June 2004 : No 3

Surgery for Arthritis: Total Hip and Knee Joint Replacement

John Fisher

Institute of Medical and Biological Engineering, University of Leeds

Reports on the Rheumatic Diseases Series 5 : Topical Reviews

  • Artificial joints have been implanted for the treatment of arthritic hips and knees for approximately 40 years
  • These remain one of the most successful forms of prosthetic surgery
  • For elderly patients the success rates are well over 90% at 10 years
  • Historically most prostheses have been manufactured with metal-on-polyethylene bearing surfaces but recently it has been discovered that adverse tissue reactions to polyethylene wear particles, accumulated in periprosthetic tissues, can cause bone resorption and loosening in the long term
  • New bearing materials and joint designs have been developed for younger, more active patients
  • In both the hip and knee, bone preserving conservative designs are being developed for earlier intervention

Introduction

Surgery and a total replacement joint is the final stage of treatment for many patients with arthritis after other approaches such as drug therapy and physical therapy have been exhausted.1,2 Joint replacement in its present form has been undertaken clinically for over 40 years in the hip3 and over 30 years in the knee,4 and remains one of the most successful applications of prosthetic surgery today. Currently over 50,000 total hip joint replacements and over 30,000 knee joint replacements are undertaken in the UK every year. The number of knee joint replacements is increasing at a greater rate than hip joint replacements, and in some countries, such as the USA, similar numbers of knee prostheses and hip prostheses are now implanted. In the early days of joint replacement3 the procedure was predominantly reserved for the more elderly patients. The success of the procedures, the increased longevity of the population, the demand for increased quality of life and more active lifestyles, the earlier onset and diagnosis of arthritis, and the correction of congenital abnormalities mean that joint replacement is now undertaken in a wider age-range of patients.5 This has placed increased demands on both the design and the performance of the prostheses, as well as the surgical technique itself. The outcome of joint replacement in the hip and knee in the elderly is excellent, with over 90% survival rates at 10 years.4,6 In younger patients, with longer-term follow-up, survival rates are lower and the incidence of and need for a second revision operation increase.7 In the UK over 10% of all joint replacement operations are for revision of failed prostheses. It is for the younger patients below 60 years of age with life expectancies beyond 20 years that new materials, designs and surgical approaches are being developed8 to further extend the lifetime of the prostheses. Additionally, it is in this younger group that alternative approaches are now being introduced. These include surgical repair and cell and tissue transplantation.1,9 In the hip, surface replacement10 is being introduced, and in the knee, arthroscopy osteotomy,11 interpositional spacers12 and unicompartmental knee arthroplasty.13

Requirements for joint replacement

The primary function of joint replacement surgery is to relieve pain and restore function, which includes transmitting physiological loads and the provision of both a physiological range of movement14 and an articulation with minimum friction and wear.15 The prosthetic materials must be anatomical in shape, be securely fixed to the surrounding bone, and remain securely fixed throughout the patient's lifespan. Preferably, the prosthetic components should be implanted with minimum bone resection. Additionally, the biomaterials from which the implant is made have to be biocompatible and any wear particles produced must also be compatible with the body and not cause adverse biological reactions.16 The joint replacements have to be compatible with a range of different patient anatomies and geometries and typically a range of different sizes is necessary. Similarly, the bone quality of patients is quite variable and the methods of fixation have to be able to accommodate different bone interface conditions. The body environment into which the joint replacement is implanted is extremely challenging. Not only does it have particular chemical, biochemical, biological and biomechanical characteristics, but also the fact that the tissue surrounding the prosthetic components remains living means that the joint replacement interface and environment can continually change with time. These changes not only are related to the natural ageing of the patient but also can occur in response to the function and properties of the prosthetic device itself. This results in a complex interactive biological and biomechanical environment involving the living tissue and prosthetic joint in the body which can determine the lifetime of the joint replacement. Over the years it has proved very difficult to predict preclinically many of these interactions, and it has only been as a result of clinical experience and research that particular clinical failure and success scenarios have emerged. This has resulted in more rigorous and demanding requirements for designs of joint replacement.

Evolution of joint replacement

Hip

Hip joint replacement as we know it today started in the early 1960s through two pioneering surgeons: John Charnley, who introduced metal-on-polyethylene hip prostheses along with bone cement for fixation,3 and George McKee, who introduced metal-on-metal hip prostheses.17 Difficulties in the manufacturing of the metal-on-metal prostheses and perceived poor fixation resulted in the discontinuation of this approach in the 1970s, only for it to re-emerge during the 1990s.18,19

The approach adopted by Charnley of bone cement fixation with a metal stem, and a polished femoral head articulating on an ultra-high molecular weight polyethylene (UHMWPE) acetabular cup3 (Figure 1), became standard for the hip6,7,20,21 and was subsequently adopted for the knee4 during the 1960s and 1970s. The early principles of Charnley, using a small-diameter femoral head which reduced frictional torque (the product of the friction coefficient and radius of the head), also reduced the sliding distance per step and hence the wear. In other devices and designs different sizes of femoral head (28 mm and 32 mm) were introduced.22 Charnley also initially implanted only in older, low-demand patients and this resulted in good initial clinical results in the first 10 years.2,20,21 More widespread use with a greater number of patients and larger follow-up began to highlight limitations. Early fracture of the femoral stem due to fatigue23 was resolved through improved metal alloys and geometrical modifications to the stem design.

FIGURE 1. An artificial hip joint with metallic femoral stem and ultra-high molecular weight polyethylene acetabular cup.

By the 1970s, wider concerns were emerging regarding localised bone resorption surrounding hip prostheses.24 These related to the long-term reaction to self-curing acrylic bone cement25,26 and, in Europe, the role of wear particles27 – although the latter received little attention for a further 10 years. Concerns about the bone cement26 led to the rapid development of cementless fixation systems, first as simple press-fit or mechanical interlocks, then through the development of porous coatings to allow bone ingrowth,28 and more recently through development in Europe of bioactive coatings such as hydroxyapatite, which can also be used with porous coatings.29 Currently approximately half the implants used worldwide do not use bone cement, and there is a trend towards using cementless fixation in younger patients with good bone quality. The introduction of cementless prostheses involved the use of larger-sized and stiffer implants, which resulted in the emergence during the 1980s of the problem of stress shielding in the proximal femur and resulting bone resorption.30 Recognition of this problem resulted in the evolution during the 1990s of stem designs with proximal fit and proximal coatings, which transfer stress more effectively to the proximal femur.31 In cemented prostheses the development of smooth taper stems also encouraged stress transfer through the cement to the proximal femur.32

Despite the introduction of cementless prostheses during the 1980s, the bone resorption and adverse tissue reactions initially reported in the late 1970s27 and early 1980s33 continued to emerge. At first bone cement and metal particles or even fluid pressure were considered to contribute to osteolysis.32,34,35 In the early 1990s, however, it became clear that polyethylene wear debris was the major cause of osteolysis and loosening in joint replacement.16,36-40 Studies of retrieved tissues showed an abundance of micron- and submicron-sized polyethylene wear particles41-45 and these were also found in laboratory wear studies.45,46 These particles were shown to stimulate macrophages to release osteolytic cytokines, which lead to osteolysis and bone resorption.16,32,47-50 More recent studies have shown that the submicron polyethylene wear particles (0.1–1 μm) are more reactive than the larger particles (>1 μm)51,52 and that the release of osteolytic cytokines is dependent on the volumetric concentration of the submicron wear particles.53 This has led to the study of the volume distribution of the wear particles in different size ranges to analyse their osteolytic potential.54,55 By the start of the 1990s nearly all hip prostheses implanted worldwide had polyethylene acetabular cups. This prompted considerable research into factors that caused acceleration of the wear of polyethylene as well as the development of alternative bearing surfaces and new technologies to reduce wear and osteolysis.

During the first three decades of hip joint replacement the majority of polyethylene components were sterilised using gamma irradiation in the presence of air. During the early 1990s it emerged that the irradiation, which causes chain scission and free radicals,56 renders the material unstable – subject to oxidative degradation causing a reduction in its mechanical properties57 and an increase in the wear rate.58 As well as a higher wear rate the oxidised materials also produce smaller particles with greater osteolytic potential.59 The majority of femoral heads were constructed from polished metal alloys which were shown to become scratched and damaged and cause accelerated wear.60 The widespread recognition of the role of polyethylene wear debris induced osteolysis in the long-term failure of hip prostheses has led to a new generation of designs and bearing materials for hip prostheses.

Knee

In many ways the evolution of knee joint replacement has followed on from developments in hip joint replacement. Early knee joint replacements were in the form of simple metal hinges. However, these were not successful as they did not allow for the 6° of motion and forces found in the knee.61-63 The total condylar knee, which utilises a polished metal femoral component and a UHMWPE tibial bearing component,4 was developed in the late 1960s from the metal-on-polyethylene experience in the hip. This formed the basis for many of the total knee joint replacements currently available (Figure 2). Early total knee designs were quite conforming, placing constraints on other motions such as internal-external rotation, abduction-adduction and anterior-posterior translation.64 These constrained devices can result in transmission of higher forces to the fixation interfaces and cause loosening. Less constrained devices which retained both the posterior65 and anterior cruciate ligaments were developed during the 1970s and 1980s66 and have been reviewed.67 The complex relationships between geometry, constraints of ligament function forces, and resulting motions continue to be studied.68 As the knee becomes less conforming the stress levels in the polyethylene tibial components rise,69,70 which has subsequently been shown to effect longevity. The advantages/disadvantages of ligament-retaining, ligament-sacrificing or ligament-substituting designs remain a matter of debate today. There is considerable support for retaining the posterior cruciate ligament when it is functional. However, in knees with unstable soft tissues and in revision knees a more conforming ligament-sacrificing design is often used.

FIGURE 2. An artificial knee joint in situ, with metallic femoral component and ultra-high molecular weight polyethylene bearing.

Whereas early knees were implanted using bone cement the development of cementless fixation and porous coating in the hip led to the introduction of this procedure in the knee, with the use of metal tibial trays to support the polyethylene insert.71 More recently hydroxyapatite coatings have also been applied in the knee. The concept of mobile bearing knees was introduced in the late 1970s.66 This technique utilised a metal tibial tray on which the polyethylene insert could slide, providing freedom of motion to translate in the anterior-posterior direction and to rotate. This permitted greater conformity of the femoral-tibial articulation, hence reducing contact stress in the polyethylene.66 The short- to medium-term success of knee joint replacements during the 1970s and 1980s was considered good.4,66,67,72

As with the hip, longer-term follow-up led to the identification of wear of polyethylene as a cause of long-term failure in the knee joint replacement.73-78 Polyethylene wear related failures presented themselves in different forms in the knee. Gross failure and delamination surface wear, polishing wear as found in the hip, and backside wear79 have been identified. As with hip prostheses, the majority of polyethylenes implanted before 1995 were sterilised with gamma irradiation in the presence of air, and oxidative degradation,57 leading to reduction in mechanical properties and increased wear, was identified as a cause of premature failure.79-88

The introduction of stabilised and oxidation-resistant polyethylenes in the late 1990s88 has led to a substantial reduction in the delamination potential of knee prostheses.89 However, the generation of wear particles remains a problem,90,91 although the stabilised and oxidation-resistant polyethylenes wear less than the historical gamma-in-air polyethylene for reasons described in the hip. Polyethylene wear rates found in knee joint simulators are up to a factor of two less than those found in the hip, but are dependent on design and kinematic conditions. Analysis of debris from retrieved knee prostheses has shown the concentration in the tissues to be similar to those found in the hip92 after similar lifetimes. However, the knee debris is slightly larger and not as reactive. Nevertheless it is predicted that with longer clinical lifetimes and more active patients, polyethylene wear debris will cause higher rates of failure in the knee. At present bearing-associated problems are cited as causing up to 30% of all failures.93


Over the last decade the recognition that wear and wear debris induced osteolysis is a major factor limiting the lifetimes of joint replacement94,95 has led to the development and introduction of a range of new bearing materials and designs into joint replacements.

Current designs and bearing solutions

Hip

For many patients over the age of 65 with a life expectancy of less than 20 years a conventional metal-on-polyethylene bearing combination6 which uses stabilised and oxidation resistant UHMWPE88,96 is an adequate solution for most patients. This would commonly be fixed to the bone using polymethyl methacrylate (PMMA) bone cement.5

For younger patients under 65 with greater than 20 years' expected lifetime, or patients with particularly active lifestyles, there are alternative bearing options which may have a better long-term success rate and less propensity to osteolysis in the longer term.8 These include the use of ceramic femoral heads, intentionally cross-linked polyethylene acetabular cups, ceramic-on-ceramic bearings, and metal-on-metal bearings.8,16,94,97,98

It should be noted that none of these combinations have clinical histories well beyond 20 years and the proposed improved wear performance and reduction in osteolytic potential has still to be demonstrated in clinical practice. Some of the questions that are appropriate for the surgeon to ask in selecting a bearing are given by Greenwald and Heim.99 It should be recognised that although bearing surface performance has traditionally been compared by measuring volumetric wear, the osteolytic potential of the bearing is dependent on the nature of the wear debris. Size, morphology, chemistry, its distribution function and volumetric concentration16,51-53,55 all have to be considered in analysis and presentation of the performance of different bearing materials.

The use of alumina ceramic femoral heads, which are smooth, hard and scratch-resistant, can reduce the wear rate of stabilised polyethylene as compared with the wear produced by metal femoral heads, which can become scratched in vivo and accelerate wear.60,100,101 Clinical studies have shown a reduction in wear from little difference to four times less wear with ceramic heads on conventional polyethylenes.101 The reduction in wear depends on the amount of damage generated to metal heads in the body, which is design- and surgeon-dependent. The use of alumina ceramic femoral heads should certainly extend the wear life by up to a factor of two.

More recently highly cross-linked polyethylene has been developed to further improve the wear resistance of UHMWPE acetabular cups.102,103 The reduction in wear produced by cross-linking is also associated with the generation of smaller and more reactive particles.96,104 Longer-term clinical results are awaited for highly cross-linked polyethylene acetabular cups105 and caution should be taken in using this material with large-sized femoral heads which themselves accelerate wear.

Alumina ceramic-on-ceramic bearings have been used in limited numbers for a considerable period.106-108 Improvements in materials over time have produced more fracture-resistant materials and designs, which are associated with extremely low wear rates.109-112 The wear debris has been shown to consist of nanometre- and micron-sized particles,113,114 and cell culture studies show it to be less reactive and have less osteolytic potential than polyethylene wear debris.115 This combined with the extremely low volumetric wear has indicated its use in younger active patients. More recently a further improved alumina ceramic composite bearing has been developed which provides further increase in fracture toughness and wear resistance and is the subject of clinical trials.116

Although metal-on-metal bearings were introduced in the 1960s,16,117 it is in the last 10 years – since the development of second generation devices – that they have been used more widely.118 As with ceramic-on-ceramic, the wear rate of cobalt-chromium metal-on-metal bearings is low, ten-fold less than cross-linked polyethylene.97,119 It is important that combinations of low carbon content alloys are avoided as they produce higher wear rates.120 Cobalt-chromium metal wear particles are extremely small – only nanometre in size121 – and at high concentration are cytotoxic to cells, causing tissue necrosis121 due to release of metal ions. These small particles are distributed widely and elevated metal ion levels have been found in blood, serum and urine.122,123 The clinical risks of metal-on-metal bearings have been reviewed.124 More recently, larger-diameter metal-on-metal surface replacement bearings have been developed.125 These are bone-preserving prostheses and are recommended for younger active patients. In theory the larger diameters may encourage lubrication and reduce wear, but this will be dependent on correct design and geometrical clearances.126 However, without improved lubrication a larger diameter will increase wear due to a larger sliding distance. As with other new bearing couples longer-term clinical results are awaited with the metal-on-metal surface replacement hips.

Knee

Advances in bearing designs and material developments have not been as rapid in the knee as the hip. At present nearly all knee prostheses comprise a polyethylene tibial bearing with a cobalt-chromium metal femoral component and a metal tibial tray.89 A substantial advance has been made in the last 5 years with the introduction of stabilised and oxidation-resistant polyethylene, which inhibits delamination failure.89 This will also reduce surface wear and osteolytic potential.88-91,127 Cross-linked polyethylene is also being considered for use in the knee. However, there remain concerns about this as the cross-linking reduces the toughness and fatigue-resistance of the polyethylene, and this may influence longevity in the more highly stressed knee components. There have also been significant developments in knee-bearing design in recent years. The recognition that increased activity and kinematic inputs can markedly increase wear89-91 and osteolytic potential92 has led to an increase in the use of rotating platform mobile bearing knees,90,127 which decouple the motions reducing rotation and wear at the femoral tibial interface by allowing rotation of the polyethylene insert at the tibial tray. Rotating platform and mobile bearing designs are being introduced for a number of different types of knee prostheses.

Equally important developments in recent years involve the unicompartmental knee joint replacement.13 Current unicompartmental knees involve less or minimally invasive surgery128 and preserve bone. Long-standing questions remain as to when to perform unicompartmental knee joint replacements129 and the subsequent progression of arthritis.130

Improvements in design can have considerable impact on function,64 and fluoroscopic investigations131 are now providing real insights into the effect of different designs on kinematic function in the knee.

Conclusions

Joint replacement has been one of the major successes in surgery over the last 40 years. Clinical success, long-term results and failure, and increased expectation and lifetimes of patients have driven the need for improved materials, bearing surfaces and designs. Joint replacement is a complex interactive biomechanical and biological system, and, while extensive preclinical development and evaluation is now undertaken, long-term clinical results remain the ultimate test for new designs, materials and bearings. For many patients, particularly low-demand patients over 65 years old, tried and tested stabilised polyethylene bearings articulating against polished metal femoral components have a very high probability of providing 20 years' successful clinical use. For younger, more active and demanding patients, a range of improved designs, materials and bearings is available for clinical use in the hip and the knee. Many of these devices are supported by preclinical simulation tests which indicate improved performance compared to traditional technologies. The ultimate test is the long-term clinical follow-up. Until this is established there will always remain a degree of uncertainty surrounding any new technology for joint replacements. However, for many younger patients the potential long-term benefits outweigh the uncertainties, and the new technological solutions are being introduced widely for high-demand patients. They do, however, require rigorous and effective clinical follow-up. Over the last decade there have been considerable advances in design, materials and technology of hip and knee prostheses. In the future, advances in computer-assisted orthopaedic surgery (CAOS) and minimally invasive surgery (MIM) are likely to be more widely introduced into clinical practice.

Acknowledgement

The Institute of Medical and Biological Engineering, University of Leeds has received support for its research into artificial joints from the Arthritis Research Campaign (arc) and the Engineering and Physical Sciences Research Council.

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