteria, etc. The release of these intracellular contents are sensed by the intracellular multi-protein inflammasome complex which is composed of NALP3 protein (NACHT-, LRR- and pyrin domain-containing protein 3), in association with ASC (apoptosis-associated specklike protein containing a CARD domain) (­Mariathasan and Monack, 2007; Petrilli et  al., 2007). Activation of the inflammasome (NALPs-ASC complex) leads to the cleavage of pro-caspase-1 into active caspase-1. Active caspase-1 is required for the processing and subsequent release of active proinflammatory cytokines such as IL-1β and IL-18 (and others) by cleaving intracellular

871

pro-IL-1β, pro-IL-18, etc., into their mature forms, IL-1β and IL-18, etc. How different implant debris can cause different immune responses through specific mechanisms, such as the inflammasome, remains unknown and is currently under study. These new understandings facilitate both direct targeting for drugs and can enhance diagnostic measurement for improving, measuring, and predicting when implant debris will result in loose total joint arthroplasties. Goldring et al. (1983) were among the first to describe the synovial-like character of the bone implant interface in patients with loose total hip replacements, and determine that the cells within the membrane have the capacity to produce large amounts of bone resorbing factors PGE2 and collagenase. However, since studies typically can only document the end stage of the loosening process, rather than the initiating processes, pharmacologic interventions have been limited. Osteolysis associated with total knee arthroplasty has been reported less frequently than that associated with total hip arthroplasty. It is unclear why. However, in addition to obvious factors such as implant/bone mechanical loading environments, other more subtle differential mechanisms of hip and knee wear, and differences in interfacial barriers to migration of debris have been postulated to account for this apparent disparity. Although polyethylene particles are generally recognized as the most prevalent particles in the periprosthetic milieu, metallic and ceramic particulate species are also present in variable amounts, and may have important repercussions. The bulk of this debris originates from the articular surface and has easy access to local bone. When present in sufficient amounts, particulates generated by wear, corrosion or a combination of these processes can induce the formation of an inflammatory, foreign-body granulation tissue with the ability to invade the bone–implant interface (Figure II.5.6.21). Localized osteolytic lesions in these areas are common, but their clinical significance is limited unless large granulomatous lesions develop.

FIGURE II.5.6.21  Photomicrograph (5 ×) of a section through an acetabular section of a femoral stem retrieved at autopsy, 89 months after implantation. Note the periprosthetic cavity surrounded development of a granuloma emanating from an unfilled screw hole.

872

SECTION II.5  Applications of Biomaterials

The common observance of particle-induced osteolysis remote from the articulation surfaces has shown there is substantial particle migration between the joint space and the distal regions of the THA implant space. Autopsy specimens of retrieved implants have demonstrated the presence of connective tissue macrophages (histiocytes) in cavities surrounding regions of the femoral component. While the overall incidence of femoral osteolysis associated with THA tends to be proximal in the initial stages, over time it tends to progress distally. The volume of debris generated from THA polyethylene is related to a number of variables, including the smoothness of the concave metallic surface of the acetabular component, the tolerance between polyethylene and metal shell, and the relative stability of the insert (LaBerge et al., 1998; Shanbhag et al., 1998; Wimmer et al., 1998).

associated with anemia, osteomalacia, and neurological dysfunction, possibly including Alzheimer’s disease. However, when considering the variety of documented toxicities of these elements, it is important to keep in mind that the toxicities generally apply to soluble forms of these elements, and may not apply to the chemical species that result from prosthetic implant degradation. At this time, the association of metal release from orthopedic implants with any metabolic, bacteriologic, immunologic or carcinogenic toxicity is conjectural, since cause and effect have not been well-established in human subjects. However, this is due in large part to the difficulty of observation, in that most symptoms attributable to systemic and remote toxicity can be expected to occur in any population of orthopedic patients (Jacobs et al., 1999b).

Remote and Systemic Effects of Wear and Corrosion

Metal Ion Release

Implant surfaces and wear debris generated from the implant may release chemically active metal ions into the surrounding tissues. While these ions may stay bound to local tissues, there is an increasing recognition that released metal products bind to specific protein moieties, and are transported in the bloodstream and/or lymphatics to remote organs. The concern about the release and distribution of metallic degradation products is justified by the known potential toxicities of the elements used in modern orthopedic implant alloys: titanium; aluminum; vanadium; cobalt; chromium; and nickel. In general terms, metal toxicity may occur through: (1) metabolic alterations; (2) alterations in host/parasite interactions; (3) immunologic interactions of metal moieties by virtue of their ability to act as haptens (specific immunological activation) or anti-chemotactic agents (non-specific immunological suppression); and (4) by chemical carcinogenesis (Luckey and Venugopal, 1979; ­Beyersmann, 1994; Goering and Klaasen, 1995; Britton, 1996; Hartwig, 1998). Cobalt, chromium, and possibly nickel and ­vanadium, are essential trace metals in that they are required for certain enzymatic reactions. In excessive amounts, however, these elements may be toxic. Excessive cobalt may lead to polycythemia, hypothyroidism, cardiomyopathy, and carcinogenesis. Excessive chromium can lead to nephropathy, hypersensitivity, and carcinogenesis. Nickel can lead to eczematous dermatitis, hypersensitivity, and carcinogenesis. Vanadium can lead to cardiac and renal dysfunction, and has been associated with hypertension and depressive psychosis. Biologically non-essential metallic elements also possess specific toxicities. Titanium, although generally regarded as inert, has been associated with pulmonary disease in patients with occupational exposure, and with platelet suppression in animal models. Aluminum toxicity is well-documented in renal failure, and has been

In the long clinical experience of permanent and temporary metallic implants there has always been concern with local tissue reactions. There is a considerable literature concerning serum and urine chromium (Cr), cobalt (Co), and nickel (Ni) levels following total joint replacement, but relatively fewer studies examining titanium (Ti), aluminum (Al), and vanadium (V) levels. Many investigations have been hampered by technical limitations of the analytical instrumentation. Normal human serum levels of prominent implant metals are approximately: 1–10 ng/ml Al; 0.15 ng/ml Cr; <0.01 ng/ml V; 0.1–0.2 ng/ ml Co; and <4.1 ng/ml Ti. Following total joint arthroplasty, levels of circulating metal (Al, Cr, Co, Ni, Ti, and V) have been shown to increase (Table II.5.6.5). Multiple studies have demonstrated chronic elevations in serum and urine cobalt and chromium following successful primary total joint replacement. In addition, transient elevations of urine and serum nickel have been noted immediately following surgery. This hypernickelemia/hypernickeluria may be unrelated to the implant itself, since there is such a small percentage of nickel used in these implant alloys. Rather, this may be related to the use of stainless steel surgical instruments or the metabolic changes associated with the surgery itself. Chronic elevations in serum titanium and chromium concentrations are found in subjects with well-­ functioning titanium and/or chromium containing THR components without measurable differences in urine and serum aluminum concentrations. Vanadium concentrations have not been found greatly elevated in patients with TJA (Table II.5.6.5) (Michel et  al., 1984; Dorr et al., 1990; Jacobs et al., 1994c; Stulberg et al., 1994; Jacobs et al., 1998b). Metal ion levels within serum and urine of TJA patients can be affected by a variety of factors. For example, patients with total knee replacement components containing titanium-based alloy and carbon fiber reinforced polyethylene wear couples demonstrated

Chapter ii.5.6  Orthopedic ­Applications

873

TAB L E I I . 5 . 6 . 6    Concentrations of Metal in Body Tissue of Humans With and Without Total Joint Replacements (μg/g) Cr

Co

Ti

Al

V

Skeletal Muscle

Normal TJA

<12 570

<12 160

* *

* *

* *

Liver

Normal TJA

<14 1130

120 15200

100 560

890 680

14 22

Lung

Normal TJA

* *

* *

710 980

9830 8740

26 23

Spleen

Normal TJA

10 180

30 1600

70 1280

800 1070

<9 12

Psuedocapsule

Normal TJA

150 3820

50 5490

<65 39400

120 460

<9 121

Kidney

Normal TJA

<40 <40

30 60

* *

* *

* *

Lymphatic tissue

Normal

690

10

*

*

*

TJA

690

390

*

*

*

Normal TJA

30 90

30 280

* *

* *

* *

Heart

TJA: Subjects with a well-functioning total joint arthroplasty. * Not tested.

a 10-fold elevation in serum titanium concentrations at an average of four years after implantation. Up to a hundred times higher than normal control values of serum ­titanium elevations have also been reported in patients with failed metal-backed patellar components where unintended metal/metal articulation was possible. However, even among these TJA patients there was no elevation in serum or urine aluminum, serum or urine vanadium levels, or urine titanium levels. Mechanically assisted crevice corrosion in patients with modular femoral stems from total hip arthroplasty has been associated with elevations in serum cobalt and urine chromium. It has been previously assumed that extensively porous coated cementless stems would give rise to higher serum and urine chromium concentrations, due to the larger surface area available for passive dissolution. Recent studies suggest that disseminated chromium can predominantly come from fretting corrosion of the modular head–neck junction. However, wear of the articulating surface remains the purported predominant source of metallic implant debris (Jacobs et al., 1998a,b, 1999b). Homogenates of remote organs and tissue obtained postmortem from subjects with cobalt-based alloy total joint replacement components have indicated that significant increases in cobalt and chromium concentrations occur in the heart, liver, kidney, spleen, and lymphatic tissue (Table II.5.6.6). Similarly, patients with titaniumbased alloy implants demonstrated elevated titanium, aluminum, and vanadium levels in joint pseudocapsules (with up to 200 ppm of titanium six orders of magnitude greater than that of controls, 880 ppb of aluminum, and 250 ppb of vanadium). Spleen aluminum levels and liver

titanium concentrations can also be markedly elevated in patients with failed titanium-alloy implants (Jacobs et al., 1994c).

Systemic Particle Distribution Variables influencing accumulation of wear debris in remote organs are not clearly identified. When the magnitude of particulate debris generated by a prosthetic device is increased, it seems likely that a corresponding elevation in both the local and systemic burden of particles may be expected. Thus, component loosening, duration of implantation, and the modular designs of contemporary hip and knee replacement prostheses provide the potential for increased generation of metallic and polymeric debris (Figure II.5.6.22). Wear particles found disseminated beyond the periprosthetic tissue are primarily in the submicron size range. Numerous case reports document the presence of metallic, ceramic, or polymeric wear debris from hip and knee prostheses in regional and pelvic lymph nodes (Figure II.5.6.23), along with the findings of lymphadenopathy, gross pigmentation due to metallic debris, fibrosis (build-up of fibrous tissue), lymph node necrosis, and histiocytosis (abnormal function of tissue macrophages), including complete effacement of nodal architecture. The inflammatory response to metallic and polymeric debris in lymph nodes has been demonstrated to include immune activation of macrophages and associated production of cytokines. Metallic wear particles have been detected in the paraaortic lymph nodes in up to 70% of patients with total joint replacement components.

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Lymphatic transport is thought to be a major route for dissemination of wear debris. Wear particles may migrate via perivascular lymph channels as free or phagocytosed particles within macrophages. Within the abdominal para-aortic lymph nodes, the majority of disseminated particles are submicron in size; however, metallic particles as large as 50 micrometers, and polyethylene particles as large as 30 micrometers, have also been identified. These particles may further disseminate to the liver or spleen where they are found within

FIGURE II.5.6.22  Polarized light micrograph (190 ×) of para-aortic lymph node demonstrates the abundance and morphology of birefringent particles within macrophages. The larger filamentous particles were identified by infrared spectroscopy to be polyethylene.

macrophages or, in some cases, as epithelioid granulomas throughout the organs. Within liver and spleen, the maximum size of metallic wear particles are nearly an order of magnitude less than that in lymph nodes, indicating there may be additional stages of filtration preceding the lymphatic system or alternate routes of particle migration. In the liver and spleen, as in the lymph nodes, cells of the mononuclear–phagocyte system may accumulate small amounts of a variety of foreign materials without apparent clinical significance. However, accumulation of exogenous particles can induce granulomas or granulomatoid lesions in the liver and spleen (Figure II.5.6.23). It is likely that the inflammatory reaction to particles in the liver, spleen, and lymph nodes is modulated, as it is in other tissues by: (1) material composition; (2) the number of particles; (3) their rate of accumulation; (4) the duration that they are present; and (5) the biologic reactivity of cells to these particles. Metallic particles in the liver or spleen have been more prevalent in subjects with previously failed arthroplasties when compared with cases of well-functioning primary joint replacements. Metal particles, unlike polyethylene debris, can be characterized using an electron microprobe, which allows identification of individual, submicron metallic wear particles against a background of particulates from environmental or sources other than the prosthetic components. Overall, the smallest identifiable disseminated particles using the microprobe are approximately 0.1  micrometers in diameter. However, metallic wear debris may extend into the nanometer range, suggesting that additional methods of specimen preparation and analytic instrumentation may be required to more fully define the high burden of metallic wear particles in remote tissues (Urban et al., 2000).

FIGURE II.5.6.23  Epithelioid granulomas: (a) within the portal tract of the liver (40 ×); and (b) within the splenic parenchyma (15 ×) in a patient

with a failed titanium-alloy total hip replacement and symptomatic hepatitis. (c) Backscattered SEM of a granuloma in the spleen (3000 ×) demonstrating titanium alloy particles.

Chapter ii.5.6  Orthopedic ­Applications Polyethylene particles comprise a substantial fraction of the disseminated wear particles both in subjects with revision and primary TJAs. While the presence of these polyethylene particles in lymph nodes can be confirmed by Fourier Transform Infrared Spectroscopy microanalyses, polyethylene particulates in liver and spleen have so far precluded unequivocal identification. In these sites, the size of wear particles may be much smaller than 0.1 micrometers, making differentiation impossible by polarized light microscopy or infrared spectroscopy. Diseases which cause obstruction of lymph flow through lymph nodes, such as metastatic tumor, or which cause generalized disturbances of circulation, such as chronic heart disease or diabetes, may be expected to decrease particle migration to remote organs. Other diseases, such as acute or chronic-active inflammation in the periprosthetic tissues may increase particle migration (Urban et al., 1995; Jacobs et al., 1999b, 2001).

Hypersensitivity Some adverse responses to orthopedic biomaterials are subtle, and continue to foster debate and investigation. One of these responses is “metal allergy” or hypersensitivity to metallic biomaterials. Released ions, while not sensitizers on their own, can activate the immune system by forming complexes with native proteins. These metal–protein complexes are considered to be candidate antigens (or allergens) in human clinical applications. Polymeric wear debris is not easily chemically degraded in vivo and has not been implicated as sources of allergic type immune responses. This is presumably due to the relatively large degradation products associated with the mechanical wear of polymers in vivo, which may be large enough to prevent the formation of polymer–­protein haptenic complexes with human antibodies (­Hallab et al., 2000a,b, 2001a,b). Metal hypersensitivity is a well-established phenomenon. Moreover, dermal hypersensitivity to metal is common, affecting about 10–15% of the population. Dermal contact and ingestion of metals have been reported to cause immune reactions which most typically manifest as skin hives, eczema, redness, and itching. Although little is known about the short- and long-term pharmacodynamics and bioavailability of circulating metal degradation products in  vivo, there have been many reports of immunologic type responses temporally associated with implantation of metal components. Individual case reports link hypersensitivity immune reactions with adverse performance of metallic clinical cardiovascular, orthopedic, plastic surgical, and dental implants. Metals accepted as sensitizers (haptenic moieties in antigens) include beryllium, nickel, cobalt, and chromium, while occasional responses have been reported to tantalum, titanium, and vanadium. Nickel is the most

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common metal sensitizer in humans, followed by cobalt and chromium. Cross-sensitivity reactions between metals are common. Nickel and cobalt are, reportedly, the most frequently cross-reactive. Type IV Delayed Type Hypersensitivity (DTH) is a cell mediated type of response with which orthopedic implant associated hypersensitivity reactions (metal sensitivity or metal allergy) are generally associated. Metalantigen sensitized T-DTH lymphocytes release various cytokines which result in the accumulation and activation of macrophages. The majority of DTH participating cells are macrophages. Only 5% of the participating cells are antigen specific T lymphocytes (T-DTH cells), with a fully developed DTH response. The effector phase of a DTH response is initiated by contact of sensitized T cells with antigen. In this phase T cells, which are antigen-activated, are termed T-DTH cells and secrete a variety of cytokines that recruit and activate macrophages, monocytes, neutrophils, and other inflammatory cells. These released cytokines include IL-3 and GM-CF, which promote hematopoesis of granulocytes; monocyte chemotactic activating factor (MCAF) which promotes chemotaxis of monocytes toward areas of DTH activation; INF-γ and TNF-β which produce a number of effects on local endothelial cells facilitating infiltration; and migration inhibitory factor (MIF), which inhibits the migration of macrophages away from the site of a DTH reaction. Activation, infiltration, and eventual migration inhibition of macrophages is the final phase of the DTH response. Activated macrophages, because of their increased ability to present class II MHC (Major Histocompatibility Complex) and IL-2, can trigger the activation of more T-DTH cells, which in turn activates more macrophages, which activates more T-DTH cells, and so on. This DTH self-perpetuating response can create extensive tissue damage. The first apparent correlation of eczematous dermatitis to metallic orthopedic implants was reported in 1966 by Foussereau and Laugier (1966), where nickel was associated with hypersensitivity responses. Over the past 20 years, growing numbers of case reports link immunogenic reactions with adverse performance of metallic cardiovascular orthopedic, plastic surgical, and ­dental implants. In some instances clinical immunological symptoms have led directly to device removal. In these cases reactions of severe dermatitis (inflammation of the skin), urticaria (intensely sensitive and itching red round wheels on the skin), and/or vasculitis (patch inflammation of the walls of small blood vessels) have been linked with the relatively more general phenomena of metallosis (metallic staining of the surrounding tissue), excessive periprosthetic fibrosis, and muscular necrosis. The temporal and physical evidence leaves little doubt that the phenomenon of hypersensitivity to metal released from orthopedic implants does occur in some patients. These cases of severe metal sensitivity raise the greatest concern.

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Incidence of Hypersensitivity Responses Among Patients With Metal Implants The incidence of metal sensitivity among patients with both well- and poorly-functioning implants is roughly twice as high as that of the general population, approximately 25% (Figure II.5.6.24). Furthermore, the average incidence of metal sensitivity among patients with a “failed” implant (in need of revision surgery) is approximately 50–60% (Figure II.5.6.24). This is greater than five times the incidence of metal sensitivity found in the general population, and two to three times that of patients with metal implants. This increased prevalence of metal sensitivity among patients with a loose prosthesis has prompted the speculation that immunological processes may be a factor in implant loosening. Specific types of implants with greater propensity to release metal in  vivo may be more prone to induce metal sensitivity. Failures of total hip prostheses with metal-on-metal bearing surfaces were associated with greater incidence of metal allergy than similar designs with metal-on-ultra-high molecular weight polyethylene bearing surfaces. Alternatively, several published reports have indicated that after total joint replacement with metallic components some patients show an induction of metal tolerance; that is, previously diagnosed metal sensitivity abated after implantation of a metallic prosthetic. Additionally confounding to any clear connection between metal sensitivity and implant failure is the lack of

any reported correlation between incidence of metal sensitivity and implant residence time, infection, reason for removal or pain. This lack of causal evidence implicating cell-mediated immune responses has prompted some to conclude that implantation of cemented metal-to-plastic joint prosthesis is safe, even in the case of a pre-existing metal allergy. However, this is not a consensus opinion. At this time, however, it is unclear whether metal sensitivity causes implant loosening or whether implant loosening results in the development of metal sensitivity. The majority of investigations conclude that metal sensitivity can be a contributing factor to implant failure. Such cases include instances in which clinical immunological symptoms lead directly to the need for device removal. In these or similar cases there have been reported reactions of severe dermatitis, urticaria, and/or vasculitis, all presumably linked to what has been reported as metallosis, excessive periprosthetic fibrosis, and muscular necrosis. The clinical observation of apparent immune sensitivity to metallic implants is not limited to orthopedic surgery. Some case reports suggest metal sensitivity to pacemakers, heart valves, reconstructive, dental, and general surgical devices. The temporal and physical evidence associated with such cases leaves little doubt that the phenomenon of metal-induced hypersensitivity does occur in some cases, currently accepted within the orthopedic community to be <1% of patients. However, it is currently unclear whether metal sensitivity exists only as an unusual complication in a few susceptible patients or

FIGURE II.5.6.24  A compilation of investigations show the averaged percentages of metal sensitivity among the general population for nickel, cobalt, and chromium, among patients after receiving a metal-containing implant, and among patient populations with failed implants. All subjects were tested by means of a patch or metal-LTT (lymphocyte transformation test).

Chapter ii.5.6  Orthopedic ­Applications is a more subtle and common phenomenon, which over time plays a significant role in implant failure. It is likely that cases involving implant-related metal sensitivity have been under-reported, because alternate causes were attributed to the failure of the implant. Mechanisms by which in vivo metal sensitivity occurs have not been wellcharacterized. Thus, the degree to which a precondition of metal hypersensitivity may elicit an over-­aggressive immune response in a patient receiving an implant remains unpredictable. Continuing improvements in immunologic testing methods will likely enhance future assessment of patients susceptible to hypersensitivity responses (Hallab et al., 2000a,b, 2001a,b).

Carcinogenesis The carcinogenic potential of the metallic elements used in TJA remains an area of concern. Animal studies have documented the carcinogenic potential of orthopedic implant materials. Small increases in rat sarcomas were noted to correlate with high serum cobalt, chromium or nickel content from metal implants. Furthermore, lymphomas with bone involvement were also more common in rats with metallic implants. Implant site tumors in dogs and cats – primarily osteosarcoma and fibrosarcoma – have been associated with stainless steel internal fixation devices. Initially, epidemiological studies implicated cancer incidence in the first and second decades following total hip replacement. However, larger more recent studies have found no significant increase in leukemia or lymphoma; although these studies did not include as large a proportion of subjects with a metal-on-metal prosthesis. There are constituitive differences in the populations with and without implants that are independent of the implant itself, which confound the interpretation of epidemiological investigations. The association of metal release from orthopedic implants with carcinogenesis remains conjectural, since causality has not been definitely established in human subjects. The identification of such an association depends both on the availability of comparative epidemiology, and on the ability to perform tests on the patient before and after device removal. The actual number of cases of tumors associated with orthopedic implants is likely under-reported. However, with respect to the number of devices implanted on a yearly basis, the incidence of cancer at the site of implantation is relatively rare. Continued surveillance and longer-term epidemiological studies are required to fully address these issues (Gillespie et  al., 1988; Visuri and Koskenvuo, 1991; Matheisen et al., 1995; Nyren et al., 1995).

Preventive Strategies and Future Directions Current strategies designed to address the problem of biomaterial-related implant failure are primarily aimed

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at decreasing the amount of periprosthetic particulate burden and any subsequent effects. Recently there has been a great deal of innovation regarding stronger, more wear-resistant polyethylene. These more highly cross-linked UHMWPE polymers are currently in various phases of clinical trials. However, initial results show demonstrable decrease in polyethylene wear, with potential for less particulate-induced bioreactivity/osteolysis, and therefore greater implant performance. In the same vein, femoral heads with diameters of 32 mm have been associated with increased volumetric polyethylene wear; to combat this smaller 28 mm heads are currently extolled as more biocompatible. Manufacturing flaws, such as fusion defects and foreign-body inclusions, have also been suggested to contribute to adverse polyethylene wear properties. The elimination of polyethylene is another approach being investigated clinically in various centers. With the realization that early problems may have been related to the design and not the articulation, there has been a renewed interest in the application of metal–metal and ceramic–ceramic bearings. Future designs which attempt to reduce wear include: improved tolerances between polyethylene inserts and their metal backing; improved surface finish on the metallic concave surfaces; secure locking mechanisms; and the avoidance of holes on the convex portion of the acetabular prosthesis. Metallic wear is also being addressed through techniques such as nitriding and nitrogen ion implantation to decrease the potential for abrasive wear and fretting in titanium alloy and cobalt alloy stems. Fabrication of metallic bearing surfaces with extremely low roughness can be expected to decrease articular wear rates. A polished metal head can be made as smooth as a ceramic head. Polishing of the stem will remove surface asperities and decrease particle generation from stem/bone fretting. In addition, polishing will minimize silicate contamination. New metallic biomaterials are being proposed which attempt to improve load transfer to the bone and reduce the incidence of loosening and thigh pain. Currently used alloys (Co–Cr–Mo alloy, E = 227 GPa and Ti-6Al4V alloy, E = 115 GPa) have relatively high elastic moduli which limit smooth transfer of load to the surrounding bone in THA. Designs to improve load transfer can use a reduced cross-sectional area to increase flexibility, but at the expense of adequate stability of the implant within the bone. Additionally, the stresses may exceed the relatively low fatigue strength of Co–Cr–Mo implant alloy. Lower modulus, more corrosion-resistant implant alloys are being developed. A Ti-13Nb-13Zr (E = 79 GPa) alloy is one such alloy which contains fewer elements of questionable cell response (i.e., Co, Cr, Mo, Ni, Fe, Al, V), and which possesses comparable strength and toughness to existing Ti-6Al-4V implant alloy. The ­niobium and zirconiun constituents seek to improve bone ­biocompatibility and corrosion resistance.

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Additionally, novel surface treatments on implant alloys (such as the diffusion hardening (DH) treatment proposed for the Ti-13Nb-13Zr alloy), can produce a hardened surface with wear resistance superior to that of Co–Cr–Mo alloy, currently the industry leader. These enhanced surface properties may lead to an improvement in the resistance to micro-fretting occurring within femoral head–neck taper regions and modular interfaces of current implant designs. Electrochemical corrosion of orthopedic implants remains a significant clinical concern. Although the freely corroding implants used in the past have been replaced with modern corrosion resistant “superalloys,” deleterious corrosion processes have been observed in certain clinical settings. Attention to metallurgical processing variables, tolerances of modular connections, surface processing modalities, and appropriate material selection can diminish corrosion rates and minimize the potential for adverse clinical outcomes. For example, nitriding can reportedly significantly reduce the magnitude of fretting corrosion of Ti-6Al-4V devices. A need to further investigate the mechanical–electrochemical interactions of metal oxide surfaces in implants persists. Characterization of the stresses and motion needed to fracture passivating oxide films, as well as the effects of repeated oxide abrasion on the electrochemical behavior of the interface and ultimately the implant, remain avenues of active investigation. The clinical significance of elevated metal content in body fluids and remote organs of patients with metallic implants needs to be further elucidated. Considerably more work will be required to discern the specific chemical forms and distribution of metal degradation products associated with the various forms of implant degradation. Additionally uncharacterized is how these degradation products interact with proteins in vivo in terms of: (1) metal-ion/protein complexes; (2) nanometer-particle/protein complexes (ion-like particles); and (3) particle/protein-biofilm complexes. Although much has been revealed regarding the deleterious end effects of particulate debris (e.g., osteolysis) there remain very few ways of testing people to determine which type of implant and materials are right for them. Therefore, both an understanding of the constituents of orthopedic implant degradation and their biological effects is necessary to ultimately determine threshold levels of debris and circulating metal ions, and measures of biologic reactivity (e.g., metal-LTT testing) that can be used to clinically determine when intervention is required to fix a downward spiral before too much bone loss and inflammation occurs. The importance of this evaluation of orthopedic biomaterial performance is growing as the use of orthopedic biomaterials is increasing, as new orthopedic implants are being developed (Figure II.5.6.7), and as expectations of implant durability and performance increase (Black, 1996; Jacobs et al., 1996).

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Chapter ii.5.6  Orthopedic ­Applications Callaghan, J. J., Rosenberg, A., & Rubash, H. (2007). The Adult Hip. New York, NY: Lippincott Williams & Wilkins. Campbell, P., Ma, S., Yeom, B., McKellop, H., Schmalzried, T. P., et  al. (1995). Isolation of predominantly submicron-sized UHMWPE wear particles from periprosthetic tissues. J. Biomed. Mater. Res., 29(1), 127–131. Carlsson, A. S., Macnusson, B., & Moller, H. (1980). Metal sensitivity in patients with metal-to-plastic total hip arthroplasties. Acta. Orthop. Scand., 51, 57–62. Catelas, I., Medley, J. B., Campbell, P. A., Huk, O. L., & Bobyn, J. D. (2004). Comparison of in  vitro with in  vivo characteristics of wear particles from metal-metal hip implants. J. Biomed. Mater. Res. B. Appl. Biomater., 70(2), 167–178. Charnley, J. (1960). Anchorage of the femoral head prosthesis to the shaft of the femur. J. Bone Joint Surg. (Br.), 42, 28. Charnley, J. (1964). The bonding of prosthesis to bone by cement. J. Bone Joint Surg. (Br.), 46, 518. Charnley, J. (1979). Low Friction Arthroplasty of the Hip, Theory and Practice. Berlin: Springer-Verlag. Christiansen, K., Holmes, K., & Zilko, P. J. (1980). Metal sensitivity causing loosened joint protheses. Ann. Rheum. Dis., 39(5), 476–480. Clarke, I. C., Good, V., Williams, P., Schroeder, D., Anissian, L., et al. (2000). Ultra-low wear rates for rigid-on-rigid bearings in total hip replacements. Proc. Inst. Mech. Eng. (H), 214(4), 331–347. Collier, J. P., Mayor, M. B., Jensen, R. E., Surprenant, V. A., Surprenant, H. P., et  al. (1992a). Mechanisms of failure of modular prostheses. Clin. Orthop., 285, 129–139. Collier, J. P., Surprenant, V. A., Jensen, R. E., Mayor, M. B., & Surprenant, H. P. (1992b). Corrosion between the components of modular femoral hip prostheses. J. Bone Joint Surg. (Am.), 74-B, 511–517. Cook, S. D., Gianoli, G. J., Clemow, A. J., & Haddad, R. J. J. (1983). Fretting corrosion in orthopedic alloys. Biomater. Med. Devices Artif. Organs, 11(4), 281–292. Cowan, J. A., Jr., Dimick, J. B., Wainess, R., Upchurch, G. R., Jr., Chandler, W. F., et al. (2006). Changes in the utilization of spinal fusion in the United States. Neurosurgery, 59(1), 15–20. Cunningham, B. W., Orbegoso, C. M., Dmitriev, A. E., Hallab, N. J., Sefter, J. C., et al. (2002). The effect of titanium particulate on development and maintenance of a posterolateral spinal arthrodesis: An in vivo rabbit model. Spine, 27(18), 1971–1981. Deutman, R., Mulder, T. H., Brian, R., & Nater, J. P. (1977). Metal sensitivity before and after total hip arthroplasty. J. Bone Joint Surg[Am], 59-A, 862–865. Dorr, L. D., Bloebaum, R., Emmanual, J., & Meldrum, R. (1990). Histologic, biochemical and ion analysis of tissue and fluids retrieved during total hip arthroplasty. Clin. Orthop. Relat. Res., 261, 82–95. Dostert, C., Petrilli, V., Van, B. R., Steele, C., Mossman, B. T., et al. (2008). Innate immune activation through Nalp3 inflammasome sensing of asbestos and silica. Science, 320(5876), 674–677. Elves, M. W., Wilson, J. N., Scales, J. T., & Kemp, H. B. (1975). Incidence of metal sensitivity in patients with total joint replacements. British Medical Journal, 4, 376–378. Espehaug, B., Engesaeter, L. B., Vollset, S. E., Havelin, L. I., & Langeland, N. (1997). Antibiotic prophylaxis in total hip arthroplasty. Review of 10,905 primary cemented total hip replacements reported to the Norwegian arthroplasty register, 1987 to 1995. J. Bone Joint Surg. (Br.), 79(4), 590–595. Evans, E. M., Freeman, M. A., Miller, A. J., & Vernon-Roberts, B. (1974). Metal sensitivity as a cause of bone necrosis and loosening of the prosthesis in total joint replacement. The Journal of bone and Joint Surgery, 56-B, 626–642.

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Fernstrom, U. (1966). Arthroplasty with intercorporal endoprothesis in herniated disc and in painful disc. Acta. Chir. Scand. Suppl., 357, 154–159. Firkins, P. J., Tipper, J. L., Saadatzadeh, M. R., Ingham, E., Stone, M. H., et al. (2001). Quantitative analysis of wear and wear debris from metal-on-metal hip prostheses tested in a physiological hip joint simulator. Biomed. Mater. Eng., 11(2), 143–157. Foussereau, J., & Laugier, P. (1966). Allergic eczemas from metallic foriegn bodies. Trans. St John’s Hosp. Derm. Soc., 52, 220–225. Gilbert, J. L., & Jacobs, J. (1997). The mechanical and electrochemical processes associated with taper fretting crevice corrosion: A review. ASTM STP 1301 Modularity of Orthopedic Implants, (pp. 45–59) Philadelphia, PA: ASTM. Gilbert, J. L., Buckley, C. A., & Jacobs, J. J. (1993). In vivo corrosion of modular hip prosthesis components in mixed and similar metal combinations. The effect of crevice, stress, motion, and alloy coupling. J. Biomed. Mater. Res., 27(12), 1533–1544. Gillespie, W. J., Frampton, C. M., Henderson, R. J., & Ryan, P. M. (1988). The incidence of cancer following total hip replacement. J. Bone Joint Surg. (Br.), 70(4), 539–542. Glant, T. T., & Jacobs, J. J. (1994). Response of three murine macrophage populations to particulate debris: Bone resorption in organ cultures. J. Orthop. Res., 12, 720–732. Goering, P. L., & Klaasen, C. D. (1995). Hepatoxicity of Metals. New York, NY: Academic Press. Goldring, S. R., Schiller, A. L., Roelke, M., Rourke, C. M., O’Neill, D. A., et  al. (1983). The synovial-like membrane at the bone–cement interface in loose total hip replacements and its proposed role in bone lysis. J. Bone Joint Surg., 65A, 575–584. González, O., Smith, R. L., & Goodman, S. B. (1996). Efffect of size, concentration, surface area, and volume of polymethylmethacrylate paticles on human macrophages in  vitro. J. Biomed. Mater. Res., 30, 463–473. Goodman, S. B., Lind, M., Song, Y., & Smith, R. L. (1998). In  vitro, in  vivo, and tissue retrieval studies on particulate debris. Clin. Orthop., 352, 25–34. Granchi, D., Verri, E., Ciapetti, G., Stea, S., Savarino, L., et  al. (1998). Bone-resorbing cytokines in serum of patients with aseptic loosening of hip prostheses. J. Bone Joint Surg. (Br.), 80(5), 912–917. Greenwald, A. S., & Garino, J. P. (2001). Alternative bearing surfaces: The good, the bad, and the ugly. J. Bone Joint Surg. (Am.), 83-A(Suppl. 2 Pt 2), 68–72. Hallab, N. J., Mikecz, K., & Jacobs, J. J. (2000a). A triple assay technique for the evaluation of metal-induced, delayed-type hypersensitivity responses in patients with or receiving total joint arthroplasty. J. Biomed. Mater. Res., 53(5), 480–489. Hallab, N. J., Jacobs, J. J., Skipor, A., Black, J., Mikecz, K., et al. (2000b). Systemic metal-protein binding associated with total joint replacement arthroplasty. J. Biomed. Mater. Res., 49(3), 353–361. Hallab, N., Merritt, K., & Jacobs, J. J. (2001a). Metal sensitivity in patients with orthopedic implants. J. Bone Joint Surg. (Am.), 83-A(3), 428–436. Hallab, N. J., Mikecz, K., Vermes, C., Skipor, A., & Jacobs, J. J. (2001b). Differential lymphocyte reactivity to serum-derived metal-protein complexes produced from cobalt-based and titanium-based implant alloy degradation. J. Biomed. Mater. Res., 56(3), 427–436. Hallab, N. J., Anderson, S., Stafford, T., Skipor, A., Campbell, P., & Jocabs, J. J. (2004). Correlation between lymphocyte reactivity and metal ion levels in patients with metal-on-metal hip arthroplasty. Trans 50th Orthopaedic Research Society, 49. Hallab, N. J., Anderson, S., Caicedo, M., Brasher, A., Mikecz, K., et al. (2005). Effects of soluble metals on human peri-implant cells. J. Biomed. Mater. Res. A., 74(1), 124–140.

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Hallab, N. J., Khandha, A., Malcolmson, G., & Timm, J. P. (2008). In  vitro assessment of serum-saline ratios for fluid simulator testing of highly modular spinal implants with articulating surfaces. SAS Journal, 2(4), 171–183. Hallab, N. J. (2009). A review of the biologic effects of spine implant debris: Fact from Fiction. SAS journal, 3, 143–160. Hamby Wallace, B. (1959). Replacement of spinal intervertebral discs with locally polymerizing methyl methacrylate. J. Neurosurg., 16, 311–313. Harmon, P. H. (1963). Anterior excision and vertebral body fusion operation for intervertebral disc syndromes of the lower lumbar spine. Clin. Orthop., 26, 107–111. Harris, W. H. (1969). Traumatic arthritis of the hip after dislocation and acetabular fractures: Treatment by mold arthroplasty. An end-result study using a new method of result evaluation. J. Bone Joint Surg. (Am.), 51(4), 737–755. Harris, W. H. (1995). The problem is osteolysis. Clin. Orthop., 311, 46–53. Hartwig, A. (1998). Carcinogenicity of metal compounds: Possible role of DNA repair inhibition. Toxicol. Lett., 102–103, 235–239. Heisel, C., Silva, M., la Rosa, M. A., & Schmalzried, T. P. (2004). Short-term in vivo wear of cross-linked polyethylene. J. Bone Joint Surg. (Am.), 86-A(4), 748–751. Hellier, W. G., Hedman, T. P., & Kostuik, J. P. (1992). Wear studies for development of an intervertebral disc prosthesis. Spine, 17(Suppl. 6), S86–S96. Holt, G., Murnaghan, C., Reilly, J., & Meek, R. M. (2007). The biology of aseptic osteolysis. Clin. Orthop. Relat. Res., 460, 240–252. Hornung, V., Bauernfeind, F., Halle, A., Samstad, E. O., Kono, H., et al. (2008). Silica crystals and aluminum salts activate the NALP3 inflammasome through phagosomal destabilization. Nat. Immunol., 9(8), 847–856. Huk, O. L., Bansal, M., Betts, F., Rimnac, C. M., Lieberman, J. R., et al. (1994). Polyethylene and metal debris generated by nonarticulating surfaces of modular acetabular components. J. Bone Joint Surg. (Br.), 76(4), 568–574. Huo, M. H., Salvati, E. A., Lieberman, J. R., Betts, F., & Bansal, M. (1992). Metallic debris in femoral endosteolysis in failed cemented total hip arthroplasties. Clin. Orthop., 276, 157–168. Ingham, E., Green, T. R., Stone, M. H., Kowalski, R., Watkins, N., et  al. (2000). Production of TNF-alpha and bone resorbing activity by macrophages in response to different types of bone cement particles. Biomater., 21(10), 1005–1013. Ingram, J., Matthews, J. B., Tipper, J., Stone, M., Fisher, J., et al. (2002). Comparison of the biological activity of grade GUR 1120 and GUR 415HP UHMWPE wear debris. Biomed. Mater. Eng., 12(2), 177–188. Jacobs, J. J. (1995). Particulate wear. JAMA, 273, 1950–1956. Jacobs, J. J., & Hallab, N. J. (2006). Loosening and osteolysis associated with metal-on-metal bearings: A local effect of metal hypersensitivity? J. Bone Joint Surg. (Am.), 88(6), 1171–1172. Jacobs, J. J., Urban, R. M., Schajowicz, F., Gavrilovic, J., & Galante, J. O. (1992). Particulate-associated endosteal osteolysis in titanium-base alloy cementless total hip replacement. In: Particulate Debris from Medical Implants. Philadelphia, PA: American Society for Testing and Materials. Jacobs, J. J., Gilbert, J. L., & Urban, R. M. (1994a). Corrosion of metallic implants. In R. N. Stauffer (Ed.), Advances in Orthopedic Surgery (Vol. 2, pp. 279–319). St. Louis, IL: Mosby. Jacobs, J. J., Shanbhag, A., Glant, T. T., Black, J., & Galante, J. O. (1994b). Wear debris in total joints. J. Amer. Acad. Orthop. Surg., 2, 212–220. Jacobs, J. J., Skipor, A. K., Urban, R. M., Black, J., Manion, L. M., et  al. (1994c). Systemic distribution of metal degradation products from titanium alloy total hip replacements: An autopsy study. Trans. Orthop. Res. Soc., 838.

Jacobs, J. J., Skipor, A. K., Doorn, P. F., Campbell, P., Schmalzried, T. P., et al. (1996). Cobalt and chromium concentrations in patients with metal on metal total hip replacements. Clin. Orthop., 329(Suppl.), S256–S263. Jacobs, J. J., Gilbert, J. L., & Urban, R. M. (1998a). Corrosion of metal orthopedic implants. J. Bone Joint Surg. (Am.), 80(2), 268–282. Jacobs, J. J., Skipor, A. K., Patterson, L. M., Hallab, N. J., Paprosky, W. G., et al. (1998b). Metal release in patients who have had a primary total hip arthroplasty. A prospective, controlled, longitudinal study. J. Bone Joint Surg. (Am.), 80(10), 1447–1458. Jacobs, J., Goodman, S., Sumner, D. R., & Hallab, N. (1999a). Biologic response to orthopedic implants. In Orthopedic Basic Science (pp. 402–426). Chicago, IL: American Academy of Orthopedic Surgeons. Jacobs, J. J., Silverton, C., Hallab, N. J., Skipor, A. K., Patterson, L., et al. (1999b). Metal release and excretion from cementless titanium alloy total knee replacements. Clin. Orthop., 358, 173–180. Jacobs, J. J., Roebuck, K. A., Archibeck, M., Hallab, N. J., & Glant, T. T. (2001). Osteolysis: Basic science. Clin. Orthop., 393, 71–77. Jasty, M., Bragdon, C., Jiranek, W., Chandler, H., Maloney, W., et al. (1994). Etiology of osteolysis around porous-coated cementless total hip arthroplasties. Clin. Orthop., 308, 111–126. Jones, L. C., Frondoza, C., & Hungerford, D. S. (1999). Immunohistochemical evaluation of interface membranes from failed cemented and uncemented acetabular components. J. Biomed. Mater. Res., 48(6), 889–898. Katz, J. L. (1980a). Anisotropy of Young’s modulus of bone. Nature, 283(5742), 106–107. Katz, J. L. (1980b). The structure and biomechanics of bone. Symp. Soc. Exp. Biol., 34, 137–168. Korovessis, P., Petsinis, G., Repanti, M., & Repantis, T. (2006). Metallosis after contemporary metal-on-metal total hip arthroplasty. Five- to nine-year follow-up. J. Bone Joint Surg. (Am.), 88(6), 1183–1191. Kurtz, S. M., Hozack, W., Turner, J., Purtill, J., MacDonald, D., et al. (2005). Mechanical properties of retrieved highly crosslinked crossfire liners after short-term implantation. J. Arthroplasty, 20(7), 840–849. LaBerge, M. (1998). Wear. In J. Black, & M. C. Hastings (Eds.), Biomaterial Properties (pp. 364–405). London, UK: Chapman & Hall. Laquerriere, P., Grandjean-Laquerriere, A., Jallot, E., Balossier, G., Frayssinet, P., et  al. (2003). Importance of hydroxyapatite particles characteristics on cytokines production by human monocytes in vitro. Biomaterials, 24(16), 2739–2747. Liefeith, K., Hildebrand, G. & Schade, R. (2003). In  vitro and in  vivo evaluation of a fully resorbable calcium phosphate coatin deposited on TPS coated implants. Third International Essen Symposium on the Working Group on Biomaterials and Tissue Compatibility, Essen, 3. Luckey, T. D., & Venugopal, B. (1979). Metal Toxicity in Mammals. New York, NY: Plenum. Maloney, W. J., Smith, R. L., Castro, F., & Schurman, D. J. (1993). Fibroblast response to metallic debris in vitro. Enzyme induction cell proliferation and toxicity. J. Bone Joint Surg. (Am.), 75(6), 835–844. Mariathasan, S., & Monack, D. M. (2007). Inflammasome adaptors and sensors: Intracellular regulators of infection and inflammation. Nat. Rev. Immunol., 7(1), 31–40. Mariathasan, S., Newton, K., Monack, D. M., Vucic, D., French, D. M., et al. (2004). Differential activation of the inflammasome by caspase-1 adaptors ASC and Ipaf. Nature, 430(6996), 213–218.

Chapter ii.5.6  Orthopedic ­Applications Martinon, F., Petrilli, V., Mayor, A., Tardivel, A., & Tschopp, J. (2006). Gout-associated uric acid crystals activate the NALP3 inflammasome. Nature, 440(7081), 237–241. Matheisen, E. B., Ahlbom, A., Bermann, G., & Lindgren, J. U. (1995). Total hip replacement and cancer. J. Bone Joint Surg. (Br.), 77-B(3), 345–350. Matthews, J. B., Besong, A. A., Green, T. R., Stone, M. H., ­Wroblewski, B. M., et al. (2000). Evaluation of the response of primary human peripheral blood mononuclear phagocytes to challenge with in vitro generated clinically relevant UHMWPE particles of known size and dose. J. Biomed. Mater. Res., 52(2), 296–307. Mayor, M. B., Merritt, K., & Brown, S. A. (1980). Metal allergy and the surgical patient. The American Journal of Dermatogolgy, 139, 477–479. McKee, G. K., & Watson-Farrar, J. (1943). Replacement of the arthritic hips to the McKee-Farrar replacement. J. Bone Joint Surg. (Br.), 48, 245. McKellop, H., Park, S. H., Chiesa, R., Doorn, P., Lu, B., et  al. (1996). In vivo wear of three types of metal on metal hip prostheses during two decades of use. Clin. Orthop., 329(Suppl.), S128–S140. McKellop, H., Shen, F. W., Lu, B., Campbell, P., & Salovey, R. (2000). Effect of sterilization method and other modifications on the wear resistance of acetabular cups made of ultra-high molecular weight polyethylene. A hip-simulator study. J. Bone Joint Surg. (Am.), 82-A(12), 1708–1725. McKenzie, A. H. (1995). Fernström intervertebral disc arthroplasty: A long-term evaluation. Orthopedics International Edition, 3B, 313–324. Medzhitov, R. (2008). Origin and physiological roles of inflammation. Nature, 454(7203), 428–435. Merritt, K., & Brown, S. (1981). Metal sensitivity reactions to orthopedic implants. International Journal of Dermatology, 20, 89–94. Michel, R., Hoffman, J., Loer, F., & Zilkens, J. (1984). Trace element burdening of human tissue due to corrosion of hip-joint prostheses made of cobalt-chromium alloys. Arch. Orthop. Trama. Surg., 103, 85–95. Milavec-Puretic, V., Orlic, D., & Marusic, A. (1998). Sensitivity to metals in 40 patients with failed hip endoprosthesis. Arch Orthop Trauma Surg, 117(6–7), 383–386. Milosev, I., Trebse, R., Kovac, S., Cor, A., & Pisot, V. (2006). Survivorship and retrieval analysis of Sikomet metal-on-metal total hip replacements at a mean of seven years. J. Bone Joint Surg. (Am.), 88(6), 1173–1182. Minoda, Y., Kobayashi, A., Iwaki, H., Miyaguchi, M., Kadoya, Y., et al. (2005). Polyethylene wear particle generation in vivo in an alumina medial pivot total knee prosthesis. Biomater, 26(30), 6034–6040. Moore, A. T. (1943). Metal hip joint: A case report. J. Bone Joint Surg. (Am.), 25, 688. Munor-Ashman, D., & Miller, A. J. (1976). Rejection of metal to metal prosthesis and skin sensitivity to cobalt. Contact Dermatitis, 2, 65. Nagaya, T., Ishikawa, N., & Hata, H. (1989). Sister chromatid exchange analysis in lymphocytes of workers exposed to hexavalent chromium. Br. J. Ind. Med., 46(1), 48–51. Nalepka, J. L., Lee, M. J., Kraay, M. J., Marcus, R. E., Goldberg, V. M., et al. (2006). Lipopolysaccharide found in aseptic loosening of patients with inflammatory arthritis. Clin. Orthop. Relat. Res., 451, 229–235. Nyren, O., Mclaughlin, J. K., Anders-Ekbom, G. G., Johnell, O., & Fraumeni, A. H. (1995). Cancer risk after hip replacement with metal implants: A population-based cohort study in ­Sweden. Journal of the National Cancer Institute, 87, 28–33.

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Pare, P. E., Chan, F., Powell, M. L., & Mathews, H. H. (2007). Wear characterization of the MAVERICK total disc replacment. Trans 7th Annual Meeting Spine Arthroplasty Society (Berlin), 59. Petrilli, V., Dostert, C., Muruve, D. A., & Tschopp, J. (2007). The inflammasome: A danger sensing complex triggering innate immunity. Curr. Opin. Immunol., 19(6), 615–622. Popoola, O. O., Shen, M., Heller, M., & Seebeck, J. (2007). In vitro wear of UHMWPE inlays in Dynardi and Prodisc spine disc replacment implants. Trans 7th Annual Meeting Spine Arthroplasty Society (Berlin), 49. Rader, C. P., Sterner, T., Jakob, F., Schutze, N., & Eulert, J. (1999). Cytokine response of human macrophage-like cells after contact with polyethylene and pure titanium particles. J. Arthroplasty, 14(7), 840–848. Ramachandran, R., Goodman, S. B., & Smith, R. L. (2006). The effects of titanium and polymethylmethacrylate particles on osteoblast phenotypic stability. J. Biomed. Mater. Res. A., 77(3), 512–517. Ring, P. A. (1968). Complete replacement arthroplasty of the hip by the Ring prosthesis. J. Bone Joint Surg. (Br.), 50, 720. Saikko, V., Calonius, O., & Keranen, J. (2002). Wear of conventional and cross-linked ultra-high-molecular-weight polyethylene acetabular cups against polished and roughened CoCr femoral heads in a biaxial hip simulator. J. Biomed. Mater. Res., 63(6), 848–853. Savarino, L., Stea, S., Granchi, D., Visentin, M., Ciapetti, G., et  al. (2000). Sister chromatid exchanges and ion release in patients wearing fracture fixation devices. J. Biomed. Mater. Res., 50(1), 21–26. Schmalzried, T. P., Wessinger, S. J., Hill, G. E., & Harris, W. H. (1994). The Harris-Galante porous acetabular component press-fit without screw fixation. Five-year radiographic analysis of primary cases. J. Arthroplasty, 9(3), 235–242. Shanbhag, A. S., Jacobs, J. J., Black, J., Galante, J. O., & Glant, T. T. (1994). Macrophage/particle interactions. Effect of size, composition and surface area. J. Biomed. Mater. Res., 28, 81–90. Shanbhag, A. S., Jacobs, J. J., Black, J., Galante, J. O., & Glant, T. T. (1995). Human monocyte response to particulate biomaterials generated in vivo and in vitro. J. Orthop. Res., 13, 792–801. Shanbhag, A. S., Hasselman, C. T., Jacobs, J. J., & Rubash, H. E. (1998). Biological Response to Wear Debris. In J. J. ­Callaghan, A. G. Rosenberg, & H. Rubash (Eds.), The Adult Hip (pp. 279–288). Philadelphia, PA: Lippincott-Raven Publishers. Sieving, A., Wu, B., Mayton, L., Nasser, S., & Wooley, P. H. (2003). Morphological characteristics of total joint arthroplastyderived ultra-high molecular weight polyethylene (UHMWPE) wear debris that provoke inflammation in a murine model of inflammation. J. Biomed. Mater. Res., 64A(3), 457–464. Soloviev, A., Schwarz, E. M., Darowish, M., & O’Keefe, R. J. (2005). Sphingomyelinase mediates macrophage activation by titanium particles independent of phagocytosis: A role for free radicals, NFkappaB, and TNFalpha. J. Orthop. Res., 23(6), 1258–1265. Stillwell, W. T. (1987). The Art of Total Hip Arthroplasty. Orlando, FL: Grune & Stratton, Inc. Stulberg, B. N., Merritt, K., & Bauer, T. (1994). Metallic wear debris in metal-backed patellar failure. J. Biomed. Mat. Res. Applied Biomaterials, 5, 9–16. Taguchi, T., Mitcham, J. L., Dower, S. K., Sims, J. E., & Testa, J. R. (1996). Chromosomal localization of TIL, a gene encoding a protein related to the Drosophila transmembrane receptor Toll, to human chromosome 4p14. Genomics, 32(3), 486–488. Thompson, G. J., & Puleo, D. A. (1995). Effects of sublethal metal ion concentrations on osteogenic cells derived from bone marrow stromal cells. J. Appl. Biomater., 6(4), 249–258.

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Ting, J. P., Willingham, S. B., & Bergstralh, D. T. (2008). NLRs at the intersection of cell death and immunity. Nat. Rev. Immunol., 8(5), 372–379. Tipper, J. L., Hatton, A., Nevelos, J. E., Ingham, E., Doyle, C., et  al. (2002). Alumina-alumina artificial hip joints. Part II: Characterisation of the wear debris from in  vitro hip joint simulations. Biomaterials, 23(16), 3441–3448. Urban, R. M., Hall, D. J., Sapienza, C. I., Jacobs, J. J., Sumner, D. R., et  al. (1998). A comparative study of interface tissues in cemented vs. cementless total knee replacement tibial components retrieved at autopsy. Trans. SFB, 21. Urban, R. M., Jacobs, J. J., Tomlinson, M. J., Gavrilovic, J., & Andersen, M. (1995). Migration of Corrosion Products from the Modular Head Junction to the Polyethylene Bearing Surface and Interface Membranes of Hip Prostheses. New York, NY: Raven Press. Urban, R. M., Jacobs, J. J., Sumner, D. R., Peters, C. L., Voss, F. R., et  al. (1996a). The bone–implant interface of femoral stems with non-circumferential porous coating: A study of specimens retrieved at autopsy. J. Bone Joint Surg. (Am.), 78-A(7), 1068–1081. Urban, R. M., Jacobs, J. J., Tomlinson, M. J., Black, J., Turner, T. M., et al. (1996b). Particles of metal alloys and their corrosion products in the liver, spleen and para-aortic lymph nodes of patients with total hip replacement prosthesis. Orthop. Trans., 19, 1107–1108. Urban, R. M., Jacobs, J., Gilbert, J. L., Rice, S. B., Jasty, M., et al. (1997). Characterization of solid products of corrosion generated by modular-head femoral stems of different designs and materials. In D. E. Marlowe, J. E. Parr, & M. B. Mayor (Eds.), STP 1301 Modularity of Orthopedic Implants (pp. 33–44). Philadelphia, PA: ASTM. Urban, R. M., Jacobs, J. J., Tomlinson, M. J., Gavrilovic, J., Black, J., et al. (2000). Dissemination of wear particles to the liver, spleen, and abdominal lymph nodes of patients with hip or knee replacement. J. Bone Joint Surg. (Am.), 82(4), 457–476. van Ooij, A., Kurtz, S. M., Stessels, F., Noten, H., & van Rhijn, L. (2007). Polyethylene wear debris and long-term clinical failure of the Charite disc prosthesis: A study of 4 patients. Spine, 32(2), 223–229. Venable, C. S., Stuck, W. G., & Beach, A. (1937). The effects on bone of the presence of metals; based upon electrolysis. An experimental study. Annals of Surgery, 105, 917.

CHAPTER II.5.7  DENTAL IMPLANTATION Jack E. Lemons1 and Carl E. Misch2 1University

Professor, Schools of Dentistry, Medicine and Engineering, University of Alabama at Birmingham, Birmingham, AL, USA 2DDS, MDS, Misch International Institute, Beverly Hills, MI, USA

PATIENT PROFILES, DENTAL NEEDS, AND SURGICAL IMPLANTS: 1950S–2010S Functional, aesthetic, and general health compromises have been correlated with the loss of oral dentition. The dental profession has developed a wide range of treatments to deal with dentition losses and oral diseases; however, a significant percentage of the world population continues to lose teeth progressively with dental diseases and aging. In recent decades, since the 1950s, the modern era of treatments based on surgical implants

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has evolved. Significant advances in quality and quantity have occurred during each decade. In the USA, a larger population of completely edentulous individuals existed in the 1950s, and many implant treatments were initially designed to support full arch implant supported removable dentures (Misch, 1999). A prominent design was the subperiosteal type (Figure II.5.7.1A,B,C), where a cast cobalt alloy metallic framework was fabricated and implanted under the periosteum and fitted to surface features of the bone anatomy (Rizzo, 1988). Posts extended through the gingival and mucosal soft tissues and the implant denture was directly supported on bone without significant soft tissue contact. The surgical and implant fabrication procedures were technically demanding for the various subperiosteal implant designs. A group of dentists and supporting staff emerged as the experts in this subdiscipline. Early subperiosteal systems were shown to function through implant-to-soft tissue interfaces and many