Total hip arthroplasty (THA) is one of the most successful procedures in modern orthopaedic surgery to provide pain relief and restoration of limb function, but instability is a troublesome complication. Dislocation occurred in 0.7% to greater than 10% of patients in series including primary and revision procedures,9,16-18,31-34,36,37,41 and there is no evidence the prevalence of dislocation has declined despite design and technique improvements.9 Many dislocations can be treated successfully with closed reduction and bracing, but approximately ⅓ of patients may experience recurrence.25,41 Revision for recurrent dislocation is a challenging problem as the redislocation rate after reoperation for hip instability ranges from 20% to 40%.1,20,24,41
Numerous surgical options have been described to treat recurrent dislocation. Implant designs have been altered to improve range of motion (ROM) before impingement and stability. Modular components, elevated rim liners or socket wall additions, conversion to bipolar hemiarthroplasty, and constrained acetabular components have been used with mixed clinical results.9 These limitations led to the development of tripolar hip prostheses.7,8,23,29,35 These devices introduce an additional bearing surface, using a bipolar component with a large inside-diameter acetabular cup. Successful clinical results have been reported for treating recurrent dislocation7,8,21-23,29,35,39,40 and preventing dislocation.3,19,21 Success rates for stability range from 95% to 100%.3,7,8,19,29 These encouraging results have led to the increasing use of tripolar devices, particularly in Europe.
The ROM of hip implants is closely related to the prosthetic head size and head to neck ratio: increasing the head to neck ratio resulted in a greater ROM before impingement.2,4,6,11,14 By including an additional bearing, the tripolar design has a more effective head design. Therefore, we expected tripolar implants would provide greater ROM before impingement than conventional implants, and subsequently improve stability in patients at risk for dislocation.
We asked whether the theoretical increase in ROM of such implants could be experimentally confirmed.
MATERIALS AND METHODS
We mounted a tripolar hip implant to an automated hip simulator to assess the in vitro ROM to impingement. We tested conventional implants with various head sizes and also compared ROM.
We tested a conventional acetabular component (Equateur®, Amplitude, Neyron, France) and a same-size tripolar cup (Saturne®, Amplitude). The cups were tested with a 22.2-mm and a 28-mm femoral head diameter (neutral neck length) using the same femoral stem (Generic®, Amplitude) with a constant neck length, 10-mm neck diameter, 137°-neck-shaft angle, 10/12 Morse taper, and 42-mm offset. Simulations were performed using a Size 4 stem and 54-mm outer diameter conventional and tripolar cups (Fig 1). The tripolar cup currently is not available in the United States as it is not FDA approved. The tripolar cup has a stainless steel outer shell with a highly polished inner surface. The outer shell is anatomically designed and has superior and posterior lips that are greater than hemispheric, and anterior and inferior cut-outs that are smaller than hemispheric. This shell articulates with a mobile intermediate ultrahigh molecular weight polyethylene (UHMWPE) component (Fig 2). To constrain the femoral head, the UHMWPE component envelops more than 50% of the femoral head and its opening diameter is smaller than the femoral head. When reduced, the head is captured in the polyethylene.
We designed an automated hip simulator to assess the in vitro ROM to impingement of THA implants (Fig 3). A bidirectional arm was mounted on a rotating table which allowed flexion/extension and internal/external rotation. The rotating table allowed us to select a plane for flexion and extension (plane of elevation). These three axes were controlled independently with stepper motors (Servo Systems, Montville, NJ). Acetabular and femoral components were inserted into Sawbones® model human pelves and femurs (ref. 1301 and 1130, Pacific Research Laboratories, Vashon, WA) after an appropriate femoral neck osteotomy. The acetabular components were positioned in 45° abduction (lateral opening) and 20° anteversion.30 The stem was positioned in 20° anteversion and in neutral varus/valgus alignment. We determined the degree of femoral stem version according to the bicondylar axis using the posterior surfaces of the condyles as 0° version.2 We used modeling clay to stabilize the components until confirming correct positioning with respect to the pelvis and femur using a three-dimensional (3-D) digitizer (Faro Arm Gold series CMM, Faro Technologies Inc, Lake Mary, FL). The components then were locked in place with a collar of acrylic bone cement, and their correct positioning was confirmed using the digitizer. To detect impingement, the rim of the acetabular liner and cup, the bony acetabular rim, and the proximal femur were coated with conductive copper foil (3M, Minneapolis, MN) connected to an external power source.6,38 The femoral stem and acetabular cups were conductive. Impingement was detected when electrical continuity was recorded. The pelvis and femur were mounted in the apparatus after removing the distal portion of the femur so the model could fit into the device. The pelvis was mounted and held in place in an orientation simulating the pelvis during standing, with the anterior superior iliac spines and the pubis in a vertical plane.2 Before each test we used a plumb line and a custom fixture (a plastic ball with an appropriate diameter to simulate the prosthetic head and two orthogonal rigid Kirschner wires passing through its center) to adjust the pelvic position and orientation (Fig 4). The femur then was mounted to the bidirectional arm. A spring mounted to the tip of the femur applied the necessary compressive force across the joint to maintain the THA in consistent alignment while measuring ROM. The sequence of motion was controlled using LabVIEW® 7 Express (National Instruments, Austin, TX). From the neutral position (flexion 0°, adduction 0°, and rotation 0°), we performed internal and external rotations until impingement. Next, the motor (back to its initial position) elevated the femur for flexion and extension until impingement, and rotated back to initial position. Impingement was reached at different planes of elevation while simulating flexion/extension at various adduction angles. Pure adduction and pure abduction were simulated at the plane of 90° elevation. Angles were recorded after impingement. We assessed ROM to impingement for the conventional and the tripolar implants in flexion, extension, abduction, adduction, internal rotation, and external rotation. To investigate stability, we also measured ROM at clinically relevant positions including: (1) the maximum internal rotation before impingement (with the hip flexed at 90°) at 20° adduction increments from 0° to 40°; and (2) the maximum extension before impingement at 10° external rotation increments from 0° to 50° at neutral position (0° flexion).
We repeated measurements three times to determine the average. We also recorded the standard deviations. The results of the various implants tested were compared using analysis of variance (ANOVA) (p = 0.05) and Bonferroni tests.
The prosthetic head size (22.2 mm or 28 mm) of the tripolar design did not affect global ROM. However, there was difference in global ROM in any direction between the 22.2-mm and 28-mm tripolar implants (Table 1). Compared with conventional implants, the tripolar implant (22.2- or 28-mm prosthetic head) provided greater (p < 0.05) arcs of motion in the sagittal, coronal, and transverse planes (Fig 5). The tripolar implant increased arcs of motion in all directions that were not limited by bony impingement (Fig 6). Compared with the 22.2-mm conventional implant, ROM increased 30.5° (30%) in flexion, 18.8° (44%) in extension, 12.1° (20%) in abduction, 15.4° (37%) in adduction, and 22.4° (38%) in external rotation. Compared with the 28-mm conventional implant, ROM increased 8.2° (7%) in flexion, 13.6° (28%) in extension, 8.4° (17%) in adduction, and 15.1° (23%) in external rotation.
Compared with the conventional implants, the tripolar implant provided a larger (p < 0.05) ROM in flexion during adduction (Fig 7). For both implants, progressive increments of hip adduction resulted in a progressive decrease in maximal flexion. The tripolar implant had a larger maximal internal rotation in adduction (from 0° to 40°) compared with the conventional implants. Compared with the 22.2-mm conventional implant, the internal rotation increase reached 45.2° at 40° adduction, whereas it was 27.5° compared with the 28-mm conventional implant (Fig 8).
The tripolar implant also provided greater ROM in extension during external rotation (0°-50°) compared with the conventional implants. At 30° external rotation, extension increased 7.8° compared with the 22.2-mm conventional implant, and 5.7° compared with the 28-mm conventional implant. Increasing the external rotation angle resulted in a progressive decrease in maximal extension for all implants (Fig 9).
Range of motion simulations for THA implants are used to assess the effects of component position and implant shape on the femoral and acetabular sides. Comparative analyses indicate substantial differences between implant designs.2,5,6,11,13,26,27 Range of motion studies have been reported for conventional implants2,5,11,27 and for constrained devices.10 Effects of prosthetic head diameter, neck shape (length and diameter), modular components, and elevated rim acetabular components have been studied.2,4,5,10-13,26,28 However, ranges of motion provided by nonconstrained tripolar devices have not been reported. Different methods have been described to assess ROM including laboratory evaluations,2,6,11,12,26-28 computational models,4,5,13 and mathematical formulas.42
Our study has several limitations. The use of electrical impingement detection theoretically improved the accuracy of the tests and avoided bias from manual detection, but dislocation was not measured as it requires the simulation of compressive forces across the hip. The experimental apparatus could not fully simulate the complex and changing force vectors of the hip musculature during ROM. The experimental test was performed with bony structures, but effects of muscles and soft tissues were not included. Therefore, the extremes of motion measured in certain directions may exceed what occurs in vivo. Once soft tissues have settled around the prosthesis in vivo, fibrosis could be responsible for limiting the motion at the inner and/or outer bearing surface. Such a situation has not been reported for a tripolar device. However, Drinker and Murray15 reported fibrosis around the head/neck area of a bipolar implant which limited motion at the inner bearing surface. Some surgeons advocate even larger diameter femoral heads (eg, 32-mm or larger) for instability. Our study does not address the issue of whether even larger head sizes would have even greater effects. Despite these limitations, the tripolar device (22.2-mm or 28-mm femoral head) provided improved ROM before impingement compared with 22.2-mm and 28-mm conventional implants in any direction not limited by bony impingement.
Absolute quantitative comparison with previous studies is not possible as pelves, implant sizes, and orientations varied. The increased ROM provided by the tripolar implants was limited by bony impingement. There were two major anatomic constraints: the ischial tuberosity and the proximal femur impinged to limit external rotation, and the greater trochanter impinged with the pubic and iliac bones to limit internal rotation and abduction. Prosthetic impingement and ROM are markedly influenced by the configuration of the head to neck area. A large neck diameter decreases prosthetic ROM. Smaller anteroposterior neck widths provide increased ROM.2,5,27,28 In contrast, a larger head diameter increases ROM for a constant neck diameter and shape.2,5,6,11,13,14,27 An increase of the head to neck ratio allows a larger ROM before neck-socket impingement occurs, and can result in a transition from prosthetic to bony impingement.6,11 The 22.2-mm and 28mm tripolar implants had the same global ROM. As the outer diameter of the mobile polyethylene component represented the effective head diameter, using a 22.2-mm or a 28-mm prosthetic femoral head did not influence global ROM. However, the inner-bearing ROM was affected by the prosthetic femoral head (Fig 10). The tripolar device provided a similar effect as conversion to an ultralarge head. Therefore, tripolar devices provide a larger head to neck ratio than conventional implants. The tripolar implant head to neck ratio was 4.7, compared with 2.2 and 2.8 for the 22.2-mm and 28-mm femoral head conventional implants, respectively. The larger the tripolar implant outer cup diameter, the greater the head to neck ratio. With a 54-mm outer cup, using the same head diameter as for conventional implants, the tripolar implant provided an increased head to neck ratio: 114% more than the conventional 22.2-mm design, and 68% more than the 28-mm conventional design. The head to neck ratio for the tripolar implants was greater than ratios reported for other designs. Conventional implants have been reported with maximal head to neck ratios ranging from 2.29 to 3.4.2,11,13 Increasing the effective head diameter with the tripolar device resulted in an increased head to neck ratio and in delayed prosthetic impingement or a transition from prosthetic to bony impingement.
We showed the in vitro benefit of tripolar devices increasing ROM compared with conventional implants. In patients at risk for dislocation, the larger head to neck ratio and outer-shell anatomic design of the tripolar implant provided an increased of arc of motion before impingement. We expect tripolar implants to provide greater stability and reduce the dislocation rates by delaying prosthetic impingement. Such an assumption is reinforced by encouraging clinical results with unconstrained tripolar devices.3,7,8,19,21,23,29,35,40
We thank Geraldine Bernard (Department of Engineering, Mayo Clinic), and Frederic Schultz and Prashanth Prabhakar (Biome-chanics Laboratory, Division of Orthopedic Research, Mayo Clinic) for contributions to this study.
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