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A five-station hip joint simulator for wear rate studies V Saikko, MSc Laboratory of Machine Design, Helsinki University of Technology, Espoo, Finland P Paavolainen, MD, PhD Orthopaedic Hospital of the Invalid Foundation, Helsinki, Finland
M Kleimola, PhD Laboratory of Machine Design, Helsinki University of Technology,Espoo, Finland P Slatis, MD, PhD
Orthopaedic Hospital of the Invalid Foundation, Helsinki, Finland
The aim of the work has been the development o j a hip joint simulator for comparative wear rate studies of long duration. Ajive-station apparatus has been designed, constructed and tested. Five total hip joints can be tested at the same time in identical conditions. The flexion-extension motion and the superior-inferior component of the joint contact force are incorporated. The motion is electromechanical and the loading pneumatic. The angle and load waveforms arefixed and simulate level walking. For accurate wear measurements each station employs a control prosthesis. The conditions of the control prosthesis in regard to loading, exposure to lubricant and environment temperature (37 f lac)are identical to those of the test prosthesis. The acetabular cups can be readily removed for periodic wear measurements and reassembled in exactly the original position. Extensive tests have shown that the simulator is a practical and reliuble instrument in the wear rate studies of various designs of total hip joint.
1 INTRODUCTION
The success of the total hip replacement is limited by late loosening of the femoral and acetabular components. The loosening process is assumed to be a biological response to wear and corrosion products of the implant. To avoid the use of patients in the testing of tribological performance of total hip joints, hip joint simulators are a necessity. A hip joint simulator is a laboratory apparatus that simulates the conditions of the hip joint in regard to motion, load, lubrication and temperature. Hip joint simulators that have been used for wear studies are described by Duff-Barclay and Spillman (l), Dowson et al. (Z), Dumbleton et al. (3), Walker and Salvati (4), Weightman et al. (5), Swanson et al. (6), Ungethiim et al. (7), Beutler et aE. (8), Swikert and Johnson (9), Cappozzo et al. (lo), Wright and Scales (II), Dumbleton (12), McKellop and Clarke (13) and Dowson and Jobbins (14). There are apparently not very many laboratories currently working on this laborious subject. It is obvious that the present state of knowledge of the wear phenomena of total hip joints is still inadequate, and consequently more simulators and wear studies are needed. The duration of a proper wear test is long, at least a few weeks, and so a multi-station apparatus or several single-station apparatuses are necessary. In this paper, a new five-station hip joint simulator is described. The motion, load, attachment of femoral heads and acetabular cups, lubrication system, temperature control, wear measurement and safety cut-outs are described in detail, and a summary of wear rate results is presented. The simulator was designed, constructed and tested in the Laboratory of Machine Design at Helsinki University of Technology. The main features of the apparatus are The M S was received on 13 March I992 and w a ~uccepted for publication on 13 Ortober 1992 H01692 0 IMechE 1992
in accord with I S 0 TR 9325 (15). The layout of the simulator showing the centre station is depicted in Fig. 1, Fig. 2 being a general view of the simulator. The apparatus does not include a frictional torque measurement system, because a separate servohydraulic microcomputer-controlled hip joint simulator, which will be described in another paper, was constructed especially for frictional torque studies. 2 MOTION
The motion is uniaxial and simulates the flexionextension of the hip joint in level walking. The abduction-adduction and the internal-external rotation are distinctly smaller in level walking and are omitted to avoid excessive complexity of the apparatus. The motion is implemented by means of an electric motor (Fig. 1, part 16), gear (part 15), crank (part 14) and linear bearing (part 13). The angular velocity o of the crank is 6.79 rad/s and so the cycle frequency of the simulator becomes 1.08 s-'. The structure of the crank mechanism is such that the resulting flexion-extension angle of the femoral head a(t) = arctan {sin at/ (2 - cos cot)}, where t is time. The range of motion becomes z/3,which is about 15 per cent more than the motion measured for normal subjects by Johnston and Smidt (16). The range of motion is increased with a view to compensating for the part of frictional energy that is attributable to the abduction-adduction and the internakxternal rotation. The maximum extension and flexion velocities are 2.26 and 6.79 rad/s respectively. The duration of the extension is 66.7 per cent of the length of the cycle, which is close to that measured for normal subjects by Johnston and Smidt. The motion is transmitted from the centre cradle (Fig. 1, part 10) to the other four by means of transmission levers (part 12) and connecting rods. The range of motion was checked with a clinometer.
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I Femoral head
2 Acetabular cup 3 Femoral stem 4 Adjustable femoral stem holder 5 Clamp 6 Acetabular cup holder 7 Femoral head of control prosthesis 8 Acetabular cup of control prosthesis 9 Water replenishmcnt bottle 10 Cradle I 1 Support bearing
12 Motion transmission lcver 13 Linear bearing 14 Crank 15 Gear 16 Electric motor 17 Loading arm 18 Pneumatic cylinder 19 Counterweight of loading arm 20 Pivot of loading arm 21 Acrylic hood 22 Connecting part
Fig. 1 Layout of five-station hip joint simulator from anterior-posterior view showing centre station
3 AlTACHMENT OF THE FEMORAL HEADS
The femoral head (Fig. 1, part 1) is attached to the cradle by means of the femoral stem (part 3). The stem is attached to the cradle by means of an adjustable stem holder (part 4) and a clamp (part 5). The stem holder is cast and machined of low-temperature melting aluminium-zinc alloy. The alignment of the femoral
head with respect to the axis of swinging is made possible by the two degrees of freedom of the cylindrical stem holder with respect to the cradle prior to clamping. The alignment of the head is checked with displacement indicators in the superior-inferior and in the anterior-posterior directions. The permissible ranges of radial displacements in the alignment procedure are 0.03 mm as the cradle is swung manually f30". The
Fig. 2 Five-station hip joint simulator. Part numbering as in Fig. 1 Part H : Journal of Engineering in Medicine
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implemented by the simulator. The oscilloscope measurements on the print refer to the load signal; 1 V (1 division) represents 1 kN. The load was measured so that one of the femoral components together with its cradle was replaced with a fixed frame, to which a force transducer was attached. A femoral head was mounted on the lower end of the force transducer. The flexionextension angle was measured simultaneously by an angle transducer from one of the cradles; 5 V (1 division) on the print represents 15".
HI=!
5 ATTACHMENT OF THE ACETABULAR CUPS
1
Fig. 3 Oscilloscope print of variation of flexionxxtension angle and superior-inferior load (nearly square wave) with time in simulation. Heel strike is located at the end of the zero-load phase and toe-off at the start of the zero-load phase
misali nment of the head due to the loading is much greater. The greatest displacements are attributable to the elastic deformation of the neck of the stem. However, the effect of the misalignments on the load is negligible since the loading arrangement is self-centring (see Section 8). Thc exit of possible contaminants from the cradle-stem holder interface and from the stem holder-stem interface is prevented by a layer of silicone sealant (see Fig. 1). 4 LOAD
The load is also uniaxial and simulates the most important of the three orthogonal components of the hip joint contact force in level walking, viz. the superior-inferior. The anterior-posterior and the medial-lateral force components are distinctly smaller in level walking and are omitted to avoid excessive complexity of the apparatus. The omission hardly decreases the relevance of the data obtained, since the examination of acetabular cups removed from patients has revealed considerable scatter in the direction of wear (17). The load is implemented by a pneumatic cylinder (Fig. 1, part 18) and a pivoted loading arm (part 17). The loading arm magnifies the force of the pneumatic cylinder by a factor of 3.2. The loading is of the on/off type. Each station has a loading arrangement of its own. The load is switched on at maximum flexion and off at maximum extension. The load waveform was taken from Crowninshield et al. (18), except that the valley between the two peaks is omitted for simplicity. In their waveform the duration of the loaded part of the cycle, from heel strike to toe off, is about 63 per cent of the length of the cycle. The load applied to the hip joint is adjusted to 3.5 kN, which is five times body weight of 71.4 kg body mass. The direction of the load remains virtually vertical and stationary with respect to the acetabular cup (Fig. 1, part 2), regardless of the position of the cradle, thus truly simulating the superior-inferior component. The maximum load that can be applied is about 8 kN. Figure 3 is an oscilloscope print of the flexionextension angle and the load (nearly square wave)
The acetabular cup is attached to an acrylic holder (Fig. 1, part 6) so that the angle between the horizontal plane and the plane of the rim of the cup is 45". The Young's modulus of the acrylic (about 3 GPa) is close to that of cancellous bone [from 0.004 to 1.8 GPa (1911, which is important if the cup is not metal-backed. The thickness of the holder at the point of the load vector is 5 mm. Since the wear measurement is gravimetric, a substitutive method for the attachment of cups intended for cemented application was developed. The method does not hamper the removal of the cup from the holder for periodic weighings, does not affect the weight of the cup during any phase of the wear test procedure and makes it possible to reassemble the cup in exactly the original position. The cup is locked into the holder with two polyacetal pegs, 5.5 mm in diameter. They are drawn with a dashed line Fig. 1, and two of the pegs can partly be seen in Fig. 2. The shape of the recess of the holder is a hemisphere, the radius of which is equal to the radius of the exterior surface of the cup. All the protrusions on the exterior surface of the cup are machined off. Two 5.5 mm bores are machined into the cup for the pegs in the superior-inferior direction through the bores in the bottom of the holder. The bores are located away from the load-bearing area of the cup, symmetrically on both sides, so that the distance between the bores is r1 + r 2 , where r I is the radius of the interior surface of the cup and r2 is the radius of the exterior surface of the cup. Extensive tests with the simulator have shown that no perceivable wear takes place at the cupholder interface, nor at the c u p p e g interface. If the design has a metallic acetabular shell, the acetabular insert is simply put into the shell, which is attached to the holder by means of polymethylmethacrylate bone cement. The sharp snap-lock edges of the insert are cut off prior to the test because they are prone to become detached as the insert is removed for wear measurements, which would reduce the weight of the cup. 6 LUBRICATION
The acetabular cup holder also operates as a lubricant receptacle. The cup is immersed in the lubricant so that the articulation remains lubricated. Evaporation is compensated by water from the replenishment bottle (Fig. 1, part 9). The upside-down position of the prosthesis is necessary, because in the opposite position the articulation would be prone to drying out due to air bubbles which gather in the contact area during the test. The only metallic parts with which the lubricant is in contact are those of the prosthesis.
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7 TEMPERATURE
The simulator is covered by an acrylic hood (Fig. 1, part 21). The temperature of the air within the hood is maintained at 37 f 1°C by means of two hot-air fans, a relay and a temperature sensor. The fans can be seen in the right of Fig. 2. The hood can be seen only partly in Fig. 2, because it has been raised. The hood also protects the prostheses against airborne contaminants.
8 WEAR MEASUREMENT
In order to perform accurate wear measurements for ultra-high molecular weight polyethylene (UHMWPE) acetabular cups, control cups are needed. Each test station of the simulator employs a control prosthesis-a unique feature. The acetabular cup of the control prosthesis (Fig. 1, part 8) and the femoral head of the control prosthesis (part 7) are of the same design as those of the test prosthesis. The conditions of the control cup in regard to loading, exposure to lubricant and environment temperature are identical to those of the test cup. The only significant difference is that there is no sliding motion in the control prosthesis. The method of wear measurement is gravimetric and so the amount of water absorbed by the test cup must be discovered. It is done by means of the control cup. In some prosthesis designs the weight loss due to wear proved to be less than the weight gain due to water absorption. Loading conditions must be identical, because the absorption rate proved to be affected by the cyclic loading. Only the effect of sliding motion on the absorption rate remains unquantified. Evaporation from the receptacle of the control prosthesis is prevented by a polyvinyl chloride (PVC) membrane, which covers the receptacle (see Fig. 2). Prior to the weighings, the wear debris is first rinsed out with distilled water and stored together with the lubricant from the receptacle for later analysis. Then the test and the soak control cups are vacuum desiccated for half an hour to evaporate the water off the surfaces. There is no need for a complete desiccation, which would take much too long. The cups are weighed with a Mettler AT261 DeltaRange analytical balance to the nearest 0.01 mg. The weight loss due to wear is assumed to be, according to the recommendation of I S 0 TR 9326 (20), ( M , , - M , - M , , M,), where M,, is the initial mass of the test acetabular cup, M , is the mass of the test acetabular cup, M,, is the initial mass of the soak control cup and M , is the mass of the soak control cup. The repetition of the weighings at intervals of a few minutes showed that the weight change of the test cup from the previous weight, apparently attributable to the humidity of the atmosphere, was small and equal to the weight change of the soak control cup, and so the wear value given by (M,, - M , - M,, M,) remained unchanged. This indicates that the gravimetric method of wear measurement is a sound one. Considering also the accuracy of the analytical balance, the error of the values of wear of acetabular cups obtained using the method is estimated to be kO.1 mg at most. In addition, the control prosthesis renders the loading arrangement self-centring. Rotation about the vertical axis of the connecting part (Fig. 1, part 22), to which the holder of the test cup and the femoral head of the
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control prosthesis are attached, is prevented. The leaning of the connecting part, which is mainly due to the deformations of the prostheses caused by the loading, is small, because the deformations are much less than the distance between the two femoral heads. The leaning due to the frictional torque is also small, because the coefficient of friction is usually in the order of 0.05. Consequently, the effect of the leaning of the connecting part on the load, on the relative position between the head and the cup, and on the resulting tribological behaviour of the hip joint is apparently negligible. The connecting part is machined of stainless steel or titanium. If the method of wear measurement were based on dimensional changes, the proportion of creep could be discovered by means of the loaded control cups. 9 CUT-OUTS
The operational reliability of the simulator is of the utmost importance, since the apparatus runs 24 hours a day, seven days a week, mostly without surveillance. The simulator has cut-outs, which interrupt the test automatically, if one of the following conditions is met : 1. A load-on or a load-off transition does not occur. The total number of load-on and load-off transitions in the five-station apparatus is ten per cycle, 930 528 per 24 hours. The length of the stroke of the pneumatic cylinder during the simulation is usually 23 mm, caused mainly by the deformation of the prostheses. Included in the stroke is the magnification by a factor of 3.2 caused by the loading arm. The reciprocating motion of the loading arm makes a microswitch, mounted above the cylinder, go on and off. These on/off and off/on transitions are monitored by a programmable logic controller, and even a single missing transition causes an interruption of the test. If the load was stuck in the on position, the joint would be rapidly destroyed due to overheating; if it was stuck in the off position, the wear would stop. The system also senses the loss of pneumatic pressure. 2. An excessive rise of a loading arm during the stance phase occurs. Another microswitch is mounted above the pneumatic cylinder a few millimetres higher than the load transition detector to sense yielding of some of the numerous loaded parts of the system, for example the neck of the femoral stem, the clamp of the stem holder, the cradle or the loading arm. 3. Excessive loading of the electric motor occurs, which is likely to be an indication of a jam in the motion mechanism. A test period is normally ended by a timer. The number of cycles is checked from a cycle counter. Regardless of the reason for the stop, the load is switched off automatically so that the undesirable creep deformation of the UHMWPE cups is avoided. In the case of an interruption of electric power, the load is switched off by a battery back-up, 10 TESTS
Five different total hip systems have been under scrutiny. The wear of UHMWPE cups articulating against
Part H Journal of Fngmeenng in Medicine
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Table 1 Wear rates of UHMWPE cups in five-station hip joint simulator; three of each head--cup
combination were worn
mm
Wear rates mg/1o6 cycles
Stainless steel (35.32.01) Alumina (12.32.06)
32 32
176, 146, 212 0.3, 2.1, 0.2
104308
Ti-6A1-4V, ion implanted (163185) Alumina (131407)
32 32
0.3, 0.1, 0.2 0.4, 11.6, 5.0
Link
102-1 10
Co-Cr-Mo (128-705) Alumina (128-709)
32 32
39.7, 48.2, 24.0 1.5, 0.5, 0.0
Howmedka
6285-0-525
Co-Cr-Mo (6284-0-132) Alumina (6290-1-032)
32 32
2.6, 4.7, 4.3 0.1, 0.0, 0.0
Thackray
62-37 17
Stainless steel (62-5671) Alumina (62-6392)
22.2 22.2
50.9, 62.4, 46.0 0.8, 1.0, 0.7
Manufacturer of prosthesis
Catalogue number (CUP)
Protek
62.32.50
Biomet
Head material and catalogue number
modular metallic and modular alumina ceramic heads was measured. The test with each combination was performed three times and so the total number of prostheses tested was 30. In two total hip systems the metallic head material was Co-Cr-Mo, in two stainless steel and in one ion implanted Ti-6A1-4V. The duration of one test was three million cycles and so the total running time for 18 million cycles was 4643 hours. The lubricant was distilled, deionized water. Wear measurement was performed at half a million cycle intervals. A summary of the wear rates calculated by the method of least squares linear regression is presented in Table 1. A detailed report of the tests forms a separate paper that has been submitted to another journal. The mean wear rates ranged from virtually zero (Co-Cr-Mo-backed cup articulating against an alumina head) to 178 mg/ 106 cycles (cup articulating against a stainless steel head). The observed superiority of alumina is consistent with clinical findings (21), although clinical findings show a less pronounced difference. The excellent behaviour of alumina is apparently attributable to its superior resistance to abrasion and corrosion, and good wettability that decreases adhesive wear of UHMWPE. Stainless steel heads were corroded, which explains the high wear rates of the cups they articulated against. As for the tribological study of prostheses that are or have been in patients, the most studied design is apparently Charnley. The mean wear rate obtained on the simulator for the modular prostheses with 22.2 mm head diameter corresponds to 0.15 mm/yr, assuming that the penetration is the wear volume (= mass loss/0.94 mg/ mm3) divided by the projected area of the head and that lo6 cycles in the simulator corresponds to one year in vivo. The value is not too far from 0.21 mm/yr, the mean value for 87 retrieved Charnley cups, obtained by measuring the internal volume changes (22). In addition, the third test with alumina heads, started in November 1991, has been considerably extended and it reached 24 million cycles in September 1992, which is apparently a world record. 11 CONCLUSIONS
A five-station hip joint simulator for comparative wear rate studies of long duration has been designed, constructed and tested. The distressing problem of late loosening of femoral and acetabular components is apparently a biological response to wear and corrosion
Head diameter
products of the implant. By means of the simulator described in this paper, new information about wear of total hip joints can be obtained. The apparatus has proved to be a practical and reliable instrument in the wear studies of various designs of total hip joint. During the first 17 months of operation the running of the simulator was interrupted virtually only for periodic wear measurements. ACKNOWLEDGEMENTS
Two grants and a three-year post as assistant researcher awarded to Vesa Saikko by the Academy of Finland made this work possible and are gratefully acknowledged. Vesa Saikko thanks also the Emil Aaltonen Foundation and the Foundation of Technology for financial support. Mr. I. Pajamaki and his associates in the workshop of the Machine Laboratory Building, Helsinki University of Technology, skilfully machined the numerous parts of the simulator. REFERENCES Duff-Barclay, I. and Spillman, D. T. Total human hip joint prostheses-a laboratory study of friction and wear. Proceedings of an IMechE Symposium on Lubrication and wear in living and artifkid human joints, London, 1967, Vol. 181, Part 35, pp. 90-103 (The Institution of Mechanical Engineers, London). Dowson, D., Walker, P. S., Longfield, M. D. and Wright, V. A joint simulating machine for load-bearing joints. Med. Biolog. Engng, 1970,8(1), 37-43. Dumbleton, J. H., Miller, D. A. and Miller, E. H. A simulator for load bearing joints. Wear, 1972,20(2), 165-174. Walker, P. S. and Salvati, E. The measurement and effects of friction and wear in artificial hip joints. Journal of Biomedical Materials Research, Biomedical Materials Symposium 4 on Materials and design considerationsfor the attachment of prostheses to the musculo-skeletal system, Clemson, South Carolina, 1972, pp. 327-342 (John Wiley, New York). Weightman, B., Simon, S., Paul, I., Rose, R. and Radin, E. Lubrication mechanism of hip joint replacement prostheses. Trans. ASME, J . Lubric. Technol., 1972,94(2), 131-135. Swanson, S. A. V., Freeman, M. A. R. and Heath, J. C. Laboratory tests on total joint replacement prostheses. .I. Bone Jt Surg., 1973, 558(4), 759 773. Ungethiim, M., Hildebrandt, J., Jager, M. and Mode, H. G. Ein neuer Simulator zur Testung von Totalendoprothesen fur das Hiiftgelenk. Archiv f i r orthopadische und Unfall-Chirurgie, 1973, 77(4), 304-3 14. Beutler, H., Lehmann, M. and Stahli, G. Wear behaviour of medical engineering materials. Wear, 1975,33(2), 337-350. Swikert, M. A. and Johnson, R. L. Simulated studies of wear and friction in total hip prosthesis components with various ball sizes
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and surface finishes, NASA TN D-8174, 1976 (National Aeronautics and Space Administration, Washington, D.C.). 10 Cappozzo, A., Cini, L., Pizzoferrato, A., Trentani, C. and Cortesi, S. S. Evaluation of hip arthroproslheses by means of body environment simulators. J . Biomed. Mater. Res., 1977, 11(5), 657669. 11 Wright, K. W. J. and Scales, J. T. The use of hip joint simulators for the evaluation of wear of total hip prostheses. First European Conference on Biomaferials, evaluation of biomaterials, Strasbourg, 1977, pp. 135-146 (John Wiley, Chichester). 12 Dumbleton, J. H. Tribology of natural and arti$cial joints, 1981, pp. 263-267 (Elsevier Scientific, Amsterdam). 13 McKellop, H. A. and Clarke, 1. C. Evolution and evaluation of materials-screening machines and joint simulators in predicting in vivo wear phenomena. In Functional behavior of orthopedic biomaterials, Vol. 11, Ch. 3, 1984, pp. 51-85 (CRC Press, Boca Raton, Ha.). 14 Dowson, D. and Jobbins, B. Design and development of a versatile hip joint simulator and a preliminary assessment of wear and creep in Charnley total replacement hip joints. Engng in Medicine, 1988, 17(3), 1-6. 15 I S 0 TR 9325-1989 Implants for surgery-partial and total hip joint prostheses-recommendations for simulators for evaluation of
hip joint prostheses, 1989 (International Organization for Standardization), 16 Johnston, R. C. and Smidt, G. L. Measurement of hip-joint motion during walking. J. Bone Jt Surg., 1969,51A(6),1083-1094. 17 Wroblewski, B. M. Direction and rate of socket wear in Charnley low-friction arthroplasty. J . Bone Jt Surg., 1985,67B(5),757-761. 18 Crowninshield, R. D., Johnston, R. C., Andrews, J. G. and Brand, R. A. A biomechanical investigation of the human hip. J . Biomechanics, 1978, 11(2),75-85. 19 Hodgskinson, R. and Currey, J. D. The effect of variation in structure on the Young's modulus of cancellous bone: a comparison of human and non-human material. Proc. Instn Mech. Engrs, Part H , l990,204(H2), 115-121. 20 I S 0 TR 9326-1989 Implants for surgery--partial and total hip joint prostheses-guidance for laboratory evaluation of change of form of bearing surfaces, 1989 (International Organization for Standardization). 21 Oonishi, H., Tsji, E, Hanatate, Y. and Mizokoshi, T. Tribological studies on retrieved total joint prostheses outlines. Jap. 1. Tribol., 1991,36(12),1345-1355. 22 Isaac, G. H., Wroblewski, B. M, Atkinson, J. R. and Dowson, D. A tribological study of retrieved hip prostheses. Clin. Orthop. Related Res., 1992,216, 115-125.
Part H: Journal of Engineering in Medicine
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A five-station hip joint simulator for wear rate studies V Saikko, MSc Laboratory of Machine Design, Helsinki University of Technology, Espoo, Finland P Paavolainen, MD, PhD Orthopaedic Hospital of the Invalid Foundation, Helsinki, Finland
M Kleimola, PhD Laboratory of Machine Design, Helsinki University of Technology,Espoo, Finland P Slatis, MD, PhD
Orthopaedic Hospital of the Invalid Foundation, Helsinki, Finland
The aim of the work has been the development o j a hip joint simulator for comparative wear rate studies of long duration. Ajive-station apparatus has been designed, constructed and tested. Five total hip joints can be tested at the same time in identical conditions. The flexion-extension motion and the superior-inferior component of the joint contact force are incorporated. The motion is electromechanical and the loading pneumatic. The angle and load waveforms arefixed and simulate level walking. For accurate wear measurements each station employs a control prosthesis. The conditions of the control prosthesis in regard to loading, exposure to lubricant and environment temperature (37 f lac)are identical to those of the test prosthesis. The acetabular cups can be readily removed for periodic wear measurements and reassembled in exactly the original position. Extensive tests have shown that the simulator is a practical and reliuble instrument in the wear rate studies of various designs of total hip joint.
1 INTRODUCTION
The success of the total hip replacement is limited by late loosening of the femoral and acetabular components. The loosening process is assumed to be a biological response to wear and corrosion products of the implant. To avoid the use of patients in the testing of tribological performance of total hip joints, hip joint simulators are a necessity. A hip joint simulator is a laboratory apparatus that simulates the conditions of the hip joint in regard to motion, load, lubrication and temperature. Hip joint simulators that have been used for wear studies are described by Duff-Barclay and Spillman (l), Dowson et al. (Z), Dumbleton et al. (3), Walker and Salvati (4), Weightman et al. (5), Swanson et al. (6), Ungethiim et al. (7), Beutler et aE. (8), Swikert and Johnson (9), Cappozzo et al. (lo), Wright and Scales (II), Dumbleton (12), McKellop and Clarke (13) and Dowson and Jobbins (14). There are apparently not very many laboratories currently working on this laborious subject. It is obvious that the present state of knowledge of the wear phenomena of total hip joints is still inadequate, and consequently more simulators and wear studies are needed. The duration of a proper wear test is long, at least a few weeks, and so a multi-station apparatus or several single-station apparatuses are necessary. In this paper, a new five-station hip joint simulator is described. The motion, load, attachment of femoral heads and acetabular cups, lubrication system, temperature control, wear measurement and safety cut-outs are described in detail, and a summary of wear rate results is presented. The simulator was designed, constructed and tested in the Laboratory of Machine Design at Helsinki University of Technology. The main features of the apparatus are The M S was received on 13 March I992 and w a ~uccepted for publication on 13 Ortober 1992 H01692 0 IMechE 1992
in accord with I S 0 TR 9325 (15). The layout of the simulator showing the centre station is depicted in Fig. 1, Fig. 2 being a general view of the simulator. The apparatus does not include a frictional torque measurement system, because a separate servohydraulic microcomputer-controlled hip joint simulator, which will be described in another paper, was constructed especially for frictional torque studies. 2 MOTION
The motion is uniaxial and simulates the flexionextension of the hip joint in level walking. The abduction-adduction and the internal-external rotation are distinctly smaller in level walking and are omitted to avoid excessive complexity of the apparatus. The motion is implemented by means of an electric motor (Fig. 1, part 16), gear (part 15), crank (part 14) and linear bearing (part 13). The angular velocity o of the crank is 6.79 rad/s and so the cycle frequency of the simulator becomes 1.08 s-'. The structure of the crank mechanism is such that the resulting flexion-extension angle of the femoral head a(t) = arctan {sin at/ (2 - cos cot)}, where t is time. The range of motion becomes z/3,which is about 15 per cent more than the motion measured for normal subjects by Johnston and Smidt (16). The range of motion is increased with a view to compensating for the part of frictional energy that is attributable to the abduction-adduction and the internakxternal rotation. The maximum extension and flexion velocities are 2.26 and 6.79 rad/s respectively. The duration of the extension is 66.7 per cent of the length of the cycle, which is close to that measured for normal subjects by Johnston and Smidt. The motion is transmitted from the centre cradle (Fig. 1, part 10) to the other four by means of transmission levers (part 12) and connecting rods. The range of motion was checked with a clinometer.
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I Femoral head
2 Acetabular cup 3 Femoral stem 4 Adjustable femoral stem holder 5 Clamp 6 Acetabular cup holder 7 Femoral head of control prosthesis 8 Acetabular cup of control prosthesis 9 Water replenishmcnt bottle 10 Cradle I 1 Support bearing
12 Motion transmission lcver 13 Linear bearing 14 Crank 15 Gear 16 Electric motor 17 Loading arm 18 Pneumatic cylinder 19 Counterweight of loading arm 20 Pivot of loading arm 21 Acrylic hood 22 Connecting part
Fig. 1 Layout of five-station hip joint simulator from anterior-posterior view showing centre station
3 AlTACHMENT OF THE FEMORAL HEADS
The femoral head (Fig. 1, part 1) is attached to the cradle by means of the femoral stem (part 3). The stem is attached to the cradle by means of an adjustable stem holder (part 4) and a clamp (part 5). The stem holder is cast and machined of low-temperature melting aluminium-zinc alloy. The alignment of the femoral
head with respect to the axis of swinging is made possible by the two degrees of freedom of the cylindrical stem holder with respect to the cradle prior to clamping. The alignment of the head is checked with displacement indicators in the superior-inferior and in the anterior-posterior directions. The permissible ranges of radial displacements in the alignment procedure are 0.03 mm as the cradle is swung manually f30". The
Fig. 2 Five-station hip joint simulator. Part numbering as in Fig. 1 Part H : Journal of Engineering in Medicine
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.4 FIVE-STATION HIP JOINT SIMULATOR FOR WEAR RATE STUDIES
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implemented by the simulator. The oscilloscope measurements on the print refer to the load signal; 1 V (1 division) represents 1 kN. The load was measured so that one of the femoral components together with its cradle was replaced with a fixed frame, to which a force transducer was attached. A femoral head was mounted on the lower end of the force transducer. The flexionextension angle was measured simultaneously by an angle transducer from one of the cradles; 5 V (1 division) on the print represents 15".
HI=!
5 ATTACHMENT OF THE ACETABULAR CUPS
1
Fig. 3 Oscilloscope print of variation of flexionxxtension angle and superior-inferior load (nearly square wave) with time in simulation. Heel strike is located at the end of the zero-load phase and toe-off at the start of the zero-load phase
misali nment of the head due to the loading is much greater. The greatest displacements are attributable to the elastic deformation of the neck of the stem. However, the effect of the misalignments on the load is negligible since the loading arrangement is self-centring (see Section 8). Thc exit of possible contaminants from the cradle-stem holder interface and from the stem holder-stem interface is prevented by a layer of silicone sealant (see Fig. 1). 4 LOAD
The load is also uniaxial and simulates the most important of the three orthogonal components of the hip joint contact force in level walking, viz. the superior-inferior. The anterior-posterior and the medial-lateral force components are distinctly smaller in level walking and are omitted to avoid excessive complexity of the apparatus. The omission hardly decreases the relevance of the data obtained, since the examination of acetabular cups removed from patients has revealed considerable scatter in the direction of wear (17). The load is implemented by a pneumatic cylinder (Fig. 1, part 18) and a pivoted loading arm (part 17). The loading arm magnifies the force of the pneumatic cylinder by a factor of 3.2. The loading is of the on/off type. Each station has a loading arrangement of its own. The load is switched on at maximum flexion and off at maximum extension. The load waveform was taken from Crowninshield et al. (18), except that the valley between the two peaks is omitted for simplicity. In their waveform the duration of the loaded part of the cycle, from heel strike to toe off, is about 63 per cent of the length of the cycle. The load applied to the hip joint is adjusted to 3.5 kN, which is five times body weight of 71.4 kg body mass. The direction of the load remains virtually vertical and stationary with respect to the acetabular cup (Fig. 1, part 2), regardless of the position of the cradle, thus truly simulating the superior-inferior component. The maximum load that can be applied is about 8 kN. Figure 3 is an oscilloscope print of the flexionextension angle and the load (nearly square wave)
The acetabular cup is attached to an acrylic holder (Fig. 1, part 6) so that the angle between the horizontal plane and the plane of the rim of the cup is 45". The Young's modulus of the acrylic (about 3 GPa) is close to that of cancellous bone [from 0.004 to 1.8 GPa (1911, which is important if the cup is not metal-backed. The thickness of the holder at the point of the load vector is 5 mm. Since the wear measurement is gravimetric, a substitutive method for the attachment of cups intended for cemented application was developed. The method does not hamper the removal of the cup from the holder for periodic weighings, does not affect the weight of the cup during any phase of the wear test procedure and makes it possible to reassemble the cup in exactly the original position. The cup is locked into the holder with two polyacetal pegs, 5.5 mm in diameter. They are drawn with a dashed line Fig. 1, and two of the pegs can partly be seen in Fig. 2. The shape of the recess of the holder is a hemisphere, the radius of which is equal to the radius of the exterior surface of the cup. All the protrusions on the exterior surface of the cup are machined off. Two 5.5 mm bores are machined into the cup for the pegs in the superior-inferior direction through the bores in the bottom of the holder. The bores are located away from the load-bearing area of the cup, symmetrically on both sides, so that the distance between the bores is r1 + r 2 , where r I is the radius of the interior surface of the cup and r2 is the radius of the exterior surface of the cup. Extensive tests with the simulator have shown that no perceivable wear takes place at the cupholder interface, nor at the c u p p e g interface. If the design has a metallic acetabular shell, the acetabular insert is simply put into the shell, which is attached to the holder by means of polymethylmethacrylate bone cement. The sharp snap-lock edges of the insert are cut off prior to the test because they are prone to become detached as the insert is removed for wear measurements, which would reduce the weight of the cup. 6 LUBRICATION
The acetabular cup holder also operates as a lubricant receptacle. The cup is immersed in the lubricant so that the articulation remains lubricated. Evaporation is compensated by water from the replenishment bottle (Fig. 1, part 9). The upside-down position of the prosthesis is necessary, because in the opposite position the articulation would be prone to drying out due to air bubbles which gather in the contact area during the test. The only metallic parts with which the lubricant is in contact are those of the prosthesis.
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7 TEMPERATURE
The simulator is covered by an acrylic hood (Fig. 1, part 21). The temperature of the air within the hood is maintained at 37 f 1°C by means of two hot-air fans, a relay and a temperature sensor. The fans can be seen in the right of Fig. 2. The hood can be seen only partly in Fig. 2, because it has been raised. The hood also protects the prostheses against airborne contaminants.
8 WEAR MEASUREMENT
In order to perform accurate wear measurements for ultra-high molecular weight polyethylene (UHMWPE) acetabular cups, control cups are needed. Each test station of the simulator employs a control prosthesis-a unique feature. The acetabular cup of the control prosthesis (Fig. 1, part 8) and the femoral head of the control prosthesis (part 7) are of the same design as those of the test prosthesis. The conditions of the control cup in regard to loading, exposure to lubricant and environment temperature are identical to those of the test cup. The only significant difference is that there is no sliding motion in the control prosthesis. The method of wear measurement is gravimetric and so the amount of water absorbed by the test cup must be discovered. It is done by means of the control cup. In some prosthesis designs the weight loss due to wear proved to be less than the weight gain due to water absorption. Loading conditions must be identical, because the absorption rate proved to be affected by the cyclic loading. Only the effect of sliding motion on the absorption rate remains unquantified. Evaporation from the receptacle of the control prosthesis is prevented by a polyvinyl chloride (PVC) membrane, which covers the receptacle (see Fig. 2). Prior to the weighings, the wear debris is first rinsed out with distilled water and stored together with the lubricant from the receptacle for later analysis. Then the test and the soak control cups are vacuum desiccated for half an hour to evaporate the water off the surfaces. There is no need for a complete desiccation, which would take much too long. The cups are weighed with a Mettler AT261 DeltaRange analytical balance to the nearest 0.01 mg. The weight loss due to wear is assumed to be, according to the recommendation of I S 0 TR 9326 (20), ( M , , - M , - M , , M,), where M,, is the initial mass of the test acetabular cup, M , is the mass of the test acetabular cup, M,, is the initial mass of the soak control cup and M , is the mass of the soak control cup. The repetition of the weighings at intervals of a few minutes showed that the weight change of the test cup from the previous weight, apparently attributable to the humidity of the atmosphere, was small and equal to the weight change of the soak control cup, and so the wear value given by (M,, - M , - M,, M,) remained unchanged. This indicates that the gravimetric method of wear measurement is a sound one. Considering also the accuracy of the analytical balance, the error of the values of wear of acetabular cups obtained using the method is estimated to be kO.1 mg at most. In addition, the control prosthesis renders the loading arrangement self-centring. Rotation about the vertical axis of the connecting part (Fig. 1, part 22), to which the holder of the test cup and the femoral head of the
+
+
control prosthesis are attached, is prevented. The leaning of the connecting part, which is mainly due to the deformations of the prostheses caused by the loading, is small, because the deformations are much less than the distance between the two femoral heads. The leaning due to the frictional torque is also small, because the coefficient of friction is usually in the order of 0.05. Consequently, the effect of the leaning of the connecting part on the load, on the relative position between the head and the cup, and on the resulting tribological behaviour of the hip joint is apparently negligible. The connecting part is machined of stainless steel or titanium. If the method of wear measurement were based on dimensional changes, the proportion of creep could be discovered by means of the loaded control cups. 9 CUT-OUTS
The operational reliability of the simulator is of the utmost importance, since the apparatus runs 24 hours a day, seven days a week, mostly without surveillance. The simulator has cut-outs, which interrupt the test automatically, if one of the following conditions is met : 1. A load-on or a load-off transition does not occur. The total number of load-on and load-off transitions in the five-station apparatus is ten per cycle, 930 528 per 24 hours. The length of the stroke of the pneumatic cylinder during the simulation is usually 23 mm, caused mainly by the deformation of the prostheses. Included in the stroke is the magnification by a factor of 3.2 caused by the loading arm. The reciprocating motion of the loading arm makes a microswitch, mounted above the cylinder, go on and off. These on/off and off/on transitions are monitored by a programmable logic controller, and even a single missing transition causes an interruption of the test. If the load was stuck in the on position, the joint would be rapidly destroyed due to overheating; if it was stuck in the off position, the wear would stop. The system also senses the loss of pneumatic pressure. 2. An excessive rise of a loading arm during the stance phase occurs. Another microswitch is mounted above the pneumatic cylinder a few millimetres higher than the load transition detector to sense yielding of some of the numerous loaded parts of the system, for example the neck of the femoral stem, the clamp of the stem holder, the cradle or the loading arm. 3. Excessive loading of the electric motor occurs, which is likely to be an indication of a jam in the motion mechanism. A test period is normally ended by a timer. The number of cycles is checked from a cycle counter. Regardless of the reason for the stop, the load is switched off automatically so that the undesirable creep deformation of the UHMWPE cups is avoided. In the case of an interruption of electric power, the load is switched off by a battery back-up, 10 TESTS
Five different total hip systems have been under scrutiny. The wear of UHMWPE cups articulating against
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Table 1 Wear rates of UHMWPE cups in five-station hip joint simulator; three of each head--cup
combination were worn
mm
Wear rates mg/1o6 cycles
Stainless steel (35.32.01) Alumina (12.32.06)
32 32
176, 146, 212 0.3, 2.1, 0.2
104308
Ti-6A1-4V, ion implanted (163185) Alumina (131407)
32 32
0.3, 0.1, 0.2 0.4, 11.6, 5.0
Link
102-1 10
Co-Cr-Mo (128-705) Alumina (128-709)
32 32
39.7, 48.2, 24.0 1.5, 0.5, 0.0
Howmedka
6285-0-525
Co-Cr-Mo (6284-0-132) Alumina (6290-1-032)
32 32
2.6, 4.7, 4.3 0.1, 0.0, 0.0
Thackray
62-37 17
Stainless steel (62-5671) Alumina (62-6392)
22.2 22.2
50.9, 62.4, 46.0 0.8, 1.0, 0.7
Manufacturer of prosthesis
Catalogue number (CUP)
Protek
62.32.50
Biomet
Head material and catalogue number
modular metallic and modular alumina ceramic heads was measured. The test with each combination was performed three times and so the total number of prostheses tested was 30. In two total hip systems the metallic head material was Co-Cr-Mo, in two stainless steel and in one ion implanted Ti-6A1-4V. The duration of one test was three million cycles and so the total running time for 18 million cycles was 4643 hours. The lubricant was distilled, deionized water. Wear measurement was performed at half a million cycle intervals. A summary of the wear rates calculated by the method of least squares linear regression is presented in Table 1. A detailed report of the tests forms a separate paper that has been submitted to another journal. The mean wear rates ranged from virtually zero (Co-Cr-Mo-backed cup articulating against an alumina head) to 178 mg/ 106 cycles (cup articulating against a stainless steel head). The observed superiority of alumina is consistent with clinical findings (21), although clinical findings show a less pronounced difference. The excellent behaviour of alumina is apparently attributable to its superior resistance to abrasion and corrosion, and good wettability that decreases adhesive wear of UHMWPE. Stainless steel heads were corroded, which explains the high wear rates of the cups they articulated against. As for the tribological study of prostheses that are or have been in patients, the most studied design is apparently Charnley. The mean wear rate obtained on the simulator for the modular prostheses with 22.2 mm head diameter corresponds to 0.15 mm/yr, assuming that the penetration is the wear volume (= mass loss/0.94 mg/ mm3) divided by the projected area of the head and that lo6 cycles in the simulator corresponds to one year in vivo. The value is not too far from 0.21 mm/yr, the mean value for 87 retrieved Charnley cups, obtained by measuring the internal volume changes (22). In addition, the third test with alumina heads, started in November 1991, has been considerably extended and it reached 24 million cycles in September 1992, which is apparently a world record. 11 CONCLUSIONS
A five-station hip joint simulator for comparative wear rate studies of long duration has been designed, constructed and tested. The distressing problem of late loosening of femoral and acetabular components is apparently a biological response to wear and corrosion
Head diameter
products of the implant. By means of the simulator described in this paper, new information about wear of total hip joints can be obtained. The apparatus has proved to be a practical and reliable instrument in the wear studies of various designs of total hip joint. During the first 17 months of operation the running of the simulator was interrupted virtually only for periodic wear measurements. ACKNOWLEDGEMENTS
Two grants and a three-year post as assistant researcher awarded to Vesa Saikko by the Academy of Finland made this work possible and are gratefully acknowledged. Vesa Saikko thanks also the Emil Aaltonen Foundation and the Foundation of Technology for financial support. Mr. I. Pajamaki and his associates in the workshop of the Machine Laboratory Building, Helsinki University of Technology, skilfully machined the numerous parts of the simulator. REFERENCES Duff-Barclay, I. and Spillman, D. T. Total human hip joint prostheses-a laboratory study of friction and wear. Proceedings of an IMechE Symposium on Lubrication and wear in living and artifkid human joints, London, 1967, Vol. 181, Part 35, pp. 90-103 (The Institution of Mechanical Engineers, London). Dowson, D., Walker, P. S., Longfield, M. D. and Wright, V. A joint simulating machine for load-bearing joints. Med. Biolog. Engng, 1970,8(1), 37-43. Dumbleton, J. H., Miller, D. A. and Miller, E. H. A simulator for load bearing joints. Wear, 1972,20(2), 165-174. Walker, P. S. and Salvati, E. The measurement and effects of friction and wear in artificial hip joints. Journal of Biomedical Materials Research, Biomedical Materials Symposium 4 on Materials and design considerationsfor the attachment of prostheses to the musculo-skeletal system, Clemson, South Carolina, 1972, pp. 327-342 (John Wiley, New York). Weightman, B., Simon, S., Paul, I., Rose, R. and Radin, E. Lubrication mechanism of hip joint replacement prostheses. Trans. ASME, J . Lubric. Technol., 1972,94(2), 131-135. Swanson, S. A. V., Freeman, M. A. R. and Heath, J. C. Laboratory tests on total joint replacement prostheses. .I. Bone Jt Surg., 1973, 558(4), 759 773. Ungethiim, M., Hildebrandt, J., Jager, M. and Mode, H. G. Ein neuer Simulator zur Testung von Totalendoprothesen fur das Hiiftgelenk. Archiv f i r orthopadische und Unfall-Chirurgie, 1973, 77(4), 304-3 14. Beutler, H., Lehmann, M. and Stahli, G. Wear behaviour of medical engineering materials. Wear, 1975,33(2), 337-350. Swikert, M. A. and Johnson, R. L. Simulated studies of wear and friction in total hip prosthesis components with various ball sizes
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and surface finishes, NASA TN D-8174, 1976 (National Aeronautics and Space Administration, Washington, D.C.). 10 Cappozzo, A., Cini, L., Pizzoferrato, A., Trentani, C. and Cortesi, S. S. Evaluation of hip arthroproslheses by means of body environment simulators. J . Biomed. Mater. Res., 1977, 11(5), 657669. 11 Wright, K. W. J. and Scales, J. T. The use of hip joint simulators for the evaluation of wear of total hip prostheses. First European Conference on Biomaferials, evaluation of biomaterials, Strasbourg, 1977, pp. 135-146 (John Wiley, Chichester). 12 Dumbleton, J. H. Tribology of natural and arti$cial joints, 1981, pp. 263-267 (Elsevier Scientific, Amsterdam). 13 McKellop, H. A. and Clarke, 1. C. Evolution and evaluation of materials-screening machines and joint simulators in predicting in vivo wear phenomena. In Functional behavior of orthopedic biomaterials, Vol. 11, Ch. 3, 1984, pp. 51-85 (CRC Press, Boca Raton, Ha.). 14 Dowson, D. and Jobbins, B. Design and development of a versatile hip joint simulator and a preliminary assessment of wear and creep in Charnley total replacement hip joints. Engng in Medicine, 1988, 17(3), 1-6. 15 I S 0 TR 9325-1989 Implants for surgery-partial and total hip joint prostheses-recommendations for simulators for evaluation of
hip joint prostheses, 1989 (International Organization for Standardization), 16 Johnston, R. C. and Smidt, G. L. Measurement of hip-joint motion during walking. J. Bone Jt Surg., 1969,51A(6),1083-1094. 17 Wroblewski, B. M. Direction and rate of socket wear in Charnley low-friction arthroplasty. J . Bone Jt Surg., 1985,67B(5),757-761. 18 Crowninshield, R. D., Johnston, R. C., Andrews, J. G. and Brand, R. A. A biomechanical investigation of the human hip. J . Biomechanics, 1978, 11(2),75-85. 19 Hodgskinson, R. and Currey, J. D. The effect of variation in structure on the Young's modulus of cancellous bone: a comparison of human and non-human material. Proc. Instn Mech. Engrs, Part H , l990,204(H2), 115-121. 20 I S 0 TR 9326-1989 Implants for surgery--partial and total hip joint prostheses-guidance for laboratory evaluation of change of form of bearing surfaces, 1989 (International Organization for Standardization). 21 Oonishi, H., Tsji, E, Hanatate, Y. and Mizokoshi, T. Tribological studies on retrieved total joint prostheses outlines. Jap. 1. Tribol., 1991,36(12),1345-1355. 22 Isaac, G. H., Wroblewski, B. M, Atkinson, J. R. and Dowson, D. A tribological study of retrieved hip prostheses. Clin. Orthop. Related Res., 1992,216, 115-125.
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