- Email: [email protected]
Initial stability of fully and partially cemented femoral stems
Clinical Biomechanics 15 (2000) 750±755
www.elsevier.com/locate/clinbiomech
Initial stability of fully and partially cemented femoral stems Lutz Claes a,*, Stefan Fiedler a, Michael Ohnmacht a, Georg N. Duda b b
a Institut f ur Unfallchirurgische Forschung und Biomechanik, Universit at Ulm, Helmholtzstraûe 14, 89081 Ulm, Germany Virchow-Klinikum, Abteilung Unfallchirurgie und Wiederherstellungschirurgie, Humboldt Universit at zu Berlin, 13353 Berlin, Germany
Received 26 April 2000; accepted 13 July 2000
Abstract Objective. To test the initial stability of a newly designed partially cemented femoral stem in comparison with a fully cemented conventional stem. Design. An in vitro study to determine the interface motion between femoral stem and bone as a response to loading. Background. The aim of the new prosthesis design is a proximal load transfer by a de®ned partial cement ®xation in the proximal femur region and a slim prosthesis stem in the distal region. Before a clinical study can be started, the new stem has to show an initial stability comparable to that of fully cemented prostheses. Method. Six paired fresh cadaveric femora were used for the testing of the new partially cemented stem (Option 3000, Mathys Orthopaedics, Bettlach, Switzerland) and a fully cemented stem (Weber Shaft, AlloPro, Baar, Swizerland). Under cyclic loading up to 1600 N hip joint forces, the interface motion between implants and bone was measured at six locations. Results. Both stems showed uncritical interface motions below 43 lm. However, the Option 3000 stem exhibited signi®cantly smaller motions in the proximal region and slightly larger movements in the distal regions than the Weber prosthesis. Conclusions. The new type of partially cemented stem provided a comparable initial stability to the fully cemented Weber prosthesis. Relevance The high initial stability of the Option 3000 stem justi®ed the clinical use of the new implant. More than 100 implantations in the last three years, with very good preliminary clinical results, support the preclinical ®ndings. Ó 2000 Published by Elsevier Science Ltd. Keywords: Total hip replacement; Initial stability; Prosthesis design; Mechanical testing
1. Introduction Under physiological conditions, the loads acting on the femoral head are transferred to the femoral shaft by the trabecular system of the proximal femur [1,2]. Distally anchoraged femoral stems often demonstrate a proximal bone loss. One of the possible reasons for this eect is due to the unphysiological load transfer and corresponding stress shielding of the proximal femur region [1,2]. Recently, a new femoral stem was developed which is mainly ®xed in the proximal part of the femur by press-®t and partial
*
Corresponding author. E-mail address: [email protected] (L. Claes).
0268-0033/00/$ - see front matter Ó 2000 Published by Elsevier Science Ltd. PII: S 0 2 6 8 - 0 0 3 3 ( 0 0 ) 0 0 0 4 4 - 9
cementing (Option 3000, Mathys Orthopaedics, Bettlach, Switzerland, Patent PCT CH 95/00260). The distal uncemented part of this prosthesis is very slim, ®lls the intramedullary canal only partly, and avoids axial load transfer (Fig. 1). However, the question remained open whether this partially cemented stem would provide enough initial stability. It is well accepted that initial stability and lack of motion between bone and implant is one of the most important prerequisites for the long-term success of prostheses [3±5]. Therefore, one of the ®rst steps in the preclinical testing of new prostheses should be the measurement of the interface motion under sucient loading conditions. Such investigations have been carried out for cemented and cementless femoral stems by several groups [3±8]. It could be shown that the risk for interface motion and rotation was higher for cementless stems than
L. Claes et al. / Clinical Biomechanics 15 (2000) 750±755
751
metal cup using methylmethacrylate (Technovit 4040, Merck, Darmstadt, Germany). 2.2. Implants The new type of prosthesis (Option 3000, Mathys Orthopaedics, Bettlach, Switzerland) is designed to be ®xed only in the proximal region of the femur. It has an oval collarless cross-sectional design at the most proximal region and a very slim distal stem (Fig. 1). The small diameter of the distal stem avoids a distal ®tting and has a polished surface. In the mid-proximal region of the prosthesis, a circular groove allows the locally de®ned application of bone cement when the prosthesis is already press-®tted in the femur. The bone cement can be admitted by a longitudinal channel running from the proximal end of the prosthesis into the circular groove. There are two channels, one on the ventral and one on the dorsal surfaces of the prosthesis. Pilot studies demonstrated that a good cement ®lling of the circular groove and the adjacent spongy bone could be achieved when the cement which was injected into the ventral channel came out of the dorsal channel. Proximal of the circular groove, the surface of the stainless steel prosthesis has a sand-blasted rough structure to promote better bone bonding. The prosthesis head is made of Al2 O3 -ceramic. 2.3. Implantation Fig. 1. Photograph of the partially cemented Option 3000 prosthesis with a groove for the cemented area in the proximal region of the implant.
for cemented implants [4,7,9]. Partially cemented stems therefore have to show that they can provide a sucient initial stability. The aim of the study was to examine whether the new partially cemented stem would be as stable as a fully cemented conventional stem. 2. Methods 2.1. Specimens Six paired human cadaveric femora were used in this study. After the autopsy, they were stored at )24°C in double-sealed plastic bags prior to implantation of the prosthesis. The average age of the donors was 68 yrs, ranging from 49 to 78 yrs. Femora with bone-related diseases were excluded from the study. X-rays in two planes were performed to select the right implant size. The length of the entire specimen was adjusted to 370 mm by shortening the distal femoral shaft to the appropriate length, and the distal part moulded in a
One femur from each of the six pairs to be implanted with a fully cemented implant was chosen at random. The contralateral femur was implanted with a partially cemented prosthesis (Fig. 2). Each stem was implanted by the same orthopaedic surgeon using the standard surgical technique and normal bone cement (Sul®x, Sulzer, Baar, Switzerland). The Option 3000 prosthesis was implanted like a cementless prosthesis using special guided rasps to achieve a good proximal press-®t ®xation. After press-®tting, the cement was injected to ®ll the proximal groove of the prosthesis and the adjacent spongy bone as described above. X-rays performed after implantation showed that the cement was fully surrounding the Weber shaft whereas the cement in the Option 3000 prosthesis was only located at the level of the prosthesis groove in the proximal region of the femur (Fig. 2). 2.4. Micromotion measurement Micromotion of the prosthesis stem relative to the outer cortex of the femur was measured using 4 LVDT (HBM W1, Hottinger, Darmstadt, Germany) with a resolution of 1 lm and two LVDT (Typ GTL222, TESA Brown and Sharp, Renens, Switzerland) with a resolution of 0.5 lm. As a reference point for the locations of
752
L. Claes et al. / Clinical Biomechanics 15 (2000) 750±755
Fig. 2. Anteroposterior contact radiograph of the fully cemented Weber Shaft prosthesis (left side) and a partially cemented Option 3000 prosthesis (right side) implanted in a pair of cadaveric bones.
the transducers one point at the corner of the medial surface of the minor trochanter and the resection plane was marked. Two orthogonal transducers were located at the level 31 mm below the minor trochanter at the lateral and the ventral surfaces of the femur (S3, S5, c, Fig. 3). Two other orthogonal transducers were placed 50 mm below (S4, S6, d, Fig. 3). The transducers were held by custom made devices which were ®xed to the femur by cable binder (Fig. 4). Holes of 4 mm diameter were drilled in longitudinal axes of the transducer, perpendicular to the bone surface, up to the interface between bone and implant surface to allow the contact between the tip of the transducer and the shaft surface. Springs included in the transducer guaranteed a contact between transducer and implant surface during all loading conditions. An other transducer was ®xed at the greater trochanter by a 4.5 mm SchanzÕ screw which measured the movement of the prosthesis in the longitudinal axis of the femur shaft (S1, b 34 mm, Fig. 3). To determine the rotation of the
Fig. 3. Location of the transducers relative to a point at the resection plane on the medial side of the bone.
prosthesis around its longitudinal axis another transducer was ®xed at the ventral part of the greater trochanter, measuring the relative movement between the trochanter and the medial/proximal surface of the prostheses (S2, e 10 mm, Fig. 3). Length a (a 30 mm) was the distance between the longitudinal axis of the distal shaft and the ventral/medial reference point S2 (Fig. 3). 2.5. Loading procedure The specimens were ®xed in a specially designed rig of a material testing machine (Zwick Typ 5040, Einsingen, Germany) for loading in craniocaudal direction. The femora were tilted 8° lateral in the fronal plane and 6° dorsal in the sagittal plane to simulate a single leg stance loading [10] and create bending and torsional moments in all planes (Fig. 4). Three load cycles with a frequency of 0.5 Hz were applied, starting with a load of 200 N. Then the load was increased in steps of 200 N up to a load of 1600 N. The measured displacement at all transdurcers was taken at the maximum of the signal for
L. Claes et al. / Clinical Biomechanics 15 (2000) 750±755
753
Table 1 Micromotion at the prosthesis±bone interface (lm) (means (SD))
Fig. 4. Illustration of the testing setup showing the moulded and instrumented femur ®xed in the testing machine.
the third loading cycle at each load level. Only the data of the last load cycle at 1600 N were used for the statistical analysis. During testing, specimens were kept wet by saline solution to avoid drying of the bone. From the measured translations, the rotation of the prostheses relative to the bone was calculated using trigonometric rules. The data were tested by a Wilcoxon signed rank test to investigate dierences between the two types of prostheses. 3. Results There was a small irreversible deformation of a few microns (0±4) between the ®rst and the second loading cycles of each loading level. For statistical analysis, only the signal of the third loading cycle of the highest load (1600 N) was used, which did not show dierences to the second loading cycle. Therefore the elastic, reversible relative movement between prosthesis and bone was determined. The measured translation showed signi®cant dierences between the two types of prostheses
Transducer
Weber shaft
Option 3000
P-value
S1 S2 S3 S4 S5 S6
23.0 224.8 14.2 6.3 8.2 7.1
4.5 51.5 2.3 10.5 7.8 17.1
0.043 0.043 0.043 0.498 0.500 0.043
(13.5) (163.8) (10.4) (2.3) (6.4) (3.1)
(2.5) (26.9) (2.3) (8.8) (9.4) (7.9)
(Table 1). In general, the micromotions measured in the proximal region (S1, S2, S3, Fig. 3) were always significantly lower (P 0:04) for the Option 3000 than for the Weber Shaft prosthesis. For the distal region the micromotion measured (S4, S6, Fig. 3, Table 1) was small, but higher for the Option 3000 prosthesis than for the Weber Shaft prosthesis. The anteroposterior translation (S6) was signi®cantly (P < 0:04) dierent. The craniocaudal motion (S1) was about 4-times higher for the Weber Shaft than for the Option 3000 (Table 1). About the same dierence was found for the position S2 (Table 1) at the proximal medial aspect of the prosthesis (S2) which describes a combined anterior± posterior translation and a rotation around the longitudinal axis of the prosthesis. Assuming the axis of rotation is in the longitudinal axes of the prosthesis, a mean rotation of the Weber Shaft of 0.61° (S.D. 0.42°) can be calculated whereas the Option 3000 prostheses showed only 0.12° (S.D. 0.04°) of rotation. The tilting of the prosthesis in the frontal plane and sagittal plane was very small for both types of prosthesis and always below 0.02°. 4. Discussion The new partially cemented prosthesis stem showed an initial stability which was comparable to or even better than that of a fully cemented prosthesis. The measurement at various locations around and along the stem axis demonstrated the signi®cantly better proximal ®xation of the new prosthesis against anterior±posterior and medial±lateral translations as well as a signi®cantly smaller rotation around the longitudinal axis. At the non-cemented distal stem area the partially cemented prosthesis showed very small, but larger interface motions than the fully cemented stem. The largest dynamic translation measured for both types of stems in the interface was 43 lm in axial direction and seems to be uncritical with regard to bone resorption. Results from in vivo studies indicate that a critical interface motion during dynamic loading is probably above 100 lm [11,12]. In postmortem studies, it was shown that femoral stems which had been
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L. Claes et al. / Clinical Biomechanics 15 (2000) 750±755
clinically successful for many years had a micromotion of 40±70 lm [13,14]. The larger translation measured at the proximal ventro-medial surface of the Weber Shaft (S2, Table 1) does not describe an interface translation but a combined movement of posterior translation and rotation around the longitudinal stem axis outside the bone-implant interface and with an oset to the axis of rotation. Since neither the bone nor the prosthesis are rigid a small part of the relative movement between the bone and the prosthesis might have resulted from the deformation of the materials itself and not the interface. However, because the stiness of the bone and the prosthesis was very similar for both groups the dierences between groups were mainly caused by the different interface motion in the proximal ®xation. The comparison with other studies is limited because of various conditions in load application, bone density, measuring techniques, measurement locations, types of prostheses and number of load cycles applied. However, the dynamic interface motion measured in this study seems to be relatively small when compared to similar other studies [3,4]. The loading conditions used in vitro depend on the type of in vivo loading which should be simulated. There are numerous possibilities of loading conditions such as single leg stance, various phases of gait or stair climbing. In our opinion, a load case should be used which has a sucient hip load of 2±3 times the body weight [10] and should apply bending moments in both planes (sagittal and frontal) as well as a torsional moment. In this study, a hip joint force of 2.2 times the body weight was used. Because of the tilted femur in both planes, bending moments and torsional moments were produced which were similar to single leg stance loading. However, because of the anatomical variability (i.e., antecurvature), even standardised conditions for the ®xation of the bones in the materials testing machine will lead to different loading conditions of individual bones along the femur axis. To minimise the in¯uence of anatomical variations we always used pairs of bones for the comparison between the two types of prothesis stems and used only the dierence between the measured translations of one pair of bones for statistical analysis. To our knowledge, there is only one other study comparing micromotions of fully and proximally cemented femoral stems [5]. In contrast to the partially cemented stem (Option 3000) in this study, the proximal cemented stem in the other study (Bridge Hip) was cementlessly ®xed by press-®t in the distal stem region. The cemented area for the Bridge Hip was the whole proximally region, whereas in the Option 3000 stem only a part of the proximal region was cemented. However, the dynamic micromotion under similar loading conditions (1500 N) was comparable (12±43 lm) in both
studies. Both studies with proximal cemented stems showed dynamic micromotions similar to the results for fully cemented stems [7,13]. Besides the dynamic elastic movement at the interface between bone and implant, the irreversible deformation accumulated during a large number of loading cycles could be measured. In our investigation, this irreversible deformation was very small, 6 4 lm during 24 load cycles (8 loading levels times 3 cycles) and was not presented in the results. However, the irreversible deformation measured in vitro cannot be compared to the clinically visual subsidence of prosthesis. The testing conditions for in vitro long-term measurements are not representative of the conditions in patients, and in vivo biological processes such as bone remodelling and resorption play a more important role. Subsidence of some degree occurs in clinical cases and does not lead to prostheses loosening [15,16]. The biomechanically caused bone resorption, however, seems to be more in¯uenced by the dynamic interface motion [3,11]. The high initial stability of the new Option 3000 stem shown by the small dynamic interface motions justi®ed a clinical application. In the last 3 yrs, more than 100 Option 3000 stems implanted in patients with very good results were found sofar.
Acknowledgements We would like to thank Dr. Buschor, St. Gallen for the implantation of the stems and the Mathys company for providing the implants.
References [1] Engh CA, Mc TF, Govern JD, Bobyn WH. A quantitative evaluation of periprosthetic bone-remodeling after cementless total hip. J Bone Joint Surg 1992;74-A:1009±20. [2] Turner TM, Sumner DR, Urban RM, Igloria R, Galante JO. Maintenance of proximal cortical bone with use of a less sti femoral component in hemiarthroplasty of the hip without cement. J Bone Joint Surg (Am) 1997;79:1381±90. [3] B uhler DW, Berlemann U, Lippuner K, Jaeger P, Nolte LP. Three-dimensional primary stability of cementless femoral stems. Clin Biomech 1997;12:75±86. [4] Schneider E, Kinast C, Eulenberger J, Wyder D, Eskilsson G, Perren SM. A comparative study of the initial stability of cementless hip prostheses. Chir Orthop Rel Res 1989;248:200±9. [5] Bachus KN, Bloebaum RD, Jones RE. Comparative micromotion of fully and proximally cemented femoral stems. Clin Orthop Rel Res 1999;366:248±57. [6] Whiteside LA, McCarthy DS, White MS. Rotational stability on noncemented total hip femoral components. Am J Orthop 1996:276±80. [7] Burke DW, OÕConnor DO, Zalenski EB. Micromotion of cemented and uncemented femoral components. J Bone Joint Surg 1991;73-B:33±7.
L. Claes et al. / Clinical Biomechanics 15 (2000) 750±755 [8] Hua J, Walker PS. Relative motion of hip stems under load. An in vitro study of symmetrical, asymmetrical, and custom asymmentrical designs. J Bone Joint Surg 1994;76-A:95±103. [9] Charnley J, Kettlewell J. The elimination of slip between prosthesis and femur. J Bone Joint Surg 1965;47-B:56±60. [10] Bergmann G, Graichen F, Rohlmann A. Hip joint loading during walking and running, measured in two patients. J Biomech 1993;26:969±90. [11] Sùballe K, Hansen ES, Rasmussen H. Tissue ingrowth into titanium and hydroxyapatite-coated implants during stable and unstable mechanical conditions. J Orthop Res 1991;10:285±99. [12] Pilliar RM, Lee JM, Maniatopoulus C. Observations on the eect of movement on bone ingrowth into porous-surfaced implants. Clin Orthop 1986;208:108±13.
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[13] Maloney WJ, Burke DW, OÕConnor DO, Zalenski EB, Bragdon C, Harris WH. Biomechanical and histologic investigation of cemented total hip arthroplasties. A study of autopsy-retrieved femurs after in vivo cycling. Clin Orthop 1989;249:129±40. [14] Wilke HJ, Seitz RS, Bombelli M, Claes L, D urselen L. Biomechanical and histomorphological investigations on a isoelastic prosthesis after eight months implantation. J Mater Sci Mat Med 1995;5:384±6. [15] Blaha JD, Spotorno L, Romagnoli S. CLS press-®t total hip arthroplasty. Tech Orthop 1991;6:80±6. [16] Mj oberg B, Brismar J, Hansson LI, et al. De®nition of endoprosthetic loosening ± Comparison of arthrography, scintigraphy and roentgen stereo photogrammetry in prosthetic hips. Acta Orthop Scand 1985;56:469±73.
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Initial stability of fully and partially cemented femoral stems Lutz Claes a,*, Stefan Fiedler a, Michael Ohnmacht a, Georg N. Duda b b
a Institut f ur Unfallchirurgische Forschung und Biomechanik, Universit at Ulm, Helmholtzstraûe 14, 89081 Ulm, Germany Virchow-Klinikum, Abteilung Unfallchirurgie und Wiederherstellungschirurgie, Humboldt Universit at zu Berlin, 13353 Berlin, Germany
Received 26 April 2000; accepted 13 July 2000
Abstract Objective. To test the initial stability of a newly designed partially cemented femoral stem in comparison with a fully cemented conventional stem. Design. An in vitro study to determine the interface motion between femoral stem and bone as a response to loading. Background. The aim of the new prosthesis design is a proximal load transfer by a de®ned partial cement ®xation in the proximal femur region and a slim prosthesis stem in the distal region. Before a clinical study can be started, the new stem has to show an initial stability comparable to that of fully cemented prostheses. Method. Six paired fresh cadaveric femora were used for the testing of the new partially cemented stem (Option 3000, Mathys Orthopaedics, Bettlach, Switzerland) and a fully cemented stem (Weber Shaft, AlloPro, Baar, Swizerland). Under cyclic loading up to 1600 N hip joint forces, the interface motion between implants and bone was measured at six locations. Results. Both stems showed uncritical interface motions below 43 lm. However, the Option 3000 stem exhibited signi®cantly smaller motions in the proximal region and slightly larger movements in the distal regions than the Weber prosthesis. Conclusions. The new type of partially cemented stem provided a comparable initial stability to the fully cemented Weber prosthesis. Relevance The high initial stability of the Option 3000 stem justi®ed the clinical use of the new implant. More than 100 implantations in the last three years, with very good preliminary clinical results, support the preclinical ®ndings. Ó 2000 Published by Elsevier Science Ltd. Keywords: Total hip replacement; Initial stability; Prosthesis design; Mechanical testing
1. Introduction Under physiological conditions, the loads acting on the femoral head are transferred to the femoral shaft by the trabecular system of the proximal femur [1,2]. Distally anchoraged femoral stems often demonstrate a proximal bone loss. One of the possible reasons for this eect is due to the unphysiological load transfer and corresponding stress shielding of the proximal femur region [1,2]. Recently, a new femoral stem was developed which is mainly ®xed in the proximal part of the femur by press-®t and partial
*
Corresponding author. E-mail address: [email protected] (L. Claes).
0268-0033/00/$ - see front matter Ó 2000 Published by Elsevier Science Ltd. PII: S 0 2 6 8 - 0 0 3 3 ( 0 0 ) 0 0 0 4 4 - 9
cementing (Option 3000, Mathys Orthopaedics, Bettlach, Switzerland, Patent PCT CH 95/00260). The distal uncemented part of this prosthesis is very slim, ®lls the intramedullary canal only partly, and avoids axial load transfer (Fig. 1). However, the question remained open whether this partially cemented stem would provide enough initial stability. It is well accepted that initial stability and lack of motion between bone and implant is one of the most important prerequisites for the long-term success of prostheses [3±5]. Therefore, one of the ®rst steps in the preclinical testing of new prostheses should be the measurement of the interface motion under sucient loading conditions. Such investigations have been carried out for cemented and cementless femoral stems by several groups [3±8]. It could be shown that the risk for interface motion and rotation was higher for cementless stems than
L. Claes et al. / Clinical Biomechanics 15 (2000) 750±755
751
metal cup using methylmethacrylate (Technovit 4040, Merck, Darmstadt, Germany). 2.2. Implants The new type of prosthesis (Option 3000, Mathys Orthopaedics, Bettlach, Switzerland) is designed to be ®xed only in the proximal region of the femur. It has an oval collarless cross-sectional design at the most proximal region and a very slim distal stem (Fig. 1). The small diameter of the distal stem avoids a distal ®tting and has a polished surface. In the mid-proximal region of the prosthesis, a circular groove allows the locally de®ned application of bone cement when the prosthesis is already press-®tted in the femur. The bone cement can be admitted by a longitudinal channel running from the proximal end of the prosthesis into the circular groove. There are two channels, one on the ventral and one on the dorsal surfaces of the prosthesis. Pilot studies demonstrated that a good cement ®lling of the circular groove and the adjacent spongy bone could be achieved when the cement which was injected into the ventral channel came out of the dorsal channel. Proximal of the circular groove, the surface of the stainless steel prosthesis has a sand-blasted rough structure to promote better bone bonding. The prosthesis head is made of Al2 O3 -ceramic. 2.3. Implantation Fig. 1. Photograph of the partially cemented Option 3000 prosthesis with a groove for the cemented area in the proximal region of the implant.
for cemented implants [4,7,9]. Partially cemented stems therefore have to show that they can provide a sucient initial stability. The aim of the study was to examine whether the new partially cemented stem would be as stable as a fully cemented conventional stem. 2. Methods 2.1. Specimens Six paired human cadaveric femora were used in this study. After the autopsy, they were stored at )24°C in double-sealed plastic bags prior to implantation of the prosthesis. The average age of the donors was 68 yrs, ranging from 49 to 78 yrs. Femora with bone-related diseases were excluded from the study. X-rays in two planes were performed to select the right implant size. The length of the entire specimen was adjusted to 370 mm by shortening the distal femoral shaft to the appropriate length, and the distal part moulded in a
One femur from each of the six pairs to be implanted with a fully cemented implant was chosen at random. The contralateral femur was implanted with a partially cemented prosthesis (Fig. 2). Each stem was implanted by the same orthopaedic surgeon using the standard surgical technique and normal bone cement (Sul®x, Sulzer, Baar, Switzerland). The Option 3000 prosthesis was implanted like a cementless prosthesis using special guided rasps to achieve a good proximal press-®t ®xation. After press-®tting, the cement was injected to ®ll the proximal groove of the prosthesis and the adjacent spongy bone as described above. X-rays performed after implantation showed that the cement was fully surrounding the Weber shaft whereas the cement in the Option 3000 prosthesis was only located at the level of the prosthesis groove in the proximal region of the femur (Fig. 2). 2.4. Micromotion measurement Micromotion of the prosthesis stem relative to the outer cortex of the femur was measured using 4 LVDT (HBM W1, Hottinger, Darmstadt, Germany) with a resolution of 1 lm and two LVDT (Typ GTL222, TESA Brown and Sharp, Renens, Switzerland) with a resolution of 0.5 lm. As a reference point for the locations of
752
L. Claes et al. / Clinical Biomechanics 15 (2000) 750±755
Fig. 2. Anteroposterior contact radiograph of the fully cemented Weber Shaft prosthesis (left side) and a partially cemented Option 3000 prosthesis (right side) implanted in a pair of cadaveric bones.
the transducers one point at the corner of the medial surface of the minor trochanter and the resection plane was marked. Two orthogonal transducers were located at the level 31 mm below the minor trochanter at the lateral and the ventral surfaces of the femur (S3, S5, c, Fig. 3). Two other orthogonal transducers were placed 50 mm below (S4, S6, d, Fig. 3). The transducers were held by custom made devices which were ®xed to the femur by cable binder (Fig. 4). Holes of 4 mm diameter were drilled in longitudinal axes of the transducer, perpendicular to the bone surface, up to the interface between bone and implant surface to allow the contact between the tip of the transducer and the shaft surface. Springs included in the transducer guaranteed a contact between transducer and implant surface during all loading conditions. An other transducer was ®xed at the greater trochanter by a 4.5 mm SchanzÕ screw which measured the movement of the prosthesis in the longitudinal axis of the femur shaft (S1, b 34 mm, Fig. 3). To determine the rotation of the
Fig. 3. Location of the transducers relative to a point at the resection plane on the medial side of the bone.
prosthesis around its longitudinal axis another transducer was ®xed at the ventral part of the greater trochanter, measuring the relative movement between the trochanter and the medial/proximal surface of the prostheses (S2, e 10 mm, Fig. 3). Length a (a 30 mm) was the distance between the longitudinal axis of the distal shaft and the ventral/medial reference point S2 (Fig. 3). 2.5. Loading procedure The specimens were ®xed in a specially designed rig of a material testing machine (Zwick Typ 5040, Einsingen, Germany) for loading in craniocaudal direction. The femora were tilted 8° lateral in the fronal plane and 6° dorsal in the sagittal plane to simulate a single leg stance loading [10] and create bending and torsional moments in all planes (Fig. 4). Three load cycles with a frequency of 0.5 Hz were applied, starting with a load of 200 N. Then the load was increased in steps of 200 N up to a load of 1600 N. The measured displacement at all transdurcers was taken at the maximum of the signal for
L. Claes et al. / Clinical Biomechanics 15 (2000) 750±755
753
Table 1 Micromotion at the prosthesis±bone interface (lm) (means (SD))
Fig. 4. Illustration of the testing setup showing the moulded and instrumented femur ®xed in the testing machine.
the third loading cycle at each load level. Only the data of the last load cycle at 1600 N were used for the statistical analysis. During testing, specimens were kept wet by saline solution to avoid drying of the bone. From the measured translations, the rotation of the prostheses relative to the bone was calculated using trigonometric rules. The data were tested by a Wilcoxon signed rank test to investigate dierences between the two types of prostheses. 3. Results There was a small irreversible deformation of a few microns (0±4) between the ®rst and the second loading cycles of each loading level. For statistical analysis, only the signal of the third loading cycle of the highest load (1600 N) was used, which did not show dierences to the second loading cycle. Therefore the elastic, reversible relative movement between prosthesis and bone was determined. The measured translation showed signi®cant dierences between the two types of prostheses
Transducer
Weber shaft
Option 3000
P-value
S1 S2 S3 S4 S5 S6
23.0 224.8 14.2 6.3 8.2 7.1
4.5 51.5 2.3 10.5 7.8 17.1
0.043 0.043 0.043 0.498 0.500 0.043
(13.5) (163.8) (10.4) (2.3) (6.4) (3.1)
(2.5) (26.9) (2.3) (8.8) (9.4) (7.9)
(Table 1). In general, the micromotions measured in the proximal region (S1, S2, S3, Fig. 3) were always significantly lower (P 0:04) for the Option 3000 than for the Weber Shaft prosthesis. For the distal region the micromotion measured (S4, S6, Fig. 3, Table 1) was small, but higher for the Option 3000 prosthesis than for the Weber Shaft prosthesis. The anteroposterior translation (S6) was signi®cantly (P < 0:04) dierent. The craniocaudal motion (S1) was about 4-times higher for the Weber Shaft than for the Option 3000 (Table 1). About the same dierence was found for the position S2 (Table 1) at the proximal medial aspect of the prosthesis (S2) which describes a combined anterior± posterior translation and a rotation around the longitudinal axis of the prosthesis. Assuming the axis of rotation is in the longitudinal axes of the prosthesis, a mean rotation of the Weber Shaft of 0.61° (S.D. 0.42°) can be calculated whereas the Option 3000 prostheses showed only 0.12° (S.D. 0.04°) of rotation. The tilting of the prosthesis in the frontal plane and sagittal plane was very small for both types of prosthesis and always below 0.02°. 4. Discussion The new partially cemented prosthesis stem showed an initial stability which was comparable to or even better than that of a fully cemented prosthesis. The measurement at various locations around and along the stem axis demonstrated the signi®cantly better proximal ®xation of the new prosthesis against anterior±posterior and medial±lateral translations as well as a signi®cantly smaller rotation around the longitudinal axis. At the non-cemented distal stem area the partially cemented prosthesis showed very small, but larger interface motions than the fully cemented stem. The largest dynamic translation measured for both types of stems in the interface was 43 lm in axial direction and seems to be uncritical with regard to bone resorption. Results from in vivo studies indicate that a critical interface motion during dynamic loading is probably above 100 lm [11,12]. In postmortem studies, it was shown that femoral stems which had been
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L. Claes et al. / Clinical Biomechanics 15 (2000) 750±755
clinically successful for many years had a micromotion of 40±70 lm [13,14]. The larger translation measured at the proximal ventro-medial surface of the Weber Shaft (S2, Table 1) does not describe an interface translation but a combined movement of posterior translation and rotation around the longitudinal stem axis outside the bone-implant interface and with an oset to the axis of rotation. Since neither the bone nor the prosthesis are rigid a small part of the relative movement between the bone and the prosthesis might have resulted from the deformation of the materials itself and not the interface. However, because the stiness of the bone and the prosthesis was very similar for both groups the dierences between groups were mainly caused by the different interface motion in the proximal ®xation. The comparison with other studies is limited because of various conditions in load application, bone density, measuring techniques, measurement locations, types of prostheses and number of load cycles applied. However, the dynamic interface motion measured in this study seems to be relatively small when compared to similar other studies [3,4]. The loading conditions used in vitro depend on the type of in vivo loading which should be simulated. There are numerous possibilities of loading conditions such as single leg stance, various phases of gait or stair climbing. In our opinion, a load case should be used which has a sucient hip load of 2±3 times the body weight [10] and should apply bending moments in both planes (sagittal and frontal) as well as a torsional moment. In this study, a hip joint force of 2.2 times the body weight was used. Because of the tilted femur in both planes, bending moments and torsional moments were produced which were similar to single leg stance loading. However, because of the anatomical variability (i.e., antecurvature), even standardised conditions for the ®xation of the bones in the materials testing machine will lead to different loading conditions of individual bones along the femur axis. To minimise the in¯uence of anatomical variations we always used pairs of bones for the comparison between the two types of prothesis stems and used only the dierence between the measured translations of one pair of bones for statistical analysis. To our knowledge, there is only one other study comparing micromotions of fully and proximally cemented femoral stems [5]. In contrast to the partially cemented stem (Option 3000) in this study, the proximal cemented stem in the other study (Bridge Hip) was cementlessly ®xed by press-®t in the distal stem region. The cemented area for the Bridge Hip was the whole proximally region, whereas in the Option 3000 stem only a part of the proximal region was cemented. However, the dynamic micromotion under similar loading conditions (1500 N) was comparable (12±43 lm) in both
studies. Both studies with proximal cemented stems showed dynamic micromotions similar to the results for fully cemented stems [7,13]. Besides the dynamic elastic movement at the interface between bone and implant, the irreversible deformation accumulated during a large number of loading cycles could be measured. In our investigation, this irreversible deformation was very small, 6 4 lm during 24 load cycles (8 loading levels times 3 cycles) and was not presented in the results. However, the irreversible deformation measured in vitro cannot be compared to the clinically visual subsidence of prosthesis. The testing conditions for in vitro long-term measurements are not representative of the conditions in patients, and in vivo biological processes such as bone remodelling and resorption play a more important role. Subsidence of some degree occurs in clinical cases and does not lead to prostheses loosening [15,16]. The biomechanically caused bone resorption, however, seems to be more in¯uenced by the dynamic interface motion [3,11]. The high initial stability of the new Option 3000 stem shown by the small dynamic interface motions justi®ed a clinical application. In the last 3 yrs, more than 100 Option 3000 stems implanted in patients with very good results were found sofar.
Acknowledgements We would like to thank Dr. Buschor, St. Gallen for the implantation of the stems and the Mathys company for providing the implants.
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