Joint bearing surfaces and replacement joint design

8

Joint bearing surfaces and replacement joint design R L A P P A L A I N E N and M S E L E N I U S , University of Kuopio, Finland

8.1

Introduction

Joint bearing surfaces are critical for the painless movement of limbs. When they are severely damaged, eventually joints must be replaced by either tissue implants or artificial materials. The operation can normally restore basic joint movements and allow sufficient range of motion of the joint. In spite of a good outcome, the performance of artificial materials is poor compared with natural joints, and a large number of wear debris particles are released from the articulating surfaces. They irritate tissues and lead to aseptic loosening of the implant components. This chapter introduces some basics related to joint design, describes different artificial joint bearing materials and compares their performance. Although this discussion focuses on the hip joint and implants, similar concepts can be applied to other joints such as the knee, ankle, shoulder and elbow.

8.2

Articulating surfaces in natural joints

In natural healthy joints, the smoothness and elasticity of the cartilage-covered and bone-supported counterfaces and joint lubricants are essential for the absorption of the impact loads and the low friction movement of bones against each other. The excellent tribological properties of the human and animal joints are supposed to be due to a mixed mode lubrication, which includes pressure film lubrication and in particular contact point lubrication. Intact joint surfaces and high-quality joint lubricants seem to be essential for proper joint function. Normally, joint structure and function are maintained by proper use of joints for healthy living, including walking, bicycling and swimming, to mention a few of our favourite activities. Inactivity leads to atrophy, and overuse to an accelerated degeneration. It is not movement per se that maintains the joint architecture and physiology but it is the changes that these movements lead to in joint structures. These include interstitial fluid flow, which has been well recognised for its importance for joint nutrition. However, owing to water and ion movements this

Joint bearing surfaces and replacement joint design

177

also leads to associated electrical and electromagnetic phenomena, i.e. joint mechanics are closely related to electrodynamics. Owing to the extraordinary structure of joint articulating surfaces, joints are able to carry weights that are 8±12 times the body weight and glide with extremely low friction of 0.002±0.005 at 3000 kg/cm2 load (Charnley, 1959). Furthermore, even after stable loading with squeezing out of the liquid film between articulating surfaces, as the patient moves the joint, excellent lubrication is quickly achieved. From the engineering point of view this is an excellent achievement, since one of the best synthetic boundary lubricants, Teflon, can only provide lubrication with a friction coefficient at least 10 times higher than the natural joint, i.e. around 0.1 (Adamson, 1967).

8.3

Demands for the bearing surfaces

It is evident that the performance of natural joints is hard to achieve using synthetic materials. Therefore, conventionally, artificial joint bearing surfaces are designed to achieve two main goals, i.e. low wear rate and sufficiently low friction. Low wear rate is necessary to maintain wear debris generation at low level, i.e. below a threshold to avoid aseptic loosening owing to a large number of foreign particles from the sliding surfaces being released into the joint capsule and surrounding tissues. Low friction at the articulating surfaces guarantees low bending torque values on the fixation surfaces of implants, e.g. on the back surface of a hip joint cup. As demonstrated also in clinical use, the above criteria can be met by different combinations of basic material types, i.e. polymers, metals and ceramics. Friction and wear rate values for different kinds of combination are schematically illustrated in Fig. 8.1.

8.1 Comparison of friction and wear properties of different sliding pairs (modified from Schmidt et al., 1996).

178

Joint replacement technology

In principle, additional requirements for articulating surface materials include biocompatibility, good corrosion resistance in synovial fluid and degradation resistance in long-term use. However, these demands have been compromised since CoCr alloys are commonly used in metal components of sliding surfaces in spite of the fact that both Co and Cr are known to be toxic and cause allergic reactions. Ultra-high molecular weight polyethylene (UHMWPE) is also the most common plastic material for implant bearing surfaces, even though there is long-term degradation of UHMWPE components and fractures are common especially in knee implants.

8.4

Different solutions available

Two very relevant factors related to tribological performance of different sliding pairs are wear mechanisms and lubrication. There are five generally recognised wear mechanisms which can be summarised as follows: 1. Abrasive ± the displacement of materials by hard particles. 2. Adhesive ± the transference of material from one surface to another during relative motion by the process of solid-phase welding. 3. Fatigue ± the removal of materials as a result of cyclic stress variations. 4. Erosive ± the loss of material from a solid surface due to relative motion in contact with a fluid that contains solid particles. This is often subdivided into impingement erosion and abrasive erosion. If no solid particles are present, erosion can still take place, such as rain erosion and cavitation. 5. Corrosive ± a process in which chemical or electrochemical reactions with the environment dominates, such as oxidative wear. Consideration of these mechanisms is important to the discussion which follows with respect to UHMWPE as well as the other bearing surfaces. Lubrication refers to adding a lubricant between two bearing surfaces in order to control friction and wear. Change in the lubrication conditions can have a major impact on both friction and wear of total hip replacements. This is especially true in the case of polymers, where the wear rate may vary even by several orders of magnitude in two quite similar lubricants. Furthermore, the geometry, topography and surface chemistry of bearing surfaces determine how the lubricant operates and remains on the sliding surfaces. Three alternatives are presented in Fig. 8.2 of how the lubricant operates on the sliding surface. Furthermore, especially hard third-body wear debris particles can significantly increase scratching and wear of the sliding surfaces, especially in the case of metals.

8.4.1

Polymers on bearing surfaces

UHMWPE is the polymer that has been generally used as an implant material for over four decades. Sir John Charnley reported his clinical experience with total

Joint bearing surfaces and replacement joint design

179

8.2 Schematic of lubrication regimes and associated characteristics (fluid film, boundary and mixed lubrication regimes, based on Jin, 2006).

hip replacement in 1961. Charnley's original low-friction arthroplasty was based on a polyethylene acetabular component and a metallic femoral component. Today the number of total hip replacements done worldwide each year is over 800 000, most of which still utilise polyethylene on one of the bearing surfaces. A property of UHMWPE that distinguishes it from other polymers is its structure with extremely long and highly entangled molecular chains, which

180

Joint replacement technology

makes it resistant to wear. This is one main reason why UHMWPE has been so successful on implant bearing surfaces. Furthermore, UHMWPE is biocompatible and stable in the body. Several factors affect the wear resistance of the finished polymer implant and some of them can be traced back to its processing. Different processing steps can be used to modify the characteristics which affect performance strongly and the wear of the artificial joint. Molecular weight and degree of crystallinity are among the most significant ones. Other important factors include the state of crosslinking and oxidation index. It has been proved that crosslinked UHMWPE functions better on the sliding surfaces than ordinary UHMWPE. The present manufacturing processes are carried out so that there will be as little oxygen as possible present at any stage. Also the sterilisation is performed by irradiating in inert gas or using ethylene oxide gas, minimising oxidative degradation. It has been possible to improve UHMWPE as a sliding surface but it has not been possible to prevent wear and release of particles. Although the present implants wear out considerably less, even a minor yield of particles can be extremely harmful. There is now considerable evidence implicating UHMWPE wear debris in osteolysis and loosening of prostheses. There is, therefore, much interest in understanding the wear mechanisms, so that the best solutions can be developed. The compressive yield strength of the UHMWPE is about 12 MPa. The maximum nominal contact stresses in the hip joint are in the range 3±6 MPa. It is known that the acetabular component can still deform elastically, although the surface layer may experience large plastic flow. It has been found clinically that three different contact conditions favour the occurrence of surface plastic flow. Firstly, one or both of the contact surfaces may be rough and this often results in contact stresses exceeding the yield strength at the asperities. The second case follows when third-body abrasive particles remain stuck between the sliding surfaces. In many examples, the particles come from the bone cement that is used to fix implants. Bone cement contains hard ceramic particles, such as zirconium oxide or barium sulphate. In the third poorly understood case, both sliding surfaces are extremely smooth and no particles appear between the sliding surfaces. It has been perceived clinically that in some cases there are high wear rates (10±6 mm3/N m) although the sliding surfaces of the implant can be extremely smooth (Atkinson, 1985; Hall, 1996).

8.4.2 Metal±metal bearing surfaces Oil-lubricated metal bearings are common in machinery and they were also adopted in the early phase of the development of the modern THR. These first generation metal-on-metal bearings were based on stainless steel (McKee, 1951; Wiles, 1957). In the 1960s and early 1970s, the metal-on-metal design by

Joint bearing surfaces and replacement joint design

181

McKee and Farrar became popular and successfully competed with the Charnley-type metal-on-polyethylene designs. These prostheses were made of CoCrMo alloy and had head diameters of 32±42 mm. Poor fit between head and socket resulted quite often in high friction and wear although significant improvements were achieved with time. It was assumed that high friction values measured for the sliding pairs of the early McKee±Farrar joint caused failure of the bone cement fixation due to high torque on the cup. Later, simulator studies have confirmed that the wear rates are low and the loosening was more likely caused by mechanical factors. The bearing surface qualities and tribological properties of the later McKee prostheses were actually very good. Retrospective studies have shown that in clinical use wear was very low as long as the bearing components were properly matched with 0.15 to 0.20 mm clearance to allow a lubricating fluid interface. The first generation metal-on-metal bearings and prostheses were practically abandoned in the 1970s partly due to their problems and the better overall results achieved by the Charnley-type low friction implant which had bearing solutions based on polyethylene. Weber was one of the first to realise that in fact the low wear rates in metalon-metal THRs could be related to reduced loosening (Weber et al., 1989). His findings indicated that technically well-implanted first generation metal-onmetal prostheses usually had a very good clinical and radiological outcome. Based on these findings, Weber and his industrial partners started the development of the second generation metal-on-metal sliding pairs to optimise the clearance between head and liner, to enhance the low roughness and microstructure of CoCrMo alloy by using a wrought alloy instead of cast alloys, to develop optimal sphericity on the bearing and to use modern quality control. The resulting MetasulTM metal-on-metal THR prosthesis was introduced to clinical use in 1988 and approved by the US Food and Drug Administration (FDA) in 1999. Based on the promising outcome, most of the major implant manufacturers have a metal-on-metal THR system on the market. During the past decade, most hip implant manufacturers have introduced surface replacements having metal-on-metal bearing combinations with large head diameters of about 44±54 mm. Both the metal ball and the socket are manufactured with high precision tolerance. Better CoCrMo alloys and high precision on the bearing surfaces have made it possible to utilise these larger heads without increasing the wear rate significantly in spite of longer sliding distance with increasing head size. An additional benefit is the improved stability of the hip joint and reduced risk for dislocation, by as much as a factor of 10 compared with conventional implants. Furthermore, the reduced risk of impingement and point contact damage in rim contact during separation of the head and cup lead to decreased risk for severe damage to the bearing surfaces. The wear rates of bearing surfaces of metal-on-metal prostheses have been estimated both from retrieved implants and by using hip joint simulators. Generally, during an initial running-in phase, the wear rate is around 20 m/year

182

Joint replacement technology

corresponding to a volumetric wear rate of about 2 mm3/year. In the steady state, the wear rate is much lower, i.e. about 0.1±0.2 mm3/year, and remains about the same for forthcoming years. This might be due to the fact that the bearing surfaces are to a certain degree self polishing and tough. The size of metallic wear particles is generally below 100 nm and therefore they are too small to irritate the resident macrophages (Willert et al., 1996). In patients with metalon-metal THRs, an elevation of Co and Cr concentrations in blood serum and urine has been measured (Brodner et al., 1997). These concentrations seem to decrease after the initial one-year run-in phase. It has been thought that patients with chronic renal failure should not have a metal-on-metal THR implanted. Although hypersensitivity to the metals is theoretically possible, there is not much data on its clinical relevance (Willert et al., 2000). Furthermore Visuri and Koskenvuo (1991) showed that over a period of 15 years there was no elevated cancer risk in patients with McKee±Farrar type CoCrMo metal-on-metal THRs. As a conclusion the clinical experience accumulated from current and older metal-on-metal designs is extensive. Such bearings have been developed to be tough, nearly unbreakable and in the long run generate very little wear debris.

8.4.3

Ceramic±ceramic bearing surfaces

Owing to the extreme hardness of ceramics they should provide the lowest wear rate as sliding surfaces unless critical fractures occur. In hard sliding pairs, in addition to smoothness, an accurate fit of articulating surfaces is even more important than with pairs using polyethylene. Current manufacturing techniques can provide tolerances of bearing surface sphericity better than 1 m. These well-matched ball±cup pairs should allow hydrodynamic lubrication with a continuous fluid film. However, clinical surveys have indicated that articulating surfaces are partly in contact and that adhesive and abrasive wear occur (Streicher, 1995). Therefore, in identical sliding pairs, the materials should be as hard as possible to minimise wear. Aluminium oxide (alumina) and partially stabilised ZrO2 (zirconia) are the most used ceramics on bearing surfaces and they have been systematically studied since the early 1970s. The clinical performance of first generation ceramic±ceramic implant systems was poor owing to shortcomings in fixation, design and material properties. Sudden component fractures and catastrophic wear were quite common. Through a systematic development in material properties and proper designs, more consistent properties and wear performance have been achieved. In fact, a modern aluminium oxide ceramic head paired with an acetabular component of the same material has turned out to be reliable in clinical use and has been recommended especially for young and active patients (Sedel et al., 1994). Very low linear wear rates and low concentrations of wear particles in periprosthetic tissues have been confirmed by several clinical studies (e.g. BoÈhler et al., 2000). Aluminum oxide ceramics provide

Joint bearing surfaces and replacement joint design

183

superior performance in terms of wear resistance, bioinertness and little local cellular response to wear particles compared with other bearing couples (BoÈhler et al., 2000; Mochida et al., 2001). The wear properties including lubrication, friction and wear in alumina-onalumina bearings are excellent. Clinical data of wear rates for alumina±alumina pairs vary a lot for several reasons such as stability/instability of fixation, possible catastrophic failures and poor material grades and design. For wellfixed and stable implants the wear rates of 1±3 m/year are quite typical (Clarke and Willmann, 1994; Skinner, 1999). As mentioned above, good performance of alumina ceramics on sliding surfaces requires systematic development in material characteristics, especially a fine and uniform grain size to achieve high density, elimination of flaws and greater strength, by a manufacturing method called hot isostatic pressing (HIP). In addition to better fracture resistance, this kind of fine-grained material also wears less. In the case of ceramics, a single critical flaw can lead to complete failure of an implant. Therefore, 100% proof testing during the final inspection of finished ceramic implants has improved reliability since high stresses applied in testing would immediately propagate cracks from critical flaws, for example on the bearing surface, and lead to failure of the implant in testing. The criteria for such testing are based on empirical and theoretical data. Clinical fractures have decreased from around 10% in the 1980s to 0.004% with contemporary alumina implants meeting stringent quality criteria (Willman, 1999). Alumina wear particles as well as bulk alumina are very biocompatible with an average particle size below a micrometre. They are removed from the hard sliding surface and phagocytosed easily by macrophages in tissues, leading to a benign response. Nowadays the excellent tribological properties of alumina ceramic bearings can be utilised with several successful THR systems and designs. Current modular ceramic bearings are also clinically successful with good tribological properties, very low wear rate, and good local and systemic biocompatibility.

8.4.4

Coatings for bearing surfaces

In principle, coatings can be utilised on bearing surfaces of both hard±soft and hard±hard bearing couples. Several coating methods can be effectively used to modify implant bearing surface properties such as wettability (contact angle) or scratch and wear resistance. Although the typical average surface roughness of metal implant gliding surfaces initially is in the range 0.01±0.05 m, the roughness of these surfaces increases with time because of wear, corrosion and third body particles. This roughening can significantly increase the wear rate of metal and especially a polymer counterpart. Attempts to minimise polyethylene wear have been tried using two main approaches, namely, by optimising the properties of crosslinked polyethylene

184

Joint replacement technology

and by optimising the material properties of the counterface. In addition to bulk ceramics like alumina or zirconia, coatings such as TiN (Pappas et al., 1995), CrN, CrCN (Fisher et al., 2002), zirconium oxide (OxiniumÕ, Smith&Nephew Orthopaedics, Memphis, TN), alumina (Yen and Hsu, 2001) or amorphous diamond (Lappalainen et al., 1998, 2003; Affatato et al., 2000) have given some promising results. In principle, these coatings can be readily applied to metallic implant materials to provide a high strength ceramic outer layer. This ceramic layer can improve, for example, wettability of the implant, which is known to improve tribological performance of bearing surfaces in the case of sliding pairs with polyethylene. If the ceramic layer is thick enough, it can also protect the metal surface against third body particles (Santavirta et al., 1999). These might be ceramic components from the bone cement used for implant fixation, i.e. typically ZrO2 or BaSO4, which are two agents commonly used to make bone cement radiopaque. Ceramic coating may provide good improvement in corrosion resistance as well, if its adhesion and corrosion resistance on the metallic substrate are good on long-term exposure to body fluids. Unfortunately, these requirements have not been met in all the recently used clinical coated implants, leading to poor outcome in long-term clinical use. Based on the simulator experiments and clinical trials, ceramic surfaces can remain stable and minimise long-term PE wear with typically two to four times lower wear than with Co± Cr±Mo heads (McKellop, 1998; Oonishi et al., 1989, Schuller and Marti, 1990). On the other hand, coatings can give much worse results than uncoated CoCrMo against polyethylene (e.g., Jones et al., 2001). This is because a certain class of materials and coatings prepared using different methods and set-ups may have a large variation in characteristics and behaviour. Ceramic coatings have potential for hard sliding pairs due to good wear resistance. In this combination, in addition to smoothness, an accurate fit of bearing surfaces is even more important than with pairs using PE. However, high tolerances can be achieved using current manufacturing techniques. Theoretically, the well-matched ball±cup pairs should allow hydrodynamic lubrication with a continuous fluid film at the gliding interface. However, the clinical surveys have shown that articulating surfaces are partly in contact and that adhesive and abrasive wear occur (Streicher, 1995). Therefore, the coating material should be smooth and as hard as possible to minimise wear. Even in this case, ceramic coatings could offer several advantages over bulk ceramics. For example, because of extreme hardness and good tribological characteristics of diamond, continuous film lubrication is not needed in the case of amorphous diamond coatings. When the coating is thick enough (>20 m), it can withstand high contact stresses and the wear rate is negligible (less than 10 nm per 15 million cycles in a simulator) (Lappalainen et al., 2003). As shown in Fig. 8.3, ceramic particles of bone cement cannot damage the coating in simulator testing. In contrast, a CoCr head is easily damaged, leading to increasing wear of both the head and the cup. Furthermore, the coefficient of friction is generally fairly

Joint bearing surfaces and replacement joint design

185

8.3 Comparison of total wear (cup and ball) in a hip simulator test with bone cement third body particles for 2 million cycles in the middle of 5 million series. (MOM ˆ metal on metal; COC ˆ ceramic on ceramic: AD-on-AD ˆ amorphous diamond coating on amorphous diamond coating.)

low (<0.1) for ceramic-on-ceramic sliding pairs even in the early stages of an implant life cycle. Low friction is accompanied by low bending torque on fixation surfaces of the prostheses. However, the most important advantage of ceramic coatings compared with bulk ceramics is the fact that they are less prone to sudden complete failure, which is a feared, though rare (less than 0.1%), complication of current ceramic-on-ceramic total hip sliding pairs.

8.5

Special concepts and designs for bearing surfaces

Sometimes only one surface of a natural joint is damaged, e.g. the femoral head is damaged due to necrosis after a fracture of the femoral neck. In this case, the diseased femoral head can be replaced with an artificial ball of the same dimensions, typically 45±56 mm in diameter in a surgical procedure called hemiarthroplasty. The artificial head then articulates against the healthy cartilage in the socket. The ball may be rigidly fixed on a stem (monopolar) or a large ball may articulate freely on a smaller ball with a polyethylene liner between them (bipolar design). These designs lead to bearing surfaces consisting of natural cartilage and artificial implant material. In spite of the fact that hemiarthroplasty is a well-established treatment, several animal in vivo studies and human followup studies with metal implants have shown that cartilage was severely worn

186

Joint replacement technology

especially in long-term use (e.g. van der Meulen et al., 2002; Parsons et al., 2004). Evidently, better solutions would be welcome for this application, too. Since less invasive procedures are becoming more and more popular, different procedures have been developed, for example to replace part of the joint surface due to defect. These techniques include the use of tissue grafts, osteochondral grafts and cell-based techniques. In the best cases, these implants are well incorporated in the cartilage and functional even several months or years postoperatively.

8.6

Comparison of bearing surface solutions

One piece of evidence that different main bearing solutions, namely, highly cross-linked UHMWPE combined with metal or ceramic, CoCrMo±CoCrMo and alumina±alumina pairs, have their own merits and demerits is the fact that they are still in clinical use (Table 8.1). In spite of significant developments, none of them can provide all the advantages. On the other hand, they all have been clearly improved over the years and can offer THR longevity. Correct counterface design, proper material combinations, surface finish and tolerances are necessary to optimise tribological aspects. Based on long comparative simulator studies as well as implant registers, only minor differences in performance can be assumed on average. Figure 8.4 compares the typical annual penetration rates and friction values measured for different bearing combinations. However, some recommendations can be drawn based on the properties of different materials and clinical experiences. For example, if metal hyperTable 8.1 Comparison of strengths and weaknesses of different bearing materials Joint bearing

Strengths

Weaknesses

Metal-on-UHMWPE

· Many options · Toughness

· Extreme wear

Metal-on-crosslinked UHMWPE

· Many options · Toughness

· Prone to third-body wear

Ceramic-on-crosslinked UHMWPE

· Reduced wear · Abrasion resistance · Low friction

· Fracture risk · No head exchanges · Fewer sizes

Metal-on-metal

· Reduced wear · Head size options · Toughness

· High ion levels · Fewer liner options · Sensitive to abrasion

Ceramic-on-ceramic

· Reduced wear · Abrasion resistance · Low friction

· Fracture risk · Limited options · Revision challenges

Ceramic coatings

· All those listed above

· Delamination risk

Joint bearing surfaces and replacement joint design

187

8.4 Different combinations for bearing surfaces with their annual clinical penetration values and friction coefficients.

sensitivity might be a problem or retrieval (revision surgery) is carried out due to metallosis, UHMWPE combined with a ceramic head or totally ceramic-onceramic pair might be a better choice. On the other hand, if impingement or luxation are to be expected or retrieval (revision surgery) is due to ceramic failure, polyethylene cups or metal±metal combinations are more reliable, since they are not easily fractured by impingement or impacts on the rim section of the cup.

8.7

Future trends

Although several solutions for bearing couples already exist, some further developments can be expected, based on current know-how; these are listed next. One major advantage of better materials for bearing surfaces, for example highly crosslinked polyethylene and wear-resistant CoCrMo alloys by powder metallurgy, is the possibility of using larger head sizes in hip implants without significantly increasing wear debris release and the risk of osteolysis. Larger head sizes reduce the risk for impingement and luxation. Novel polymers such as polycarbonate urethane have shown potential for use as an acetabular bearing material (Khan et al., 2005). Polyurethane elastomers

188

Joint replacement technology

have a unique combination of durability, toughness and flexibility in comparison with UHMWPE, and may provide a good alternative after several years of systematic development. In spite of the serious shortcomings of pure quality coatings on bearing surfaces, coatings such as amorphous diamond or carbon nitride have the potential to give good long-term results as hard±hard bearing couples, if the coatings fulfil stringent requirements (Lappalainen and Santavirta, 2005). Perhaps the most important advantage of ceramic coatings compared with bulk ceramics is that they are less prone to sudden complete failure, which is a feared, if rare (less than 0.1%), complication of current ceramic-on-ceramic total hip sliding pairs.

8.8

References

Adamson AW, Physical Chemistry of Surfaces, New York, Wiley, 1967. Affatato S, Frigo M, Toni A, `An in vitro investigation of diamond-like carbon as a femoral head coating', J Biomed Mater Res, 2000 53 221±226. Atkinson JR, `Laboratory wear tests and clinical observations of the penetration of femoral heads into acetabular cups in total replacement hip joints', Wear, 1985 104 225±244. BoÈhler M, Mochida Y, Bauer TW, Salzer M, Plenk Jr H, `Characterization of wear debris from alumina-on-alumina THA', J Bone Joint Surg, 2000 82-B 901±909. Brodner W, Bitzan P, Meisinger V, Kaider A, Gottsauner-Wolf F, Kotz R, `Elevated serum cobalt with metal-on-metal articulating surfaces', J Bone Joint Surg, 1997 79-B 316±321. Charnley J, `Lubrication of animal joints', In Proceedings of I Mech E Conference on Biomechanics, London, 12, 1959. Clarke IC, Willmann G, `Structural ceramics in orthopedics', in Cameron HU, Bone Implant Interface, Mosby, St. Louis, 203±252, 1994. Fisher J, Hu XQ, Tipper JL, et al., `An in vitro study of the reduction in wear of metal-onmetal hip prostheses using surface-engineered femoral heads', Proc Inst Mech Eng H , 2002 216 219±230. Hall RM, `Wear in retrieved Charnley acetabular sockets', J Eng Med, 1996 210 197±207. Jin ZM, `Biotribology', Current Orthopaedics, 2006 20 32±40. Jones VC, Barton DC, Auger DD, Hardaker C, Stone MH, Fisher J, `Simulation of tibial counterface wear in mobile bearing knees with uncoated and ADLC coated surfaces', Biomed Mater Eng, 2001 11 105±115. Khan I, Smith N, Jones E et al., `Analysis and ealuation of a biomedical polycarbonate urethane teted in an in vitro study and an ovine arthroplasty model', Biomaterials, 2005 26 633±643. Lappalainen R, Santavirta S, `Potential of coatings in total hip replacements', Clin Orthop, 2005 430 72±79. Lappalainen R, Anttila A, Heinonen H, `Diamond coated total hip replacements', Clin Orthop, 1998 352 118±127. Lappalainen R, Selenius M, Anttila A et al., `Reduction of wear in total hip replacement prostheses by amorphous diamond coatings', J Biomed Mater Res, 2003 66 410± 413.

Joint bearing surfaces and replacement joint design

189

McKee GK, `Artificial hip joint', Proceedings of the East Anglian Orthopaedic Club, J Bone Joint Surg, 1951 33-B 465. McKellop H, `Assessment of wear of materials for artificial joints', in Callaghan J, Rosenberg A, Rubash H, The Adult Hip, New York, Lippincott-Raven, 231±246, 1998. Mochida Y, BoÈhler M, Salzer M, Bauer TW, `Debris from failed ceramic-on-ceramic and ceramic-on-polyethylene hip prostheses', Clin Orthop, 2001 389 113±125. Oonishi H, Igaki H, Takayama Y, `Comparison of wear of UHMWPE sliding against metal and alumina in total hip prostheses', Bioceramics, 1989 1 272±277. Pappas MJ, Makris G, Buechel FF, `Titanium nitride ceramic film against polyethylene. A 48 million cycle wear test', Clin Orthop 1995 317 64±70. Parsons IM, Millett PJ, Warner JJ, `Glenoid wear after shoulder hemiarthroplasty: quantitative radiagraphic analysis', Clin Orthop Relat Res 2004 421 120±125. Santavirta S, Lappalainen R, Pekko P, Antila A, Konttinen YT, `The counterface, surface smoothness, tolerances, and coatings in total joint prostheses', Clin Orthop, 1999 369 92±102. Schmidt M, Weber H, SchoÈn R, `Cobalt chromium molybdenum metal combination for modular hip prostheses', Clin Orthop, 1996 329 35±47. Schuller H, Marti R, `Ten-year socket wear in 66 hip arthroplasties. Ceramic versus metal heads', Acta Orthop Scand, 1990 61 240±243. Sedel L, Nizard RS, Kerboull L, Witvoet J, `Alumina±alumina hip replacement in patients younger than 50 years old', Clin Orthop,1994 298 175±183. Skinner HB, `Ceramic bearing surfaces', Clin Orthop Rel Res 1999 369 83±91. Streicher RM, `Tribology of artificial joints', in Morscher EW, Endoprosthetics, Berlin, Springer, 34±48, 1995. van der Meulen MCH, Beaupre GS, Smith RL et al., `Factors influencing changes in articular cartilage following hemiarthroplasty in sheep', J Orthop Res, 2002 20(4) 669±675. Visuri T, Koskenvuo M, `Cancer risk after McKee total hip replacement', Orthopedics 1991 14 137±142. Wiles P, `The surgery of the osteoarthritic hip', Br J Surg 1957 45 488±497. Willert HG, Buchhorn GH, GoÈbel D, KoÈster G, Schaffner S, Schenk R, Semlitsch M, `Wear behaviour and histopathology of classic cemented metal on metal hip prostheses', Clin Orthop 1996 329 160±186. Willert HG, Buchhorn GH, Fayyazi A, Lohmann CH, `Histopathologische VeraÈnderungen bei Metall/Metall Gelenken geben Hinweis auf eine zellvermittelte È berempfindlichkeit: vorlaÈufige Untersuchungsergebnisse von 14 FaÈllen', U Osteologie 2000 9 2±16. Willman G, `Ceramic ball head retrieval data', in Sedel L, Willmann G, Reliability and Long-Term Results of Ceramics in Orthopaedics, 4th International CeramTec Symposium, Stuttgart, Georg Thieme Verlag, 1999. Weber HG, Fiechter T, `PolyaÈthylen-Verschleiss und SpaÈtlockerung der Totalprothese des HuÈftgelenkes ± Neue Perspektiven fuÈr die Metall/Metall Paarung fuÈr Pfanne und Kugel', OrthopaÈde 1989 18 370±376. Yen SK, Hsu SW, `Electrolytic Al2O3 coating on Co-Cr-Mo implant alloys of hip prosthesis', J Biomed Mater Res 2001 54 412±418.

9

Cementless fixation techniques in joint replacement

M J C R O S S and J S P Y C H E R , The Australian Institute of Musculoskeletal Research, Australia

9.1

Introduction

Modern joint replacement without the use of cement demonstrates outstanding results and compels us to question the routine of using cement to improve the quality of patient care. Survivorship has been reported to be 96% at 13 years1 for cementless total knee arthroplasty and 97.5% for femoral stems in total hip arthroplasty after 20 years.2 These excellent results are possible by applying the main principles of cementless fixation in bone. These features include implant mechanics and design, rigorous surgical technique, knowledge of surface structure properties, bone ingrowth behaviour, and biological and pharmacological enhancement methods. Without doubt, the discovery and use of poly(methylmethacrylate) (PMMA) revolutionised the history of joint replacement and implant fixation in bone when it was introduced in the late 1950s. The credit for discovering the use of cement in orthopaedic surgery belongs to Sir John Charnley, who began his pioneering work at Wrightington in the United Kingdom.3±5 After some initial failures with materials (Teflon instead of polyethylene) the combination of bone cement and high-density polyethylene rapidly became the gold standard in hip replacement. It took a number of years before the US Food and Drug Administration (FDA) allowed the use of bone cement and once the procedure was set in motion, other joints beyond the hip were considered. In 1971 Gunston introduced the first cemented arthroplasty of the knee joint.6 The first shoulder, ankle and elbow prostheses were introduced and slowly the age of cement took a firm grip on the orthopaedic community and bestowed it with unprecedented success in arthroplasty. It was not until the early 1980s that numerous reports appeared in the literature documenting the drawbacks of cement.7±11 The main reasons for prosthesis failure were osteolysis and loosening of the implants. Cement disease became apparent in revision surgery. The two biggest problems were bone loss, and cement removal, particularly in femoral stems. The combination of cement

Cementless fixation techniques in joint replacement

191

debris and polyethylene particles was an accelerating factor in osteolysis. Recently it has been shown that polyethylene particles are the major factors in osteolysis.8 These failures prompted renewed interest in cementless fixation that focused on how to improve the qualities of implants to allow permanent stability in bone, without the use of an additional interface. Extensive research in this field merits the successful long-term outcomes of large clinical studies published in current literature. This in turn will continue to produce innovative discoveries in material technology and biological modulation. This chapter advances our knowledge of uncemented prostheses, showing that excellent long-term results can be achieved by understanding and applying the basic principles of cementless fixation.

9.2

Cementless fixation

The primary goal of cementless fixation is to improve the longevity of the implant. To achieve this objective certain fundamental principles must be applied and any neglect can lead to the implant's early failure. The three main principles that guarantee the best outcome of cementless implant fixation are: 1. sound initial stability; 2. osseous integration; 3. mechanical properties of the implant. Failure to follow these principles caused many of the early uncemented prostheses to fail. A good example is the early development of cementless total knee replacements. Early designs were insufficient in both mechanical and anatomical terms. An important factor in the design of tibial prostheses is peripheral cortical support and this requires a sufficient number of differing sizes to fit the varying anatomical configurations. The porous coated anatomic (PCA) knee replacement had an insufficient range of tibial sizes, allowing sinkage into the cancellous bone of the tibial metaphysis. Other issues associated with the failure of cementless designs had nothing to do with fixation methods. Rapid osteolysis due to particulate debris of early polyethylenes, insufficient polyethylene-locking mechanisms in hips and knees as well as flat articulating surfaces in knees occurred in both cemented and uncemented implants. These were failures related to design features and particularly the materials used. A classical example of inappropriate material in total knee prostheses was the Miller Galante I that had a titanium femoral component. As a bearing surface, titanium proved to be too soft. This caused premature wear and severe synovitic metallosis as well as rapid production of early polyethylene particulate debris. This in turn led to massive osteolysis and bone loss, making revision difficult.

192

9.3

Joint replacement technology

Initial stability

Rigid initial stability of a prosthesis is achieved when the forces required to dislodge the components are equal to or less than the physiological forces to which the implant is exposed upon postoperative loading. This involves an immediate tight fit in or on the bone, either by a three-dimensional press-fit configuration or by additional fixation using screws. The specific loading mechanics are different for each joint and depend on their anatomy and function. When a joint is subjected to compressive forces it is more amenable to cementless fixation. The most successful cementless implants in joint replacement have been the acetabular cup and the femoral stem (Fig. 9.1), followed by the distal femur. For primary total hip arthroplasty, the uncemented press fit acetabular cup has emerged as the most common standard of fixation. The compressive forces directed from the centre of rotation to the cup±bone interface result in an ideal mechanical equilibrium where shear and rotational forces are practically negligible. In this case the cup is press fitted (the reamed diameter is slightly smaller than the actual diameter of the cup). The proximal femur, however, is exposed to more complicated patterns of rotational, shear, tensile and compressive forces that place higher demands on implant design. Bourne et al.12 evaluated the stability of fixation, stress shielding and thigh pain of cementless stem designs. They concluded that the tapered stem showed better results than anatomical or cylindrical stems. A proximal femoral prosthesis must resist initial `sinking' (taper) and rotational forces (cross-section, in-built anteversion). In the knee it is more difficult to achieve stable initial fixation of the tibial component than of the femoral component. Resurfacing of the distal femur allows for a component, designed to be fixed in a press-fit fashion with the anterior and posterior bevels at an angle of 1±3ë. This creates a minimal wedge for the femoral epiphysis to fit into distributing forces over the whole component and avoids the problem of too much loading of the distal surface. The pattern of compressive forces along the tibial component during normal gait is complex. Due to rolling, sliding, toggling and rotation of the femur on the

9.1 Uncemented acetabular cup and femoral stem.

Cementless fixation techniques in joint replacement

193

tibia, application of an unsupported flat tray on the proximal tibial surface does not provide sufficient resistance to compressive failure of cancellous bone, lift off and micromotion. Miura et al.13 showed that the addition of a stem or screws are mandatory to achieve stable initial fixation. Surface roughness of the implant interface is an additional determinant to initial stability.14±16 The resistance to frictional motions is substantially increased by the presence of a rough surface. A minimum roughness of 3.5 m has proven favourable for the promotion of sufficient primary fixation and for growth stimulating micro-movements. After establishing the optimal design and instrumentation of a given prosthesis, the next step in gaining initial rigid fixation is careful surgical technique. Preoperative planning is key. Special attention should be given to deformities, bone loss and correcting malalignments. Instrumentation of the prosthesis must yield precise bone cuts, and bone moulding must accurately accommodate the implant. Undersizing the mould of an acetabular cup or femoral metaphysis in short-stem prosthesis significantly increases the stability of implant fixation.16±20 Special care must be given not to apply deforming forces on jigs or instruments that can jeopardise accuracy. For implants that rely on cortical support, e.g. the proximal tibia, efforts must be made to choose the ideal implant size that maximises cortical contact area without over sizing the component. Therefore, tibial trays need to have at least 10 sizes (personal experience of over 7000 cases). Various studies have shown that the compaction of cancellous bone increases the initial fixation strength of the implant up until four weeks postoperatively.21±23 This simple technical detail must be considered in instrument design and should be applied wherever possible, especially in proximal femora and tibiae. It is also important to note that if the periphery of the implant succeeds in producing complete bone ingrowth and in creating a watertight fit, then the transport of particulate debris to the bone±implant interface is rendered virtually impossible.24 After having performed the bone cuts, no attempt should be made to clean the surfaces of blood and its osteogenic marrow components until firm implantation of the prosthesis. The quality of bone influences initial stability.25 In cases of gross osteoporosity a finger pressure test can be carried out. A positive result indicates a possible contra-indication of an uncemented design. Please refer to Section 9.6 on `Why do you still use cement?' Only if rigid initial stability is realised can the crucial secondary phase of successful osseo-integration begin, which in the long run will determine the longevity of the fixation.26,27

9.4

Osseous integration of cementless implants

The direct intimate contact of bone tissue to the surface of an implant is unique to the technology of cementless fixation and is key to preventing the effects of

194

Joint replacement technology

implant loosening. The preference for cementless fixation is that osseointegration occurs rapidly; thereby minimising the risk that early forces will act on the implant causing displacement. The main aspects affecting the quality and rate of osseo-integration are: · · · · ·

implant stability, micromotion and interface gaps; surface geometry characteristics; biocompatibility of materials; bioactive surface coatings; physiology of osseo-integration.

9.4.1

Implant stability, micromotion and interface gaps

Excessive micromotion of an implant in bone renders bone ingrowth impossible. The amount of permitted minimal movement within an interface has been reported to be 28±150 m, and repetitive higher displacements allow only ingrowth of fibrous tissue.28,29 Micromotion is a function not just of primary implant stability but also of the differences in the elastic modulus of bone and in the implant material. In tibial trays this mismatch can produce motions of up to 150 m. Retrieved implants have shown best bone ingrowth near fixation pegs and inconsistent ingrowth patterns at the periphery.30 Even for the most talented surgeons, it remains impossible to produce an absolute conforming host site that shows no gaps over the entire surface area. Those areas that lack direct contact have a negative effect on osseo-integration. Numerous models with analysis of controlled gaps in stable implants reveal that bone ingrowth is reduced up to sixfold in the presence of a 2 mm gap as compared with direct bone contact, These results demonstrate the supreme importance of meticulous surgical technique.20,31 In our study of hydroxyapatite coated total knee replacements we proved that gaps up to 2 mm usually filled in after two years, but only if the implant was absolutely stable.32

9.4.2

Surface geometry characteristics

Numerous modifications of surface geometry were studied to determine optimal pore sizes and roughness. Porous surfaces are usually applied as a coating on the implant through a variety of techniques. The most common techniques involve the application of vacuum-sputter or plasma-sprayed coatings, flame spraying techniques or, alternatively, roughening the surfaces through grit blasting (Fig. 9.2).33 Friedman et al. have demonstrated higher shear strength for arc-deposited titanium (flame spraying technique), than for grit-blasted surfaces or sprayed beads.34 The ideal pore size that allows bone ingrowth has been extensively investigated, mostly in canine models.35±38 These studies indicate that the

Cementless fixation techniques in joint replacement

195

9.2 Beaded coating on tibial tray.

optimal range of pore sizes is between 100 and 400 m. Although most prostheses are coated with a porosity size within this range, retrieval analysis has shown that in tibial plateaux bone ingrowth is not routinely achieved. Further research led to cast mesh being applied to the surface of implants, which demonstrated superior fixation strength when compared to porous coatings.39,40

9.4.3

Biological compatibility of materials

Modern bone implant materials can be classified according to their biocompatibility into biotolerant, bioinert and bioactive materials.41 Biotolerant materials such as PMMA are characterised by a thin fibrous tissue interface, whereas bioinert materials such as titanium and aluminium oxide typically integrate well into bone. Bioactive materials, such as calcium phosphate ceramics and glass, demonstrate direct chemical bonding of the implant with bone. Wilke et al.42 proved that cell proliferation as a marker for osseo-integration was highest on hydroxyapatite surfaces, followed by titanium and chrome±cobalt±molybdenum alloy. Under transmission electron microscopy the implant±bone ultrastructure of commercially available pure titanium, Ti6Al4V, and cobalt±chromium alloy are comparable. In recent years, a new implant material has produced substantially higher volumetric bone ingrowth into implants. It possesses a much higher porosity ± up to 80%, compared with previous porosity of only 35±50% ± without sacrificing the structural integrity of the implant. This new implant consists of commercially pure tantalum that is deposited on a vitreous carbon skeleton by chemical vapour deposition/infiltration. This trabecular-like metal can be made

196

Joint replacement technology

into complex shapes and used either as a bulk implant or as a surface coating. It shows a remarkably high resistance to corrosion and excellent biocompatibility. Its high frictional properties make it conducive to biological fixation and its low modulus of elasticity allows for more physiological load transfer and relative preservation of bone stock. Most clinical publications to date report very promising results with this novel material.43±47 A radiographic analysis of 86 acetabular cups showed that all gaps up to 5 mm were filled after 24 weeks of implantation indicating the strong osteoconductive, possibly even osteoinductive properties of trabecular metal.48 An experimental study conducted by Bobyn et al.49 in a canine model proved bone ingrowth into pores averaging 430 m of up to 80% with histological evidence of Haversian remodelling at 52 weeks.

9.4.4

Bioactive surface coatings

The quest for improving the rate and amount of osseo-integration of an implant in bone led to the introduction of osteo-conductive surface chemistry and osteoinductive biomodulators on or in implant surfaces. The application of calcium phosphates, especially hydroxyapatite (HA), as an osteo-conductive mediator has significantly improved the quality of implant fixation. This has been demonstrated in numerous experimental and clinical studies. Cooley et al. investigated ingrowth of HA-coated titanium compared with pure titanium at three time periods. Mechanical evaluation showed significantly greater interface bond strength, and histological analysis revealed nearly twice the percentage of direct bone contact for coated implants.50 Soballe and coworkers51±53 showed that HA coatings in dogs significantly enhanced gap healing and, when exposed to polyethylene particles in unstable implants, provided a superior sealing effect at the periphery. The same study showed shorter healing time, higher tolerance to micromotion and improved loading anchorage in bone. Many later clinical studies have confirmed the superior results of HA-coated implants.54±59 In our review of 1000 consecutively performed total knee replacement surgeries that employed HA-coated stemless implants, we demonstrated a survivorship of 99% at 10 years with excellent clinical and functional outcome.60 In our collective of patients over 75 years of age, as well as younger very active patients, we observed a low rate of revision and infection.61,62 Especially in the younger age group, where previous studies have shown unsatisfactory mid- to long-term results for cemented knee replacement, we strongly advocate the use of a HA-coated, uncemented knee replacement (Fig. 9.3). We believe the main advantage of HA coatings on implants lies in accelerated and increased bone ingrowth, earlier stability, reduced migration rates and less loosening, all of which allows earlier return to normal function. This incorporation seals the bone±prosthesis interface blocking synovial penetration which causes osteolysis.

Cementless fixation techniques in joint replacement

197

9.3 Hydroxyapatite-coated total knee replacement.

9.4.5

Physiology of osseo-integration

Bone ingrowth refers to the formation of bone within an irregular surface of an implant, which improves the implant's integration into bone. The presence of a porous-coated implant evokes a cellular and physiological response that resembles the healing cascade of cancellous defects. In porous implants, the void spaces are filled with newly formed bone tissue in a stable situation.63 This process is similar to that of primary fracture healing in stable osteosynthesis, in which haematoma develops into mesenchymal tissue that is then replaced by woven bone and eventually undergoes remodelling to lamellar bone without ever passing through an intermediate stage of fibrocartilagenous tissue. Various studies have proven that this process is enhanced in an optimal mechanical environment of suitable loading (Figs 9.4(a) and (b)).64±66

9.5

Mechanical properties of the implant

One of the key challenges in the development of implant materials is the mechanical mimicry of biologically active bone. Bone is a composite material

198

Joint replacement technology

9.4 (a) and (b) Bone ingrowth into beaded HA-coated surface on a retrieval implant.

made of a collagenous fibre matrix stiffened by HA crystals, which account for 69% of bone mass. It is constantly in dynamic interaction with its physiological and mechanical environment. One of the most important terms in implant technology is stiffness. Stiffness is an extensive material property that is defined by the resistance of an elastic body to deflection or deformation by an applied force. This is measured by Young's modulus of elasticity, E, which is a ratio for the rate of change of stress (force per area) with strain (deformation) in gigapascal (GPa). Cancellous bone has an E of 0.04±1.0 GPa, pure titanium 105 GPa, titanium alloys 55±110 GPa and cobalt chromium alloys 200±250 GPa. The regenerating and remodelling process in bone is directly related to loading. As Wolff's law states: bone subjected to loading or stress will regenerate and bone not subjected to stress will atrophy. An implant that is much stiffer than bone demonstrates an unphysiological redistribution of force transmission at the interface, which is referred to as stress shielding. Locations of high load transmission cause surrounding bone to generate (hypertrophy) and locations of reduced load transmission cause surrounding bone to degenerate. Numerous studies have shown that the degree of stress shielding is directly related to the difference in stiffness of implant and bone.67±70 In hip replacement, this adverse effect results in proximal/metaphyseal bone loss and potentially compromises the long-term stability of an implant that may have shown ideal mid-term osseo-integration. This was particularly observed in proximal femoral stems made of cobalt±chromium alloys. Titanium, which has a much lower stiffness than cobalt chromium, was used as a femoral stem in cemented implants. The results were unfavourable due to early loosening, the titanium showed excessive flexibility in a stiff cement mantle. Only when titanium was applied in cementless fashion did it gain popularity in hip replacement.

Cementless fixation techniques in joint replacement

199

One trend in recent years to minimise stress shielding has been toward components with only proximal/metaphyseal fixation in longer implants. One target for this trend has been the proximal femur, where elasticity mismatch has a higher effect. The new implants either restrict coating in stems of standard length to the metayphyseal section or produce short stem implants. Both techniques have shown an improvement of bone loss due to less stress shielding.71±73 Extensive research into material technology is being undertaken to minimise the mismatch of bone±implant stiffness. Advances in metallurgy have recently yielded new titanium alloys that undergo heat treatment and result in altered microstructures, i.e. metastable , that have reduced stiffness by 50% to 55 GPa. Tantalum, which was mentioned in Section 9.4.3, not only shows low stiffness and excellent biocompatibility, but should demonstrate a structural loading pattern similar to that of bone because the empty space of 80% volumetric porosity gets filled up by bone. An additional issue closely related to material stiffness and cementless implantation is micromotion. A study conducted by Simon et al.74 focused on the influence of implant material stiffness on stress distribution and micromotion at the interface. They found that the low-stiffness implant showed more homogeneous stress distribution with fewer peak loads than the high stiffness implants, but, contrary to their hypothesis, the low-stiffness implant showed more micromotion. The elasticity of a structural implant must be such that it resists deformation by physiological loading. This is especially true for thin, flat implants such as the tibial component in total knee replacements or for long implants such as the proximal femur in total hip replacement. Under such circumstances, low stiffness would allow for repetitive deformation, excessive micromotion at the bone±implant interface, and therefore early breakdown.

9.6

Why do you still use cement?

No prosthesis or implantation technique can fully create a biological replica of the tissues lost. This chapter attempts to shed some light on the advantages and disadvantages of uncemented fixation techniques, especially in joint replacement.

9.6.1

Uncemented implants and revision surgery

Even though the vast majority of implants survive for longer than 15 years, it is imperative that doctors consider the implications for revision surgery when inserting these devices. This is especially important in young patients. A recently published randomised, controlled trial by Khaw et al.75 of cemented versus uncemented total knee replacement showed equivalent survival rates after 10 years. Wood,76 in a recently published review of the most recent data on cemented versus uncemented total knee replacement, concluded that uncemented fixation has the potential to improve longevity of knee replacement

200

Joint replacement technology

9.5 Aseptic loosening of tibial tray before revision, demonstrating minimal bone loss.

and delay the expenses and risks of revision surgery. Another study by Nilsson et al.77 comparing three different fixation techniques for the tibial component in total knee replacement concluded that uncemented fixation of HA-coated implants seemed to be the best solution, especially for younger patients, because their migration in radiostereometric analysis came to a halt after three months, whereas in cemented implants migration continued for up to two years. Regarding this last study, it is likely that this migration is due to micro-movements taking place in the thin fibrous layer between biotolerant cement and bone. Migration creates more play in the host site, leading to more motion and eventually more bone loss. We think that this may be an additional reason why we observe less bone loss at revision for aseptic loosening in an uncemented implant than we do in cemented implants (Fig. 9.5). An advantage of revising an uncemented prosthesis is a significant reduction in operating time and, in the case of knee revisions, in tourniquet time. Cemented implants often show irregular debonding of the interfaces between the implant and the cement and between the cement and the bone. Removal of cement from bone can be both arduous and time consuming, especially in the proximal femur. A possible, albeit rare, disadvantage in revision of an uncemented prosthesis involves a solid implant that has to be removed. This is seldom the case, since a stable implant does not often need to be removed even in the case of infection. On occasion, however, retrieval is required, for example when the revision component is not compatible with the stable implant. In such cases the removal of the stable implant can cause extensive bony destruction, such as the conical Wagner and threaded stems of the proximal femur.

9.6.2

Osteoporosis ± do you really need cement?

Severe osteoporosis can be a contra-indication to uncemented fixation, especially in implants that rely on a competent cancellous structure in the presence of

Cementless fixation techniques in joint replacement

201

marginal cortical rim support, such as the tibial plateau. If osteoporosis is so severe that sufficient primary support is not given, then cementing is warranted. Age itself is not a reliable predictor of the degree of osteoporosis. Our experience in cementless total knee replacement shows that elderly patients over 75 years of age do just as well as patients in the younger age group.61 An interesting study conducted by Therbo et al.78 analysed the influence of preoperative bone mineral content of the proximal tibia in uncemented TKA on the revision rate. They found that low trabecular bone quality was not a predictor for later revision surgery. Takashi et al. also demonstrated that the type of osteoblastic response to osteoarthritis according to Bombelli79 as atrophic, normotrophic and hypertrophic had no effect on the predictive outcome of cementless cup fixation in total hip replacement after seven years. These results are confirmed by numerous other studies80±84 and relate to all forms of osteoporosis without regards to the etiology. Bone grafting in large defects, as seen in rheumatoid arthritis, haemophilia and pigmented villonodular synovitis, can with expert judgement be filled with local cuttings from the surgery, with excellent success. This will induce osteogenesis and is preferable to filling these defects with large amounts of cement (Figs 9.6 and 9.7).

9.6 Tibial bone grafting in post-traumatic osteoarthritis using an uncemented prosthesis.

9.7 Filling of a giant cyst in the distal femur with autologous bone graft.

202

Joint replacement technology

Cementless fixation can thus reliably be accomplished in all but the most severe cases of osteoporosis. However, initial stability must be guaranteed to allow permanent secure fixation.

9.6.3

Infected implant in the presence of cement

Infection is a potentially disastrous complication in joint replacement. Surprisingly few publications evaluate the influence of cement on the risk for infection. It seems that the presence of an additional material and interface should at least minimally increase this risk. There is, however, one study from the Norwegian Arthroplasty Register that compared the risk of infection for cemented versus cement + gentamycin versus uncemented fixation out of 56 275 primary total hip replacements.85 Their results showed that the risk for infection was higher in the cement-only group, whereas no increased risk could be detected for both uncemented and antibiotic-loaded cemented groups. There is sufficient evidence in the literature to conclude that the presence of cement reduces resistance to infection.86±91 Once a prosthetic joint is infected, the retention potential is higher in noncemented arthroplasties. This has been demonstrated by Freeman et al.92 and correlates with our treatment of infected total knee replacements.93 A reason for this may be that the avascular cement±prosthesis interface could be a site where bacteria are protected from debridement and antibiotic penetration. In an infected prosthesis that is stable, has no periprosthetic radiolucent lines and no signs of osteomyelitis, there is a 75% chance of eradication of the infection by arthroscopic synovectomy while retaining the prosthesis.93

9.6.4

Saving cement saves time, costs and complications

Implantation of an uncemented prosthesis saves the time needed for cement preparation and polymerisation. This process takes up to 15 minutes and can take twice as long when proceeding with staged cementing. In knee replacement, cement preparation extends the tourniquet significantly, which in turn increases the possibility of deep venous thrombosis and pulmonary emboli. Hirota et al.94 used transoesophageal echocardiography to relate the number of emboli to tourniquet time. Reducing operative time puts patients, especially elderly ones, at less risk for general complications.

9.6.5

Blood loss

One disadvantage of cementless fixation is increased blood loss due to nonwatertight fit of the implant on a relatively large surface of cancellous bone. Yet this downside can be overcome. We were able to demonstrate in a prospective

Cementless fixation techniques in joint replacement

203

randomised study that this blood loss could be reduced by 30% with the use of low vacuum drains and still ensure sufficient drainage of haematoma.95

9.6.6

Costs: modern uncemented implants are cheaper

One of the biggest arguments against cementless knee prostheses rests on the implant's higher price tag. Numerous studies have recommended cemented prosthesis on the basis of lower cost. We believe this argument does not properly consider the market. As the supply of cementless implants has increased, their price has decreased. In comparing the actual prices of a number of cemented total knee replacements and their uncemented counterparts on the market today, we notice a price increase of 0±10%. The most commonly used components are often the same price. This is even more the case in total hip replacements. An increased cost of 10% for materials is more than saved by the shorter use of the operating theatre and by obviating the cost of cement. Future expenses are also cut by reduced revision rates, higher retention rates in the case of infection and prevention of treatment costs that may arise from increased general complication rates due to longer surgery.

9.7

Future trends

Joint replacement has made major strides in recent years and now provides excellent pain relief and return of function for patients suffering from destructive joint disease. Whereas only two decades ago patients were told that their new hip joints might survive 10 years, some of these same implants are now demonstrating survival rates of over 90% after 20 years. The success rate of knee joints has recorded similar growth. Despite these promising developments, however, there is a continuing increased revision surgery. Kurtz et al.96 used the National Inpatient Sample and US Census data to quantify historical trends and to make future projections regarding primary and revision joint arthroplasty. They anticipated that the number of knee replacements would double by 2015 and hip replacements would double by 2026. The number of revision knee replacements was projected to increase 500% from 2005 to 2030. The chief objective of modern research is to minimise the revision load by developing products and techniques that maximise the longevity of primary implants. Ideally, joint replacement should be a biological replication of lost tissue. Although this goal remains distant, the most promising developments in technology and design are based on the principle of biological mimicry. This, combined with therapeutic alteration of a host's reaction to an implant, are the main concepts in current research. The pursuit of new treatment options must not deter us from advancing the even more beneficial principle of prevention. New measures preventing or slowing down the progression of joint disease must be introduced in the form of patient education and modification of diet and exercise; the promotion of

204

Joint replacement technology

protective gear and rule changes in sports will reap great benefits. Additionally, supportive funding for research into conservative treatment methods of early joint disease must be prioritised. In regard to cementless fixation the two main areas of current research are the development of new surface materials and the alteration of existing surfaces to promote better bonding. Literature today in multiple fields of research is replete with the mention of nanotechnology as a scientific breakthrough promising to be the instrument of modern invention. A double-stranded DNA helix is roughly 2 nm in diameter. Nanotechnology has many promising theoretical applications in bone fixation but it also poses some potential dangers to human health. It can be seen as an extension of existing sciences and technologies into the nanoscale, or a recasting of existing sciences into a more modern term.97 The ultrastructure of bone can also be called a nanostructure of collagen fibres and hydroxyapatite. Current implementations of nanotechnology in implant development range from construction of bulky bone-like structures to composition of `osteophilic' surfaces on implants. To our knowledge, there are no commercial uses of products consisting of nanoparticles in implant technology to date. There are numerous research groups simulating the composition of biological structures or biological processes by the use of nanomaterials such as nanorods, nanotubes and so-called fullerenes. Experiments have been able to demonstrate enhancement of osteogenesis in animal studies98 on nano-sized and nano-structured titanium surfaces. There are concerns, however, that nanoparticles could have harmful effects on human health. There is growing evidence indicating a potential toxicity of nanoparticles, especially due to their higher chemical reactivity and biological activity. Nanoparticles are capable of crossing all biological barriers in cells, tissues and organs, and because of their miniscule size are much more readily taken up by the human body. This field of research remains in its earliest stages and extensive in vitro and in vivo studies are warranted before clinical testing can be undertaken. Bisphosphonates administered locally as well as systemically have demonstrated an enhancement of the osteoconductive effect in HA-coated implants.99±103 To our knowledge, there are no published results available yet of its use in clinical trials. In animal studies the local application increases biomechanical fixation strength two-fold compared with non-treated HA coated implants. Clearly, more work remains to be done. Currently the most common experiments in osteoinduction are the application of bone growth factors on or preferably in HA coatings to prevent burst release of growth factors upon implantation.104 Because of their physiological high expression during bone ingrowth, the most studied growth factors are tumour growth factor 1 (TGF 1) and insulin-like growth factor 1 (IGF1). Both proteins show an improvement of mechanical fixation and osseointegration of titanium implants without production of fibrous tissue in the interface when delivered in a biodegradable coating.105±107

Cementless fixation techniques in joint replacement

205

Another modification of existing coatings is the addition of antibiotics to HA coatings. Alt et al.108 showed significant improvement of infection prophylaxis in gentamycin HA-coated implants infected with Staphylococcus aureus in rabbits. Chemical treatment of titanium mesh surfaces without HA coatings by a hydrogen peroxide solution containing tantalum chloride has also been shown to enhance bone ingrowth.109 Other research groups are carrying out extensive investigations into the use of such biopolymers as chitosan, a biologically produced polymer closely related to chitin, which is the second most abundant form of polymerised carbon in nature.110,111 Seung Yun Shin et al.112 were able to confirm the biocompatibility of chitosan nanofibre membranes in rats with evidence of enhanced bone regeneration and no evidence of an inflammatory reaction. Tissue engineering is a biological approach attempting to obtain osteoinductive coatings by in vitro cultivation of patient cells to form bone tissue. Bruijn et al. conducted experiments by coating implants with cultured osteogenic bone marrow cells from rats, goats and humans. Calcium phosphatecoated implants were seeded with bone marrow cells cultured for a week to facilitate osteogenic proliferation and extracellular modification. These hybrid implants were then subcutaneously implanted in nude mice and retrieved after four weeks. Histological examination revealed newly formed bone in both porous implants and on flat metallic surfaces.104 Gene therapy is another vast field of research in medicine that shows promising applications in orthopaedics. The principle of providing genetic information via a viral or non-viral delivery vehicle into a living cell is one of the most elegant ways to improve host site physiology. In cementless fixation this information could contain the codes for producing growth factors in continuous fashion by periprosthetic cells. Gene therapy could also provide breakthrough in regeneration of cartilage, which remains the crux of joint disease therapy to this day. In summary cementless fixation: · · · · · · ·

works; saves bone; reduces operative exposure and infection rates; decreases the number of interfaces; saves time; simplifies revision; with hydroxyapatite, seals the bone interface and prevents osteolysis.

9.8

References

1 Mohan, A.R. and M. Gross, Cementless total knee replacement ± a prospective 12 to 15 year follow-up study. J Bone Joint Surg Br, 2004. 86-B(Supp III): p. 318.

206

Joint replacement technology

2 Engh, C.A. Sr, Outcome of total hip arthroplasty using extensively porous-coated components at 20-year follow-up. In Annual meeting, AAOS 2006. 2006. 3 Charnley, J., Anchorage of the femoral head prosthesis to the shaft of the femur. J Bone Joint Surg Br, 1960. 42-B: pp. 28±30. 4 Charnley, J., Arthroplasty of the hip. A new operation. Lancet, 1961. 1(7187): pp. 1129±32. 5 Charnley, J., [Adherence of Prostheses to Living Bone.]. Acta Orthop Belg, 1964. 30: pp. 663±72. 6 Gunston, F.H., Polycentric knee arthroplasty. Prosthetic simulation of normal knee movement. J Bone Joint Surg Br, 1971. 53(2): pp. 272±7. 7 Mirra, J.M., et al., The pathology of the joint tissues and its clinical relevance in prosthesis failure. Clin Orthop Relat Res, 1976(117): pp. 221±40. 8 Schmalzreid, T.P. and J.J. Callaghan, Current concepts review ± wear in total hip and knee replacements. J Bone Joint Surg Am, 1999. 81(1): pp. 115±36. 9 Horowitz, S.M., et al., Studies of the mechanism by which the mechanical failure of polymethylmethacrylate leads to bone resorption. J Bone Joint Surg Am, 1993. 75(6): pp. 802±13. 10 Jones, L.C. and D.S. Hungerford, Cement disease. Clin Orthop Relat Res, 1987(225): pp. 192±206. 11 Jasty, M.J., et al., Localized osteolysis in stable, non-septic total hip replacement. J Bone Joint Surg Am, 1986. 68(6): p. 912±19. 12 Bourne, R.B., Rorabeck, G.H., Inman, K.J., Tozakoglou, E. Cementless total hip replacement: anatomic, cylindrical or tapered. In 67th Annual Meeting of the American Academy of Orthopaedic Surgeons. 2000. Orlando, FL. 13 Miura, H., et al., Effects of screws and a sleeve on initial fixation in uncemented total knee tibial components. Clin Orthop Relat Res, 1990(259): pp. 160±8. 14 Hayakawa, T., et al., Trabecular bone response to surface roughened and calcium phosphate (Ca-P) coated titanium implants. Biomaterials, 2002. 23(4): pp. 1025±31. 15 Shalabi, M.M., et al., Implant surface roughness and bone healing: a systematic review. J Dent Res, 2006. 85(6): pp. 496±500. 16 Shalabi, M.M., J.G. Wolke, and J.A. Jansen, The effects of implant surface roughness and surgical technique on implant fixation in an in vitro model. Clin Oral Implants Res, 2006. 17(2): pp. 172±8. 17 Kold, S., et al., Bone compaction enhances fixation of weightbearing titanium implants. Clin Orthop Relat Res, 2005(431): pp. 138±44. 18 Kold, S., et al., Bone compaction enhances fixation of weight-bearing hydroxyapatite-coated implants. J Arthroplasty, 2006. 21(2): pp. 263±70. 19 Kold, S., et al., Bone compaction enhances fixation of hydroxyapatite-coated implants in a canine gap model. J Biomed Mater Res B Appl Biomater, 2005. 75(1): pp. 49±55. 20 Dalton, J.E., et al., The effect of operative fit and hydroxyapatite coating on the mechanical and biological response to porous implants. J Bone Joint Surg Am, 1995. 77(1): pp. 97±110. 21 Green, J.R., et al., The effect of bone compaction on early fixation of porous-coated implants. J Arthroplasty, 1999. 14(1): pp. 91±7. 22 Channer, M.A., et al., Use of bone compaction in total knee arthroplasty. J Arthroplasty, 1996. 11(6): pp. 743±9. 23 Chareancholvanich, K., et al., In vitro stability of cemented and cementless femoral stems with compaction. Clin Orthop Relat Res, 2002(394): pp. 290±302. 24 W. Norman Scott, Insall & Scott's Surgery of the Knee e-dition. 4th edition 2006,

Cementless fixation techniques in joint replacement

207

vol. I: Churchill Livingstone. 25 Au S. Wong, G. Isaac, A.M.R. New and M. Taylor. Influence of bone quality on the initial stability of cementless hip stem in total hip arthroplasty. In 2003 Summer Bioengineering Conference. 2003. Sonesta Beach Resort in Key Biscayne, Florida. 26 Ryd, L., et al., Roentgen stereophotogrammetric analysis as a predictor of mechanical loosening of knee prostheses. J Bone Joint Surg Br, 1995. 77(3): pp. 377±83. 27 Karrholm, J., et al., Radiostereometry of hip prostheses. Review of methodology and clinical results. Clin Orthop Relat Res, 1997(344): pp. 94±110. 28 Burke, D.W. et al., Dynamic measurement of interface mechanics in vivo and the effect of micromotion on bone ingrowth into a porous surface device under controlled loads in vivo. Trans Orthop Res Soc, 1991 16, 103. 29 Pilliar, R.M., J.M. Lee, and C. Maniatopoulos, Observations on the effect of movement on bone ingrowth into porous-surfaced implants. Clin Orthop Relat Res, 1986(208): pp. 108±13. 30 Sumner, D.R., et al., Bone ingrowth and wear debris in well-fixed cementless porous-coated tibial components removed from patients. J Arthroplasty, 1995. 10(2): pp. 157±67. 31 Sandborn, P.M., et al., Tissue response to porous-coated implants lacking initial bone apposition. J Arthroplasty, 1988. 3(4): pp. 337±46. 32 Cross, M.J., et al., Enhanced fixation of uncemented knee replacement with hydroxyapatite. Key Engineering Materials, 2003, 240±242: pp. 773±776. 33 Goldberg, V.M., et al., Biology of grit-blasted titanium alloy implants. Clin Orthop Relat Res, 1995(319): pp. 122±9. 34 Friedman, R.J., et al., Influence of biomaterial surface texture on bone ingrowth in the rabbit femur. J Orthop Res, 1996. 14(3): pp. 455±64. 35 Bobyn, J.D., et al., The optimum pore size for the fixation of porous-surfaced metal implants by the ingrowth of bone. Clin Orthop Relat Res, 1980(150): pp. 263±70. 36 Cook, S.D., K.A. Walsh, and R.J. Haddad, Jr, Interface mechanics and bone growth into porous Co-Cr-Mo alloy implants. Clin Orthop Relat Res, 1985(193): pp. 271± 80. 37 Clemow, A.J., et al., Interface mechanics of porous titanium implants. J Biomed Mater Res, 1981. 15(1): pp. 73±82. 38 Jasty, M., et al., Bone ingrowth into porous coated canine total hip replacements. Quantification by backscattered scanning electron microscopy and image analysis. Scanning Microsc, 1989. 3(4): pp. 1051±6; discussion 1056±7. 39 Dammak, M., et al., Friction properties at the bone±metal interface: comparison of four different porous metal surfaces. J Biomed Mater Res, 1997. 35(3): pp. 329±36. 40 Bellemans, J., Osseointegration in porous coated knee arthroplasty. The influence of component coating type in sheep. Acta Orthop Scand Suppl, 1999. 288: pp. 1±35. 41 Osborn, J.F., [Biomaterials and their application to implantation]. SSO Schweiz Monatsschr Zahnheilkd, 1979. 89(11): pp. 1138±9. 42 Wilke, A., et al., Biocompatibility analysis of different biomaterials in human bone marrow cell cultures. J Biomed Mater Res, 1998. 40(2): pp. 301±6. 43 Levine, B., C.J. Della Valle, and J.J. Jacobs, Applications of porous tantalum in total hip arthroplasty. J Am Acad Orthop Surg, 2006. 14(12): pp. 646±55. 44 Shuler, M.S., M.D. Rooks, and J.R. Roberson, Porous tantalum implant in early osteonecrosis of the hip: preliminary report on operative, survival, and outcomes results. J Arthroplasty, 2007. 22(1): pp. 26±31. 45 Itala, A., J. Reach, F.H. Sim, K.N. An and D.G. Lewallen, Patellar tendon

208

46 47 48 49 50 51 52 53 54 55 56 57 58 59 60 61 62 63 64

Joint replacement technology attachment and healing to porous tantalum: an experimental canine study. J Arthroplasty, 2006. 21(2): p. 303. Nasser, S. and R.A. Poggie, Revision and salvage patellar arthroplasty using a porous tantalum implant. J Arthroplasty, 2004. 19(5): pp. 562±72. Nehme, A., D.G. Lewallen, and A.D. Hanssen, Modular porous metal augments for treatment of severe acetabular bone loss during revision hip arthroplasty. Clin Orthop Relat Res, 2004(429): pp. 201±8. Macheras, G.A., et al., Radiological evaluation of the metal±bone interface of a porous tantalum monoblock acetabular component. J Bone Joint Surg Br, 2006. 88(3): pp. 304±9. Bobyn, J.D., et al., Characteristics of bone ingrowth and interface mechanics of a new porous tantalum biomaterial. J Bone Joint Surg Br, 1999. 81(5): pp. 907±14. Cooley, D.R., et al., The advantages of coated titanium implants prepared by radiofrequency sputtering from hydroxyapatite. J Prosthet Dent, 1992. 67(1): pp. 93±100. Soballe, K., et al., Gap healing enhanced by hydroxyapatite coating in dogs. Clin Orthop Relat Res, 1991(272): pp. 300±7. Mouzin, O., K. Soballe, and J.E. Bechtold, Loading improves anchorage of hydroxyapatite implants more than titanium implants. J Biomed Mater Res, 2001. 58(1): pp. 61±8. Rahbek, O., et al., Superior sealing effect of hydroxyapatite in porous-coated implants: experimental studies on the migration of polyethylene particles around stable and unstable implants in dogs. Acta Orthop, 2005. 76(3): pp. 375±85. Epinette, J.A. and M.T. Manley, Hydroxyapatite-coated total knee replacement: clinical experience at 10 to 15 years. J Bone Joint Surg Br, 2007. 89-B(1): pp. 34± 38. Munro, J., et al., Total hip replacement hip taper-design stem and hydroxyapatite coating. Clinical outcome and quantitative CT-assisted osteodensitometry. J Bone Joint Surg Br, 2005. 87-B(Supp I): p. 29. Shetty, A.A., et al., Results of a hydroxyapatite-coated (Furlong) total hip replacement: a 13- to 15-year follow-up. J Bone Joint Surg Br, 2005. 87-B(8): pp. 1050±4. Landor, I., et al., Hydroxyapatite porous coating and the osteointegration of the total hip replacement. Arch Orthop Trauma Surg, 2007. 127(2): pp. 81±9. Geesink, R.G. and N.H. Hoefnagels, Six-year results of hydroxyapatite-coated total hip replacement. J Bone Joint Surg Br, 1995. 77(4): pp. 534±47. Rasquinha, V.J., C.S. Ranawat, and A.J. Mauriello, Jr, Hydroxyapatite: catalyst or conjuror? J Arthroplasty, 2002. 17(4 Suppl 1): pp. 113±17. Cross, M.J. and E.N. Parish, A hydroxyapatite-coated total knee replacement: prospective analysis of 1000 patients. J Bone Joint Surg Br, 2005. 87(8): pp. 1073±6. Dixon, P., et al., Hydroxyapatite-coated, cementless total knee replacement in patients aged 75 years and over. J Bone Joint Surg Br, 2004. 86(2): pp. 200±4. Tai, C.C. and M.J. Cross, Five- to 12-year follow-up of a hydroxyapatite-coated, cementless total knee replacement in young, active patients. J Bone Joint Surg Br, 2006. 88(9): pp. 1158±63. Kienapfel, H., et al., Implant fixation by bone ingrowth. J Arthroplasty, 1999. 14(3): pp. 355±68. Frost, H.M., A brief review for orthopedic surgeons: fatigue damage (microdamage) in bone (its determinants and clinical implications). J Orthop Sci, 1998. 3(5): pp. 272±81.

Cementless fixation techniques in joint replacement

209

65 Huiskes, R., et al., Effects of mechanical forces on maintenance and adaptation of form in trabecular bone. Nature, 2000. 405(6787): pp. 704±6. 66 Chicurel, M.E., et al., Integrin binding and mechanical tension induce movement of mRNA and ribosomes to focal adhesions. Nature, 1998. 392(6677): pp. 730±3. 67 Sychterz, C.J., et al., Effect of femoral stiffness on bone remodeling after uncemented arthroplasty. Clin Orthop Relat Res, 2001(389): pp. 218±27. 68 P. Christel, A.Meunier, and A.J.C. Lee, Biological and biomechanical performance of biomaterials. In Advances in Biomaterials, vol. 6, 1985, Fifth European Conference on Biomaterials, Paris: Elsevier. 69 Huiskes, R. and E.Y. Chao, A survey of finite element analysis in orthopedic biomechanics: the first decade. J Biomech, 1983. 16(6): pp. 385±409. 70 Long, M. and H.J. Rack, Titanium alloys in total joint replacement ± a materials science perspective. Biomaterials, 1998. 19(18): pp. 1621±39. 71 Glassman, A.H., J.D. Bobyn, and M. Tanzer, New femoral designs: do they influence stress shielding? Clin Orthop Relat Res, 2006. 453: pp. 64±74. 72 Francesco P., L. Malfetta, and M. Grandizio, Preservation of the femoral neck in hip arthroplasty: results of a 13- to 17-year follow-up. J Orthop Traumatology, 2000. 1(1): pp. 31±9. 73 Sumner, D.R. and J.O. Galante, Determinants of stress shielding: design versus materials versus interface. Clin Orthop Relat Res, 1992(274): pp. 202±12. 74 Simon, U., et al., Influence of the stiffness of bone defect implants on the mechanical conditions at the interface ± a finite element analysis with contact. J Biomech, 2003. 36(8): pp. 1079±86. 75 Khaw, F.M., et al., A randomised, controlled trial of cemented versus cementless press-fit condylar total knee replacement. Ten-year survival analysis. J Bone Joint Surg Br, 2002. 84(5): pp. 658±66. 76 Wood, J.E.J., Uncemented fixation in total knee replacement: a promising future. Curr Opin Orthop, 18(1): pp. 61±5. 77 Nilsson, K.G., et al., Uncemented HA-coated implant is the optimum fixation for TKA in the young patient. Clin Orthop Relat Res, 2006. 448: pp. 129±39. 78 Therbo, M., et al., Influence of pre-operative bone mineral content of the proximal tibia on revision rate after uncemented knee arthroplasty. J Bone Joint Surg Br, 2003. 85(7): pp. 975±9. 79 Bombelli, R., Osteoarthritis of the Hip: Classification and Pathogenesis: The Role of Osteotomy as a Consequent Therapy. 2nd edn. 1983. Berlin: Springer-Verlag. 80 Panagiotis, Z., et al., Biological fixation of cementless stems in patients over 75 years with diagnosed osteoporosis. J Bone Joint Surg Br, 2004. 86-B(Supp III): p. 320. 81 Healy, W.L., Hip implant selection for total hip arthroplasty in elderly patients. Clin Orthop Relat Res, 2002(405): pp. 54±64. 82 Ekelund, A., N. Rydell, and O.S. Nilsson, Total hip arthroplasty in patients 80 years of age and older. Clin Orthop Relat Res, 1992(281): pp. 101±6. 83 Effenberger, H., et al., Successful hip arthroplasty using cementless titanium implants in rheumatoid arthritis. Arch Orthop Trauma Surg, 2002. 122(2): pp. 80±7. 84 Jana, A.K., et al., Total hip arthroplasty using porous-coated femoral components in patients with rheumatoid arthritis. J Bone Joint Surg Br, 2001. 83-B(5): pp. 686±90. 85 Engesaeter, L.B., et al., Does cement increase the risk of infection in primary total hip arthroplasty? Revision rates in 56,275 cemented and uncemented primary THAs followed for 0±16 years in the Norwegian Arthroplasty Register. Acta Orthop, 2006. 77(3): pp. 351±8.

210

Joint replacement technology

86 Petty, W., The effect of methylmethacrylate on chemotaxis of polymorphonuclear leukocytes. J Bone Joint Surg Am, 1978. 60(4): pp. 492±8. 87 Petty, W., The effect of methylmethacrylate on bacterial phagocytosis and killing by human polymorphonuclear leukocytes. J Bone Joint Surg Am, 1978. 60(6): pp. 752±7. 88 Petty, W., The effect of methylmethacrylate on the bacterial inhibiting properties of normal human serum. Clin Orthop Relat Res, 1978(132): pp. 266±78. 89 Petty, R., Influence of skeletal implant material on infection. Trans ORS, 1983. 8: pp. 137. 90 Samuelson, K.M., D. Au, G.L. Rasmussen, and R.L. Ellingson, Evaluation of cemented versus cementless fixation and infections in caninetotal joint arthroplasties. Trans ORS, 1983. 8: p. 230. 91 Gristina, A.G. and J. Kolkin, Current concepts review. Total joint replacement and sepsis. J Bone Joint Surg Am, 1983. 65(1): pp. 128±34. 92 Freeman, M.A., et al., The management of infected total knee replacements. J Bone Joint Surg Br, 1985. 67(5): pp. 764±8. 93 Dixon, P., E.N. Parish, and M.J. Cross, Arthroscopic debridement in the treatment of the infected total knee replacement. J Bone Joint Surg Br, 2004. 86(1): pp. 39±42. 94 Hirota, K., et al., The relationship between pneumatic tourniquet time and the amount of pulmonary emboli in patients undergoing knee arthroscopic surgeries. Anesth Analg, 2001. 93(3): pp. 776±80. 95 Morgan-Jones, R.L., M.M. Perko, and M.J. Cross, Uncemented total knee replacement ± the favourable influence of low over high pressure drainage. Knee, 2000. 7(3): pp. 149±50. 96 Kurtz, S., et al., Projections of primary and revision hip and knee arthroplasty in the United States from 2005 to 2030. J Bone Joint Surg Am, 2007. 89(4): pp. 780±5. 97 http://en.wikipedia.org/wiki/Nanotechnology. 98 Meirelles, L., et al., Increased bone formation to unstable nano rough titanium implants. Clin Oral Implants Res, 2007. 18: pp. 326±32. 99 Stadelmann, V. and D.P. Pioletti, Using Orthopedic Implants for Local Bisphosphonate Delivery in Osteoporotic Animals: A Theoretical Model, Biomechanical Orthopedics Laboratory, Swiss Federal Institute of Technology: Lausanne, Switzerland. 100 Jakobsen, T., et al., Local alendronate increases fixation of implants inserted with bone compaction: 12-week canine study. J Orthop Res, 2007. 25(4): pp. 432±41. 101 Jensen, T.B., et al., Systemic alendronate treatment improves fixation of press-fit implants: a canine study using nonloaded implants. J Orthop Res, 2007. 25(6): pp. 772±8. 102 Eberhardt, C., et al., The bisphosphonate ibandronate accelerates osseointegration of hydroxyapatite-coated cementless implants in an animal model. J Orthop Sci, 2007. 12(1): pp. 61±6. 103 Eberhardt, C., et al., Osseointegration of cementless implants with different bisphosphonate regimens. Clin Orthop Relat Res, 2006. 447: pp. 195±200. 104 de Bruijn, J.D., et al., Bone induction by implants coated with cultured osteogenic bone marrow cells. Adv Dent Res, 1999. 13: pp. 74±81. 105 Lamberg, A., et al., Locally delivered TGF-beta1 and IGF-1 enhance the fixation of titanium implants: a study in dogs. Acta Orthop, 2006. 77(5): pp. 799±805. 106 Aspenberg, P., T. Albrektsson, and K.G. Thorngren, Local application of growthfactor IGF-1 to healing bone. Experiments with a titanium chamber in rabbits. Acta Orthop Scand, 1989. 60(5): pp. 607±10.

Cementless fixation techniques in joint replacement

211

107 Lind, M., et al., Transforming growth factor-beta stimulates bone ongrowth. Hydroxyapatite-coated implants studied in dogs. Acta Orthop Scand, 1996. 67(6): pp. 611±16. 108 Alt, V., et al., The effects of combined gentamicin-hydroxyapatite coating for cementless joint prostheses on the reduction of infection rates in a rabbit infection prophylaxis model. Biomaterials, 2006. 27(26): pp. 4627±34. 109 Taeseong Kim, M.S., C. Ohtsuki, K. Masuda, H. Tamai, E. Watanabe, A. Osaka, and H. Moriya, Enhancement of bone growth in titanium fiber mesh by surface modification with hydrogen peroxide solution containing tantalum chloride. J Biomed Mater Res Part B: Appl Biomater, 2003. 64B(Issue 1): pp. pp. 19±26. 110 Bumgardner, J.D., et al., The integration of chitosan-coated titanium in bone: an in vivo study in rabbits. Implant Dent, 2007. 16(1): pp. 66±79. 111 Jarry, C., et al., Effects of steam sterilization on thermogelling chitosan-based gels. J Biomed Mater Res, 2001. 58(1): pp. 127±35. 112 Shin, S.Y., et al., Biological evaluation of chitosan nanofiber membrane for guided bone regeneration. J Periodontol, 2005. 76(10): pp. 1778±84.