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The influence of implant articular thickness and glenohumeral conformity on stability of an all-metal glenoid component
The influence of implant articular thickness and glenohumeral conformity on stability of an all-metal glenoid component Ryan T. Bicknell, MD, MSc, FRCSC,a Allan S. L. Liew, MD, FRCSC,a Matthew R. Danter, MD,c Stuart D. Patterson, MBChB, FRCSC,d Graham J. W. King, MD, MSc, FRCSC,a,b,c,e David G. Chess, MD, FRCSC,a,b,c,e and James A. Johnson, PhD,a,b,c,e London, Ontario, Canada, and Lakeland, FL
The objective of this study was to determine the effect of implant thickness and glenohumeral conformity on fixation of an all-metal glenoid component. A stainless steel glenoid component was designed and implanted in 10 cadaveric scapulae. A testing apparatus capable of producing a loading vector at various angles, magnitudes, and directions was used. The independent variables included 6 directions and 3 angles of joint load, 3 implant thicknesses, and 4 glenohumeral conformities. Implant micromotion relative to bone was measured by use of 4 displacement transducers at the superior, inferior, anterior, and posterior sites. The components displayed a consistent response to loading of ipsilateral compression and contralateral distraction. Stability decreased as the load application angle increased (P ⬍ .05). A decrease in the implant thickness and glenohumeral conformity resulted in increased implant stability (P ⬍ .05). Decreasing implant thickness and glenohumeral conformity reduce the eccentric component of loading and may improve the durability of glenoid implants. (J Shoulder Elbow Surg 2007;16:631-639.)
L oosening of the glenoid component remains a fac-
tor limiting the long-term success of total shoulder arthroplasty.7,16,18,27,36,41,43,44,46 Furthermore, glenoid loosening has been shown to be significantly
From the Departments of aSurgery, bMechanical and Materials Engineering, and eMedical Biophysics, University of Western Ontario, and cHand and Upper Limb Centre, St Joseph’s Health Centre, London, and dCentral Florida Orthopaedic Surgery Associates, Lakeland. Reprint requests: Ryan T. Bicknell, MD, MSc, FRCSC, Hand and Upper Limb Centre, St Joseph’s Health Centre, 268 Grosvenor St, London, Ontario, Canada, N6A 4L6 (E-mail: rtbickne@ yahoo.ca). Copyright © 2007 by Journal of Shoulder and Elbow Surgery Board of Trustees. 1058-2746/2007/$32.00 doi:10.1016/j.jse.2006.10.014
correlated with deteriorating functional results and increasing pain.7,24,39,43 This is thought to be a result of several factors, including inadequate implant-bone fixation and eccentric loading. Implant-bone fixation has been investigated extensively. Various designs have attempted to achieve fixation to bone through the use of anchoring systems on the undersurface of the glenoid component. In vitro studies have been conducted to assess the strength of fixation and the stability or micromotion of various implant designs relative to bone.1,6,8,31,32 Theoretic finite-element studies have also analyzed the state of stress in the implant-bone structure.15,23,32,38 However, these studies have not conclusively defined the optimal methods by which to maximize the stability between the glenoid component and bone. Previous studies have also indicated that the glenoid component is very sensitive to eccentric loading (ie, loading that occurs nonperpendicular to the glenoid surface).6,9,25 Distraction or liftoff of the implant from the bone surface, potentially inhibiting bone ingrowth or producing interface tension, is a concern with eccentric loading. In extreme arm positions or with certain pathologic conditions, subluxation of the humeral head may lead to rim loading of the glenoid.3,14,20,22 It has been shown that advances in bone preparation and cementing techniques can reduce the effects of eccentric loading that act to cause glenoid component distraction and loosening.9,30 Reducing the eccentricity of the applied load (and also rim loading) by using a thinner implant has also been shown to minimize the distraction of the glenoid component.25 In addition, reducing glenohumeral conformity may also reduce eccentric loading by minimizing rim loading.21,22,37 The objectives of this study were to determine the effects of implant articular thickness and glenohumeral conformity on the stability of uncemented allmetal glenoid components, as quantified by the magnitude of micromotion between implant and bone. We hypothesized that decreasing implant articular
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Figure 1 Testing apparatus. A, The loading apparatus consisted of 2 pneumatic actuators located orthogonal to each other, with an adjustable loading stage to permit loading in any direction. Actuators 1 and 2 produced the medial and transverse (superior [S], inferior [I], anterior [A], and posterior [P]) compressive forces, respectively. B, Micromotion measurements were performed with LVDTs. These included an LVDT barrel attached to the bone and an LVDT core attached to the implant.
thickness and decreasing glenohumeral conformity would improve glenoid implant stability. MATERIALS AND METHODS Testing apparatus and instrumentation A pneumatic loading apparatus and measurement system was used, as described in a previous study6 (Figure 1). This consisted of a loading stage fitted with 2 pneumatic actuators (SR-243-D; Bimba Manufacturing, Monee, IL), capable of supplying loads up to 492.5 N at 0.345 MPa of air pressure. The actuators were controlled by 2 virtual instrument device drivers written in LabVIEW software (National Instruments, Austin, TX), and actuator pressure was controlled by Proportional Pressure Controllers (PPC5AAAA-AGCB-BBB-JB; MAC Valves, Wixom, MI). A steel ball was used to model the humeral component and to provide a surface to articulate with the glenoid component.6,9 For this study, we used 3 spheres of different diameters that enabled us to vary glenohumeral conformity. The actuators directly loaded the ball bearing placed on top of the glenoid component (Figure 1). By varying the load from each actuator, the target magnitude and line of action
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of the resultant force could be achieved. The loading stage was adjustable and allowed rotation of the glenoid in the sagittal plane, permitting loading in any direction. This allowed free displacement of the humeral component in the plane of the glenoid component. To minimize nonphysiologic friction, the articular surfaces were coated with bovine serum (HyClone, Logan, UT). To quantify micromotion of the implant with respect to bone, 4 linear variable differential transducers (LVDTs) were used (Figure 1, B). These were chosen for their highly linear response for a given displacement (error of 0.01%), availability in various sizes, cost-effectiveness, and high sensitivity.17 A support ring served as a mounting stage for 4 LVDT barrels (Lucas-Schaevitz 049-XS-B; Durham Instruments, Pickering, Ontario, Canada). The LVDT ring was fastened to the cortical glenoid neck with screws (diameter, 4.2 mm). To permit attachment of the LVDT cores, four 0.9-mm Kirschner wire outriggers were fixed to each glenoid component, two along each of the superoinferior and anteroposterior axes. This permitted measurement at the superior, inferior, anterior, and posterior locations. After the LVDT ring was fixed to the glenoid, the component was implanted and the LVDT cores aligned and placed within their respective barrels. The cores were then fastened to the Kirschner wire outriggers by use of a swivel hinge, located a distance of 5 mm from the glenoid component. Before experimental testing, each of the LVDT signal transducers was calibrated in accordance with technical manuals provided by the manufacturer. Throughout the study, data were collected and converted from LVDT output voltages to LVDT displacements by use of LabVIEW software (National Instruments).
Surgical technique Glenoid preparation followed the published protocol for the 3M Neer II prosthesis (3M, Orthopaedic Products Division, St Paul, MN). The Neer II glenoid reamer (3M, Orthopaedic Products Division) was used to complete surface preparation. To ensure the best possible conformity of the prosthesis to the glenoid surface, this reamer had a radius of curvature of 28 mm, which is equal to the radius of curvature of the component articular undersurface.9 Hand reaming of the lateral surface of the articulation was performed until all cartilage was removed. A hand bur was used to make a slot for each keel through the subchondral bone, and the component was then impacted into the glenoid. The reamed surface of the glenoid was irrigated by use of 0.9% normal saline solution throughout testing, to limit drying of the specimen. The same surgeon conducted all glenoid preparations and component implantations.
Implant design This study used a prototype all-metal stainless steel glenoid component, with a cross-keel, as developed and evaluated in a previous study6 (Figure 2). The articular surface of the component was based on the Neer II component (3M, Orthopaedic Products Division).12,16 This includes an articular surface radius of curvature of 25 mm; an undersurface radius of curvature of 28 mm; coronal and transverse articular surface widths of 32.5 and 24.5 mm, respectively; and an articular surface thickness of 4 mm. The
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Table I Loading protocol: Direction and angle of load application Direction Superior
Inferior Anterior Posterior Superoposterior Inferoposterior
Figure 2 Implant design. The glenoid implant used in this study is shown in the coronal and sagittal planes. In addition to the keel present in the coronal plane in the Neer II implant design, the prototype implant also combined a keel in the transverse plane. A, Articular surface thickness; B, Neer II keel; C, large keel from previous study6; D, articular surface; E, screw holes from previous study6; F, articular undersurface; G, coronal keel; H, transverse keel; I, coronal component width; J, transverse component width.
component utilized a cross-keel design (as opposed to a single-keeled or pegged design), to maximize contact, based on the anatomic dimensions of the glenoid vault in both planes (transverse and coronal), as determined in a previous study.5 In addition to the (standard) keel in the coronal plane, our glenoid component had another keel placed orthogonally in the transverse plane. This provided the potential to counteract superiorly and inferiorly directed joint forces and was shown in a previous study to decrease glenoid component micromotion.6 The medial edge of each keel was tapered to assist with the impaction of the component into the underlying cancellous bone.
Testing and specimens Ten fresh-frozen scapulae with a mean age of 76.1 ⫾ 10.3 years were stored at ⫺20°C. Before use, each specimen was thawed at 23.0°C ⫾ 2.0°C for 24 hours. Scapulae with any evidence of previous surgery, trauma, or
Angle (°) 10 20 30 20 20 20 20 20
glenoid dysplasia were not included. The cadaveric glenoids were potted in the specimen preparation tray by use of dental cement, with the surface of the glenoid aligned horizontally to permit proper surgical preparation and physiologic loading within the apparatus. Implant variables included 3 implant articular surface thicknesses of 2, 4, and 6 mm (measured at the center of the articular surface), with a fully congruent articulation. We assessed 4 glenohumeral surface radii of curvature conformities of 0, 3, 6, and 9 mm (ie, glenoid component radius minus humeral head radius) with a 4-mm-thick glenoid component. The sequence of testing of these implant variables was randomized for each specimen. Loading consisted of preconditioning by use of an axial cyclic (sinusoidal) load in pure compression of 0 to 250 N for 400 cycles at 1 Hz. The number of cycles was determined during a pilot study, where 400 cycles was determined to adequately precondition the glenoid subchondral surface. Each of the implant variables was subjected to block-randomized loads (Table I). This consisted of 6 directions of load application, described with respect to the anatomic position of the glenoid, and 3 angles of load application, described with respect to a normal drawn to the articular surface of the glenoid. The superoposterior and inferoposterior loads were chosen because they represent physiologic loading conditions that are known to produce posterior subluxation in the shoulder.13,26,45 The superior loads (up to 30°) were chosen because they represent common physiologic loading conditions in the shoulder.11,35 The 8 loading conditions in Table I were evaluated in random order within each block of implant variables. The protocol consisted of load application (linearly) from 0 to 250 N (at 25 N/s), a 1-second constant load, and unloading (at 25 N/s). During the constant-load phase, the 4 displacements were recorded at 10 Hz, and the mean value was determined. A constant-load magnitude of 250 N (36% of a body weight of 68 kg) was used throughout, as this is within the range of normal physiologic loading anticipated at the glenohumeral joint.11,28,35 A rate of 25 N/s and a 1-second constant load were chosen because these are representative of the loading rate and rhythm experienced during an average glenohumeral abduction arc.11,28,35
Statistical methods A power analysis was used before the initiation of the study to estimate the sample sizes needed to detect what would be considered clinically relevant differences between
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Figure 3 Direction of load application. Displacement of the glenoid component is shown as a function of the direction of load application. This is subdivided for each site of micromotion measurement about the glenoid. Micromotion is shown for the 4-mm component with a glenohumeral conformity of 0 mm, at an angle of load application of 20°. S, Superior; I, inferior; A, anterior; P, posterior; S-P, superoposterior; I-P, inferoposterior.
implant variables. By use of data from a related study in our laboratory, mean values and SDs for parameter outcomes of interest were obtained.6 The ␣ level was set at .05 to detect a desired difference of 80% of normal values, and a minimal detectable difference was set at the difference between the highest and lowest means to ensure that the worst-case scenario was considered. To detect a difference between implant variables, 9 specimens were required, by use of a repeated-measures analysis of variance. Mean values and SDs were calculated for the 10 specimens, and repeated-measures analyses of variance were performed to demonstrate the effect of each variable. Student-Newman-Keuls post hoc tests were also conducted to make pairwise multiple comparisons between individual variables. Statistical analyses and graphic illustrations were performed with SigmaStat and SigmaPlot software, respectively (Systat Software, Richmond, CA).
RESULTS
Figure 4 Angle of load application. Contralateral displacement of the glenoid component is shown as a function of the angle of load application. For superior loading, micromotion at the inferior site is shown for each angle of load application. The mean values are shown for all implant variables.
Loading in each direction produced compression at the site corresponding to the loading direction (ie, ipsilateral compression) and distraction from the bone surface on the opposite side (ie, contralateral liftoff) (Figure 3). The off-axis sites (ie, anterior and posterior sites for superior loading) were less predictable, demonstrating either compression or distraction with a magnitude less than the locations co-planar with the loading vector. This pattern of ipsilateral compression and contralateral distraction was consistent for all implant variables and loading directions and angles. Glenoid component micromotion increased when the angle of the applied load was increased (P ⬍ .05) (Figure 4). For all implant variables, mean contralateral displacement was 3.98 ⫾ 5.96 m for a 10°
angle of load application, 12.37 ⫾ 10.94 m for a 20° angle of load application, and 28.05 ⫾ 25.44 m for a 30° angle of load application (Table II). This effect of the angle of load application was found to be consistent for each implant variable. All articular thicknesses and glenohumeral conformities displayed more micromotion when the angle of the load vector was increased (P ⬍ .05). A decrease in implant thickness resulted in increased stability of the glenoid component, as measured by the variable of contralateral liftoff (P ⬍ .05) (Figure 5). For all directions and angles of load application, mean contralateral displacement was 16.73 ⫾ 22.71 m for a 2-mm articular thickness,
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19.68 ⫾ 15.38 m for a 4-mm articular thickness, and 25.22 ⫾ 18.56 m for a 6-mm articular thickness (Table II). This was found to be statistically significant for the 2-mm component relative to the 4- and 6-mm components (P ⬍ .05). The 2-mm component showed less distraction from the surface of the bone opposite the site of load application, as well as higher compression at the loading site, with respect to both the 4- and 6-mm components (P ⬍ .05). These findings with respect to thickness were found to be significant for inferior and posterior loading directions (P ⬍ .05). Nonconformity between articular surfaces resulted in increased stability of the glenoid component, as measured by the variable of contralateral liftoff (P ⬍ .05) (Figure 6). For all directions and angles of load application, mean contralateral displacement was 6.81 ⫾ 5.12 m for 0 mm of glenohumeral conformity, 6.06 ⫾ 4.65 m for 3 mm of glenohumeral conformity, 5.85 ⫾ 3.94 m for 6 mm of glenohumeral conformity, and 4.46 ⫾ 3.42 m for 9 mm of glenohumeral conformity (Table II). The 6- and 9-mm mismatches were found to produce the least contralateral micromotion (P ⬍ .05). The presence of nonconformity between articular surfaces resulted in less distraction from the surface of the bone opposite the site of load application and less compression at the loading site (P ⬍ .05). These findings with respect to glenohumeral conformity were found to be significant for superior and posterior loading directions (P ⬍ .05). DISCUSSION The glenoid implant articular thickness was found to influence implant stability. Less distraction from the surface of the bone opposite the site of load application was observed as the thickness was decreased. This can be explained by understanding the effect of eccentrically applied loads at the prosthesis-bone interface. As described earlier, increasing the load angle increases the eccentricity of applied loads, resulting in an increase in the associated micromotion. Therefore, by decreasing the eccentricity of loads that are transferred across the prosthesis-bone interface, glenoid component loosening may be decreased. As the thickness of the implant is increased, eccentrically applied loads are transferred more peripherally, when they are transferred to the bone. This is simply a result of the load line of action having to pass through a finite component thickness. When this line of action becomes more eccentric, the load transferred to the prosthesis-bone interface also becomes more peripheral. This effect is compounded if the implant thickness also increases. Decreasing implant thickness has previously been shown to be essential to reducing the eccentricity of loads and thereby minimizing liftoff of an all-polyethylene component.10,25 However, as a result of the mechanical properties of
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polyethylene, particularly with regard to wear, a lower limit exists as to the durable thickness of the component that can be used. It is possible that a substantial reduction in implant thickness can be accomplished by use of an all-metal component design. For this reason, we chose to use an all-metal prosthetic component. In addition, thicker components currently in use may overstuff the joint, increase lateral offset of the humeral head, and affect the moment arms of muscles attached to the humerus, altering the associated articular joint contact pressure.18 This will also affect the forces and moments applied to the prosthesis-bone interface. Glenohumeral conformity was also found to influence glenoid implant stability. The presence of any level of nonconformity between the radii of curvature of the 2 articular surfaces resulted in less micromotion between implant and bone. When a load is applied through a humeral head to a glenoid component that has a perfectly conforming articulation, the applied load will always result in loading at the edge of the component (rim loading), because the humeral head is, in all probability, in contact with the entire surface, including the rim. However, nonconformity results in only a small arc of the humeral head in contact with the glenoid. Therefore, edge loading will not, in all likelihood, normally occur unless loads become more eccentric. This also will allow eccentric loads, acting through the center of curvature of the humeral head, to act at a smaller angle for any given load vector, as the location of the center of curvature will be closer to the articular surface as a result of the smaller radius of the humeral head. Therefore, this will also decrease the eccentricity of the applied loads. However, this scenario may require the use of a humeral head component that may not accurately reflect the normal anatomy.29,34 This may lead to an alteration in the magnitude and direction of the moment arms of the associated muscles acting across this joint, such as the rotator cuff and the deltoid.10,35 This may affect the forces and moments applied to the prosthesisbone interface and may also alter the associated kinematics of the joint and, therefore, requires further study. This nonconformity also allows translation of the humeral head on the glenoid, perhaps more accurately reproducing physiologic motion.22 However, the implications with regard to wear may warrant exploration. This study showed that the glenoid implant is very sensitive to eccentric loading. This was evidenced by the effect of the angle of the applied load, as well as the implant articular thickness and glenohumeral conformity. Relative to the array of eccentric loading vectors imparted for the various functions of the shoulder, the size of the base of the glenoid implant is small, and thus, stability is difficult to achieve.35 To reduce the eccentricity of loading, design consider-
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Table II Summary of contralateral displacement measurements Contralateral micromotion Loading information Direction Inferior Superior
Posterior Anterior Total
Articular thickness
Angle (°)
2 mm
4 mm
6 mm
20 10 20 30 20 20
6.71 ⫾ 12.82 4.00 ⫾ 10.27 14.27 ⫾15.48 51.89 ⫾ 69.33 10.16 ⫾ 8.81 13.35 ⫾ 19.54 16.73 ⫾ 22.71
25.73 ⫾ 19.77 3.52 ⫾ 5.94 19.14 ⫾ 17.67 42.37 ⫾ 32.09 15.66 ⫾ 8.03 11.65 ⫾ 8.78 19.68 ⫾ 15.38
30.13 ⫾ 22.16 4.75 ⫾ 8.84 23.44 ⫾ 16.30 51.06 ⫾ 39.32 21.09 ⫾ 13.65 20.84 ⫾ 11.06 25.22 ⫾ 18.56
ations, which shift the load line of action more centrally, may be beneficial from the viewpoint of implant stability. This study has shown that decreasing component thickness and glenohumeral conformity can achieve this goal. To some degree, the micromotion noted previously can be likened to a rigid body on an elastic foundation. With uniform mechanical properties, this structure, when loaded within its middle third, undergoes no tension in response to an offcenter (ie, eccentric) load.33 For each load applied to the glenoid component throughout this study, a simple geometric analysis showed that the load at the component-bone interface was located outside of the middle third of the component. As a result, tension (ie, distraction) in response to an off-center compressive load is likely. The mechanical properties of the implant are assumed to be constant, because it is constructed of a homogeneous material, but the underlying cancellous bone has been shown to be both heterogeneous and anisotropic.2,4 However, this simplification is thought to introduce minimal errors. This tension in response to an off-center compressive load generates liftoff of implants from the bone interface and is thought to contribute to implant loosening. By use of a simple geometric analysis, as the loading angle is increased from 20° to 30°, for a 4-mm articular thickness, the location of the load at the component-bone interface is moved 4.6 mm peripherally away from the center of the component. Similarly, for a constant angle of load application of 20°, as the thickness of the component is increased from 2 to 4 mm, the location of the load at the componentbone interface is moved 0.7 mm peripherally away from the center of the component. Our findings are in agreement with previous investigations. Wang et al42 confirmed that eccentric loading of conforming shoulder implants produces contact near the glenoid edge. Walch et al40 showed that increased glenohumeral conformity leads to a higher incidence of radiolucent lines. Oosterom et al31 demonstrated that decreasing glenohumeral conformity
may increase component micromotion, which is in contradiction to our findings. However, they tested a cemented polyethylene component in a bone substitute model, using loads 3 times greater in magnitude than those in our study. The effect of glenohumeral conformity on contralateral liftoff was not evident at all angles of load application. This effect of decreased contralateral liftoff was evident for an angle of load application of 10° and 20° but was not represented at a load application angle of 30°. We hypothesize that this may be because 30° is sufficiently large that edge loading occurs even with 9 mm of glenohumeral nonconformity. This may effectively negate the effect of decreased glenohumeral conformity resulting in decreased contralateral liftoff. This study contains some potential limitations. First, component micromotion was measured with respect to bone, based on displacement transducers located a short distance away from the edge of the implant. As a result, the values obtained overestimate the relative motion of the component with respect to bone. However, the location of the measurement sites was consistent within a particular specimen, therefore allowing comparisons between variables within each specimen. Second, the measured micromotion represents a combination of the desired motion of the glenoid component with respect to bone, deformation of both the component and bone, and movement of the transducers with respect to both the component and bone. These additional motions are thought to be minimized through rigid attachment of the transducers to the component and bone at distances from the implant-bone interface that were as small as possible. In addition, the micromotion was measured in a single plane, although it is recognized that this is a representation of a 3-dimensional motion. Third, each specimen may have been conditioned by applying increasing loads sequentially, thereby allowing the sequence of the testing protocol to influence the results. Although the order of fixation technique was
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Table II Continued (mean ⴞ SD) (m) Glenohumeral conformity 0 mm
3 mm
6 mm
9 mm
Total
0.29 ⫾ 0.04 5.90 ⫾ 5.16 9.23 ⫾ 8.49 14.32 ⫾ 10.00 3.96 ⫾ 6.90 7.13 ⫾ 0.11 6.81 ⫾ 5.12
0.02 ⫾ 0.13 3.37 ⫾ 4.34 9.36 ⫾ 8.37 13.18 ⫾ 9.16 2.57 ⫾ 2.47 7.88 ⫾ 3.40 6.06 ⫾ 4.65
0.05 ⫾ 0.02 5.95 ⫾ 5.03 5.28 ⫾ 4.96 11.81 ⫾ 9.11 2.99 ⫾ 0.13 9.03 ⫾ 4.37 5.85 ⫾ 3.94
0.02 ⫾ 0.02 0.35 ⫾ 2.14 5.86 ⫾ 5.29 11.71 ⫾ 9.06 0.28 ⫾ 0.61 8.54 ⫾ 3.37 4.46 ⫾ 3.42
8.99 ⫾ 7.85 3.98 ⫾ 5.96 12.37 ⫾ 10.94 28.05 ⫾ 25.44 8.10 ⫾ 5.80 11.20 ⫾ 7.23
Figure 5 Articular thickness. Contralateral displacement of the glenoid component is shown as a function of implant thickness. The mean values are shown for all directions and angles of load application.
Figure 6 Glenohumeral conformity. Contralateral displacement of the glenoid component is shown as a function of glenohumeral conformity. The mean values are shown for all directions and angles of load application.
kept constant for each specimen, the loading protocol was block-randomized for each implantation technique to minimize these conditioning effects. In addition, each specimen was preconditioned to 400 cycles after component insertion and before data collection. These precautions were thought to be effective in eliminating any effect of the order of testing, as none could be detected in our analysis. Fourth, there are likely differences between the loading used in this study and that occurring in vivo. Given that the glenohumeral joint is subjected to a large range of motion in vivo, we chose to use a loading protocol that accommodated for a wide range of load directions and a variation in the line of action. Thus, this loading protocol likely enveloped a number of common in vivo loading patterns. Fifth, our in vitro model differs from the in vivo state. By using resected cadaveric glenoids, we have eliminated the significant ligamentous and muscular contributions to stabilization.15,19 Furthermore, the quality of bone may not have been fully retained because of the storage process and the lack of a physiologic environment. This may result in the mechanical properties differing from the in vivo
state. However, by use of a repeated-measures design, these concerns should be less pronounced. Finally, each of these limitations may have contributed to the relatively high variability between specimens. This is evidenced by our SD as a measure of variability for each measurement in our study. Each SD was relatively large, comprising, on average, between 50% and 75% of the mean value. However, our statistical analysis accounted for this variance, and therefore, the statistically significant effects noted are thought to be true. In conclusion, this study has shown that an all-metal glenoid implant design with reduced articular thickness and decreased glenohumeral conformity may be efficacious from the perspective of reducing the eccentricity of applied loads and thereby improving glenoid implant stability. The possibility of metal-onmetal articulations for the shoulder may be efficacious from the viewpoint of wear; however, this awaits further study. It is important to note that, although this study used an all-metal glenoid component, with respect to the relative effects of the implant variables addressed (ie, articular thickness and glenohumeral
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conformity), the influence of these variables should be equally applicable to either a metal-backed polyethylene or cemented polyethylene implant.21,37 REFERENCES
1. Anglin C, Wyss UP, Pichora DR. Mechanical testing of shoulder prostheses and recommendations for glenoid design. J Shoulder Elbow Surg 2000;9:323-31. 2. Barea C, Mansat P, Hobatho MC, Darmana R, Mansat M. Mechanical properties of the glenoid cancellous bone. Trans Orthop Res Soc 1997;22:877. 3. Barrett WP, Franklin JL, Jackins SE, Wyss CR, Matsen FA. Total shoulder arthroplasty. J Bone Joint Surg Am 1987;69:865-72. 4. Batte SWP, Johnson JA, Lee TY, Patterson SD, King GJW, Chess DG. Cancellous bone of the glenoid: relationship of quantitative CT density to mechanical strength. Presented at the 30th Annual Meeting of the Canadian Orthopaedic Research Society, Quebec City, Canada, May 1996. 5. Bicknell RT, Danter MD, Bennett J, Patterson SD, King GJW, Johnson JA, et al. An anthropometric evaluation of the glenoid: implications in the design and fixation of a shoulder prosthesis. Trans Orthop Res Soc 1998;23:270. 6. Bicknell RT, Liew A, Danter MR, Patterson SD, King GJW, Chess DG, et al. Does keel size, the use of screws, and the use of bone cement affect fixation of a metal glenoid implant? J Shoulder Elbow Surg 2003;2:268-75. 7. Boileau P, Avidor C, Krishnan SG, Walch G, Kempf JF, Mole D. Cemented polyethylene versus uncemented metal-backed glenoid components in total shoulder arthroplasty: a prospective, double-blind, randomized study. J Shoulder Elbow Surg 2002; 11:351-9. 8. Chess DG, Batte SWP, Patterson SD, Liew ASL, Danter M, King GJW, et al. Glenoid component design in shoulder arthroplasty: the effect of joint loading and a keel on intrinsic implant stability. Presented at the 31st Annual Meeting of the Canadian Orthopaedic Research Society, Hamilton, Canada, June 1997. 9. Collins D, Tencer A, Sidles J, Matsen F. Edge displacement and deformation of glenoid components in response to eccentric loading. J Bone Joint Surg Am 1992;74:501-7. 10. de Leest O, Rozing PM, Rozendaal LA, van der Helm FCT. Influence of glenohumeral prosthesis geometry and placement on shoulder muscle forces. Clin Orthop Relat Res 1996:222-33. 11. Ellenbecker TS, Mattalino AJ. Glenohumeral joint range of motion and rotator cuff strength following arthroscopic anterior stabilization with thermal capsulorraphy. J Orthop Sports Phys Ther 1999; 29:160-7. 12. Flatow EL, Ateshian GA, Soslowsky LJ, et al. Computer simulation of glenohumeral and patellofemoral subluxation. Clin Orthop Relat Res 1994:28-33. 13. Flatow EL, Miniaci A, Evans PJ, Simonian PT, Warren RF. Instability of the shoulder: complex problems and failed repairs. J Bone Joint Surg 1998;80:284-98. 14. Franklin JL, Barrett WP, Jackins SE, Matsen FA III. Glenoid loosening in total shoulder arthroplasty. Association with rotator cuff deficiency. J Arthroplasty 1988;3:39-46. 15. Friedman RJ. Prospective analysis of total shoulder arthroplasty biomechanics. Am J Orthop 1997;26:265-70. 16. Green A. Current concepts of shoulder arthroplasty. Instr Course Lect 1998;47:127-33. 17. Horowitz P, Hill W. The art of electronics. 2nd ed. New York: Cambridge University Press; 1990. p. 1001–2. 18. Iannotti JP, Williams GR. Total shoulder arthroplasty: factors influencing prosthetic design. Orthop Clin North Am 1998; 29:377-91. 19. Jobe CM, Iannotti JP. Limits imposed on glenohumeral motion by joint geometry. J Shoulder Elbow Surg 1995;4:281-5.
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20. Karduna AR, Williams GR, Iannotti JP, Williams JL. Total shoulder arthroplasty biomechanics: a study of the forces and strains at the glenoid component. J Biomech Eng 1998;120:92-9. 21. Karduna AR, Williams GR, Williams JL, Iannotti JP. Joint stability after total shoulder arthroplasty in a cadaver model. J Shoulder Elbow Surg 1997;6:506-11. 22. Karduna AR, Williams GR, Williams JL, Iannotti JP. Glenohumeral joint translations before and after total shoulder arthroplasty. J Bone Joint Surg 1997;79:1166-74. 23. Lacroix D, Murphy LA, Prendergast PJ. Three-dimensional finite element analysis of glenoid replacement prostheses: a comparison of keeled and pegged anchorage systems. J Biomech Eng 2000;122:430-6. 24. Lazarus MD, Jensen KL, Southworth C, Matsen FA III. The radiographic evaluation of keeled and pegged glenoid component insertion. J Bone Joint Surg Am 2002;84:1174-82. 25. Liew A, Batte SWP, Patterson SD, King GJW, Johnson JA, Chess DG. Glenoid component thickness has a significant effect on implant-bone interface stability in total shoulder arthroplasty. Trans Orthop Res Soc 1997;22:880. 26. Malicky DM, Soslowsky LJ, Blasier RB. Inferior glenohumeral subluxation: active and passive stabilization in a biomechanical model. Trans Orthop Res Soc 1995;20:679. 27. Martin SD, Zurakowski D, Thornhill TS. Uncemented glenoid component in total shoulder arthroplasty. Survivorship and outcomes. J Bone Joint Surg Am 2005;87:1284-92. 28. McMahon PJ, Debski RE, Thompson WO, Warner JJ, Fu FH, Woo SL. Shoulder muscle forces and tendon excursions during glenohumeral abduction in the scapular plane. J Shoulder Elbow Surg 1995;4:199-208. 29. McPherson EJ, Friedman RJ, An YH, Chokesi R, Dooley RL. Anthropometric study of normal glenohumeral relationships. J Shoulder Elbow Surg 1997;6:105-12. 30. Mileti J, Boardman ND, Sperling JW, Cofield RH, Torchia ME, O’Driscoll SW, et al. Radiographic analysis of polyethylene glenoid components using modern cementing techniques. J Shoulder Elbow Surg 2004;13:492-8. 31. Oosterom R, Rozing PM, Verdonschot N, Bersee HF. Effect of joint conformity on glenoid component fixation in total shoulder arthroplasty. Proc Inst Mech Eng [H] 2004;218:339-47. 32. Orr TE, Carter DR, Schurman DJ. Stress analysis of glenoid component designs. Clin Orthop Relat Res 1988:217-24. 33. Ozkaya N, Nordin M. Fundamentals of biomechanics: equilibrium, motion and deformation. New York: Van Nostrand Reinhold; 1991. p. 291–315. 34. Pearl ML, Volk AG. Coronal plane geometry of the proximal humerus relevant to prosthetic arthroplasty. J Shoulder Elbow Surg 1996;5:320-6. 35. Poppen NK, Walker PS. Forces at the glenohumeral joint in abduction. Clin Orthop Relat Res 1978:165-70. 36. Rahme H, Mattsson P, Larsson S. Stability of cemented allpolyethylene keeled glenoid components. A radiostereometric study with a two-year follow-up. J Bone Joint Surg Br 2004;86: 856-60. 37. Severt R, Thomas BJ, Tsenter MJ, Amstutz HC, Kabo JM. The influence of conformity and constraint on translational forces and frictional torque in total shoulder arthroplasty. Clin Orthop Relat Res 1993:151-8. 38. Skirving AP. Total shoulder arthroplasty: current problems and possible solutions. J Orthop Sci 1999;4:42-53. 39. Torchia ME, Cofield RH, Settergren CR. Total shoulder arthroplasty with the Neer prosthesis: long-term results. J Shoulder Elbow Surg 1997;6:495-505. 40. Walch G, Edwards TB, Boulahia A, Boileau P, Mole D, Adeleine P. The influence of glenohumeral prosthetic mismatch on glenoid radiolucent lines: results of a multicenter study. J Bone Joint Surg Am 2002;84:2186-91.
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41. Wallace AL, Phillips RL, MacDougal GA, Walsh WR, Sonnabend DH. Resurfacing of the glenoid in total shoulder arthroplasty. J Bone Joint Surg Am 1999;81:510-9. 42. Wang VM, Krishnan R, Ugwonali OF, Flatow EL, Bigliani LU, Ateshian GA. Biomechanical evaluation of a novel glenoid design in total shoulder arthroplasty. J Shoulder Elbow Surg 2005; 14(Suppl S):129S-40S. 43. Williams GR, Abboud JA. Total shoulder arthroplasty: glenoid component design. J Shoulder Elbow Surg 2005;14:122S-8S.
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44. Wirth MA, Rockwood CA Jr. Complications of total shoulder replacement arthroplasty. J Bone Joint Surg Am 1996;78:603-16. 45. Wuelker N, Brewe F, Sperveslage C. Passive glenohumeral joint stabilization: a biomechanical study. J Shoulder Elbow Surg 1994;3:129-34. 46. Yian EH, Werner CM, Nyffeler RW, Pfirrmann CW, Ramappa A, Sukthankar A, et al. Radiographic and computed tomography analysis of cemented pegged polyethylene glenoid components in total shoulder replacement. J Bone Joint Surg Am 2005;87:1928-36.
The objective of this study was to determine the effect of implant thickness and glenohumeral conformity on fixation of an all-metal glenoid component. A stainless steel glenoid component was designed and implanted in 10 cadaveric scapulae. A testing apparatus capable of producing a loading vector at various angles, magnitudes, and directions was used. The independent variables included 6 directions and 3 angles of joint load, 3 implant thicknesses, and 4 glenohumeral conformities. Implant micromotion relative to bone was measured by use of 4 displacement transducers at the superior, inferior, anterior, and posterior sites. The components displayed a consistent response to loading of ipsilateral compression and contralateral distraction. Stability decreased as the load application angle increased (P ⬍ .05). A decrease in the implant thickness and glenohumeral conformity resulted in increased implant stability (P ⬍ .05). Decreasing implant thickness and glenohumeral conformity reduce the eccentric component of loading and may improve the durability of glenoid implants. (J Shoulder Elbow Surg 2007;16:631-639.)
L oosening of the glenoid component remains a fac-
tor limiting the long-term success of total shoulder arthroplasty.7,16,18,27,36,41,43,44,46 Furthermore, glenoid loosening has been shown to be significantly
From the Departments of aSurgery, bMechanical and Materials Engineering, and eMedical Biophysics, University of Western Ontario, and cHand and Upper Limb Centre, St Joseph’s Health Centre, London, and dCentral Florida Orthopaedic Surgery Associates, Lakeland. Reprint requests: Ryan T. Bicknell, MD, MSc, FRCSC, Hand and Upper Limb Centre, St Joseph’s Health Centre, 268 Grosvenor St, London, Ontario, Canada, N6A 4L6 (E-mail: rtbickne@ yahoo.ca). Copyright © 2007 by Journal of Shoulder and Elbow Surgery Board of Trustees. 1058-2746/2007/$32.00 doi:10.1016/j.jse.2006.10.014
correlated with deteriorating functional results and increasing pain.7,24,39,43 This is thought to be a result of several factors, including inadequate implant-bone fixation and eccentric loading. Implant-bone fixation has been investigated extensively. Various designs have attempted to achieve fixation to bone through the use of anchoring systems on the undersurface of the glenoid component. In vitro studies have been conducted to assess the strength of fixation and the stability or micromotion of various implant designs relative to bone.1,6,8,31,32 Theoretic finite-element studies have also analyzed the state of stress in the implant-bone structure.15,23,32,38 However, these studies have not conclusively defined the optimal methods by which to maximize the stability between the glenoid component and bone. Previous studies have also indicated that the glenoid component is very sensitive to eccentric loading (ie, loading that occurs nonperpendicular to the glenoid surface).6,9,25 Distraction or liftoff of the implant from the bone surface, potentially inhibiting bone ingrowth or producing interface tension, is a concern with eccentric loading. In extreme arm positions or with certain pathologic conditions, subluxation of the humeral head may lead to rim loading of the glenoid.3,14,20,22 It has been shown that advances in bone preparation and cementing techniques can reduce the effects of eccentric loading that act to cause glenoid component distraction and loosening.9,30 Reducing the eccentricity of the applied load (and also rim loading) by using a thinner implant has also been shown to minimize the distraction of the glenoid component.25 In addition, reducing glenohumeral conformity may also reduce eccentric loading by minimizing rim loading.21,22,37 The objectives of this study were to determine the effects of implant articular thickness and glenohumeral conformity on the stability of uncemented allmetal glenoid components, as quantified by the magnitude of micromotion between implant and bone. We hypothesized that decreasing implant articular
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Figure 1 Testing apparatus. A, The loading apparatus consisted of 2 pneumatic actuators located orthogonal to each other, with an adjustable loading stage to permit loading in any direction. Actuators 1 and 2 produced the medial and transverse (superior [S], inferior [I], anterior [A], and posterior [P]) compressive forces, respectively. B, Micromotion measurements were performed with LVDTs. These included an LVDT barrel attached to the bone and an LVDT core attached to the implant.
thickness and decreasing glenohumeral conformity would improve glenoid implant stability. MATERIALS AND METHODS Testing apparatus and instrumentation A pneumatic loading apparatus and measurement system was used, as described in a previous study6 (Figure 1). This consisted of a loading stage fitted with 2 pneumatic actuators (SR-243-D; Bimba Manufacturing, Monee, IL), capable of supplying loads up to 492.5 N at 0.345 MPa of air pressure. The actuators were controlled by 2 virtual instrument device drivers written in LabVIEW software (National Instruments, Austin, TX), and actuator pressure was controlled by Proportional Pressure Controllers (PPC5AAAA-AGCB-BBB-JB; MAC Valves, Wixom, MI). A steel ball was used to model the humeral component and to provide a surface to articulate with the glenoid component.6,9 For this study, we used 3 spheres of different diameters that enabled us to vary glenohumeral conformity. The actuators directly loaded the ball bearing placed on top of the glenoid component (Figure 1). By varying the load from each actuator, the target magnitude and line of action
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of the resultant force could be achieved. The loading stage was adjustable and allowed rotation of the glenoid in the sagittal plane, permitting loading in any direction. This allowed free displacement of the humeral component in the plane of the glenoid component. To minimize nonphysiologic friction, the articular surfaces were coated with bovine serum (HyClone, Logan, UT). To quantify micromotion of the implant with respect to bone, 4 linear variable differential transducers (LVDTs) were used (Figure 1, B). These were chosen for their highly linear response for a given displacement (error of 0.01%), availability in various sizes, cost-effectiveness, and high sensitivity.17 A support ring served as a mounting stage for 4 LVDT barrels (Lucas-Schaevitz 049-XS-B; Durham Instruments, Pickering, Ontario, Canada). The LVDT ring was fastened to the cortical glenoid neck with screws (diameter, 4.2 mm). To permit attachment of the LVDT cores, four 0.9-mm Kirschner wire outriggers were fixed to each glenoid component, two along each of the superoinferior and anteroposterior axes. This permitted measurement at the superior, inferior, anterior, and posterior locations. After the LVDT ring was fixed to the glenoid, the component was implanted and the LVDT cores aligned and placed within their respective barrels. The cores were then fastened to the Kirschner wire outriggers by use of a swivel hinge, located a distance of 5 mm from the glenoid component. Before experimental testing, each of the LVDT signal transducers was calibrated in accordance with technical manuals provided by the manufacturer. Throughout the study, data were collected and converted from LVDT output voltages to LVDT displacements by use of LabVIEW software (National Instruments).
Surgical technique Glenoid preparation followed the published protocol for the 3M Neer II prosthesis (3M, Orthopaedic Products Division, St Paul, MN). The Neer II glenoid reamer (3M, Orthopaedic Products Division) was used to complete surface preparation. To ensure the best possible conformity of the prosthesis to the glenoid surface, this reamer had a radius of curvature of 28 mm, which is equal to the radius of curvature of the component articular undersurface.9 Hand reaming of the lateral surface of the articulation was performed until all cartilage was removed. A hand bur was used to make a slot for each keel through the subchondral bone, and the component was then impacted into the glenoid. The reamed surface of the glenoid was irrigated by use of 0.9% normal saline solution throughout testing, to limit drying of the specimen. The same surgeon conducted all glenoid preparations and component implantations.
Implant design This study used a prototype all-metal stainless steel glenoid component, with a cross-keel, as developed and evaluated in a previous study6 (Figure 2). The articular surface of the component was based on the Neer II component (3M, Orthopaedic Products Division).12,16 This includes an articular surface radius of curvature of 25 mm; an undersurface radius of curvature of 28 mm; coronal and transverse articular surface widths of 32.5 and 24.5 mm, respectively; and an articular surface thickness of 4 mm. The
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Table I Loading protocol: Direction and angle of load application Direction Superior
Inferior Anterior Posterior Superoposterior Inferoposterior
Figure 2 Implant design. The glenoid implant used in this study is shown in the coronal and sagittal planes. In addition to the keel present in the coronal plane in the Neer II implant design, the prototype implant also combined a keel in the transverse plane. A, Articular surface thickness; B, Neer II keel; C, large keel from previous study6; D, articular surface; E, screw holes from previous study6; F, articular undersurface; G, coronal keel; H, transverse keel; I, coronal component width; J, transverse component width.
component utilized a cross-keel design (as opposed to a single-keeled or pegged design), to maximize contact, based on the anatomic dimensions of the glenoid vault in both planes (transverse and coronal), as determined in a previous study.5 In addition to the (standard) keel in the coronal plane, our glenoid component had another keel placed orthogonally in the transverse plane. This provided the potential to counteract superiorly and inferiorly directed joint forces and was shown in a previous study to decrease glenoid component micromotion.6 The medial edge of each keel was tapered to assist with the impaction of the component into the underlying cancellous bone.
Testing and specimens Ten fresh-frozen scapulae with a mean age of 76.1 ⫾ 10.3 years were stored at ⫺20°C. Before use, each specimen was thawed at 23.0°C ⫾ 2.0°C for 24 hours. Scapulae with any evidence of previous surgery, trauma, or
Angle (°) 10 20 30 20 20 20 20 20
glenoid dysplasia were not included. The cadaveric glenoids were potted in the specimen preparation tray by use of dental cement, with the surface of the glenoid aligned horizontally to permit proper surgical preparation and physiologic loading within the apparatus. Implant variables included 3 implant articular surface thicknesses of 2, 4, and 6 mm (measured at the center of the articular surface), with a fully congruent articulation. We assessed 4 glenohumeral surface radii of curvature conformities of 0, 3, 6, and 9 mm (ie, glenoid component radius minus humeral head radius) with a 4-mm-thick glenoid component. The sequence of testing of these implant variables was randomized for each specimen. Loading consisted of preconditioning by use of an axial cyclic (sinusoidal) load in pure compression of 0 to 250 N for 400 cycles at 1 Hz. The number of cycles was determined during a pilot study, where 400 cycles was determined to adequately precondition the glenoid subchondral surface. Each of the implant variables was subjected to block-randomized loads (Table I). This consisted of 6 directions of load application, described with respect to the anatomic position of the glenoid, and 3 angles of load application, described with respect to a normal drawn to the articular surface of the glenoid. The superoposterior and inferoposterior loads were chosen because they represent physiologic loading conditions that are known to produce posterior subluxation in the shoulder.13,26,45 The superior loads (up to 30°) were chosen because they represent common physiologic loading conditions in the shoulder.11,35 The 8 loading conditions in Table I were evaluated in random order within each block of implant variables. The protocol consisted of load application (linearly) from 0 to 250 N (at 25 N/s), a 1-second constant load, and unloading (at 25 N/s). During the constant-load phase, the 4 displacements were recorded at 10 Hz, and the mean value was determined. A constant-load magnitude of 250 N (36% of a body weight of 68 kg) was used throughout, as this is within the range of normal physiologic loading anticipated at the glenohumeral joint.11,28,35 A rate of 25 N/s and a 1-second constant load were chosen because these are representative of the loading rate and rhythm experienced during an average glenohumeral abduction arc.11,28,35
Statistical methods A power analysis was used before the initiation of the study to estimate the sample sizes needed to detect what would be considered clinically relevant differences between
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Figure 3 Direction of load application. Displacement of the glenoid component is shown as a function of the direction of load application. This is subdivided for each site of micromotion measurement about the glenoid. Micromotion is shown for the 4-mm component with a glenohumeral conformity of 0 mm, at an angle of load application of 20°. S, Superior; I, inferior; A, anterior; P, posterior; S-P, superoposterior; I-P, inferoposterior.
implant variables. By use of data from a related study in our laboratory, mean values and SDs for parameter outcomes of interest were obtained.6 The ␣ level was set at .05 to detect a desired difference of 80% of normal values, and a minimal detectable difference was set at the difference between the highest and lowest means to ensure that the worst-case scenario was considered. To detect a difference between implant variables, 9 specimens were required, by use of a repeated-measures analysis of variance. Mean values and SDs were calculated for the 10 specimens, and repeated-measures analyses of variance were performed to demonstrate the effect of each variable. Student-Newman-Keuls post hoc tests were also conducted to make pairwise multiple comparisons between individual variables. Statistical analyses and graphic illustrations were performed with SigmaStat and SigmaPlot software, respectively (Systat Software, Richmond, CA).
RESULTS
Figure 4 Angle of load application. Contralateral displacement of the glenoid component is shown as a function of the angle of load application. For superior loading, micromotion at the inferior site is shown for each angle of load application. The mean values are shown for all implant variables.
Loading in each direction produced compression at the site corresponding to the loading direction (ie, ipsilateral compression) and distraction from the bone surface on the opposite side (ie, contralateral liftoff) (Figure 3). The off-axis sites (ie, anterior and posterior sites for superior loading) were less predictable, demonstrating either compression or distraction with a magnitude less than the locations co-planar with the loading vector. This pattern of ipsilateral compression and contralateral distraction was consistent for all implant variables and loading directions and angles. Glenoid component micromotion increased when the angle of the applied load was increased (P ⬍ .05) (Figure 4). For all implant variables, mean contralateral displacement was 3.98 ⫾ 5.96 m for a 10°
angle of load application, 12.37 ⫾ 10.94 m for a 20° angle of load application, and 28.05 ⫾ 25.44 m for a 30° angle of load application (Table II). This effect of the angle of load application was found to be consistent for each implant variable. All articular thicknesses and glenohumeral conformities displayed more micromotion when the angle of the load vector was increased (P ⬍ .05). A decrease in implant thickness resulted in increased stability of the glenoid component, as measured by the variable of contralateral liftoff (P ⬍ .05) (Figure 5). For all directions and angles of load application, mean contralateral displacement was 16.73 ⫾ 22.71 m for a 2-mm articular thickness,
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19.68 ⫾ 15.38 m for a 4-mm articular thickness, and 25.22 ⫾ 18.56 m for a 6-mm articular thickness (Table II). This was found to be statistically significant for the 2-mm component relative to the 4- and 6-mm components (P ⬍ .05). The 2-mm component showed less distraction from the surface of the bone opposite the site of load application, as well as higher compression at the loading site, with respect to both the 4- and 6-mm components (P ⬍ .05). These findings with respect to thickness were found to be significant for inferior and posterior loading directions (P ⬍ .05). Nonconformity between articular surfaces resulted in increased stability of the glenoid component, as measured by the variable of contralateral liftoff (P ⬍ .05) (Figure 6). For all directions and angles of load application, mean contralateral displacement was 6.81 ⫾ 5.12 m for 0 mm of glenohumeral conformity, 6.06 ⫾ 4.65 m for 3 mm of glenohumeral conformity, 5.85 ⫾ 3.94 m for 6 mm of glenohumeral conformity, and 4.46 ⫾ 3.42 m for 9 mm of glenohumeral conformity (Table II). The 6- and 9-mm mismatches were found to produce the least contralateral micromotion (P ⬍ .05). The presence of nonconformity between articular surfaces resulted in less distraction from the surface of the bone opposite the site of load application and less compression at the loading site (P ⬍ .05). These findings with respect to glenohumeral conformity were found to be significant for superior and posterior loading directions (P ⬍ .05). DISCUSSION The glenoid implant articular thickness was found to influence implant stability. Less distraction from the surface of the bone opposite the site of load application was observed as the thickness was decreased. This can be explained by understanding the effect of eccentrically applied loads at the prosthesis-bone interface. As described earlier, increasing the load angle increases the eccentricity of applied loads, resulting in an increase in the associated micromotion. Therefore, by decreasing the eccentricity of loads that are transferred across the prosthesis-bone interface, glenoid component loosening may be decreased. As the thickness of the implant is increased, eccentrically applied loads are transferred more peripherally, when they are transferred to the bone. This is simply a result of the load line of action having to pass through a finite component thickness. When this line of action becomes more eccentric, the load transferred to the prosthesis-bone interface also becomes more peripheral. This effect is compounded if the implant thickness also increases. Decreasing implant thickness has previously been shown to be essential to reducing the eccentricity of loads and thereby minimizing liftoff of an all-polyethylene component.10,25 However, as a result of the mechanical properties of
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polyethylene, particularly with regard to wear, a lower limit exists as to the durable thickness of the component that can be used. It is possible that a substantial reduction in implant thickness can be accomplished by use of an all-metal component design. For this reason, we chose to use an all-metal prosthetic component. In addition, thicker components currently in use may overstuff the joint, increase lateral offset of the humeral head, and affect the moment arms of muscles attached to the humerus, altering the associated articular joint contact pressure.18 This will also affect the forces and moments applied to the prosthesis-bone interface. Glenohumeral conformity was also found to influence glenoid implant stability. The presence of any level of nonconformity between the radii of curvature of the 2 articular surfaces resulted in less micromotion between implant and bone. When a load is applied through a humeral head to a glenoid component that has a perfectly conforming articulation, the applied load will always result in loading at the edge of the component (rim loading), because the humeral head is, in all probability, in contact with the entire surface, including the rim. However, nonconformity results in only a small arc of the humeral head in contact with the glenoid. Therefore, edge loading will not, in all likelihood, normally occur unless loads become more eccentric. This also will allow eccentric loads, acting through the center of curvature of the humeral head, to act at a smaller angle for any given load vector, as the location of the center of curvature will be closer to the articular surface as a result of the smaller radius of the humeral head. Therefore, this will also decrease the eccentricity of the applied loads. However, this scenario may require the use of a humeral head component that may not accurately reflect the normal anatomy.29,34 This may lead to an alteration in the magnitude and direction of the moment arms of the associated muscles acting across this joint, such as the rotator cuff and the deltoid.10,35 This may affect the forces and moments applied to the prosthesisbone interface and may also alter the associated kinematics of the joint and, therefore, requires further study. This nonconformity also allows translation of the humeral head on the glenoid, perhaps more accurately reproducing physiologic motion.22 However, the implications with regard to wear may warrant exploration. This study showed that the glenoid implant is very sensitive to eccentric loading. This was evidenced by the effect of the angle of the applied load, as well as the implant articular thickness and glenohumeral conformity. Relative to the array of eccentric loading vectors imparted for the various functions of the shoulder, the size of the base of the glenoid implant is small, and thus, stability is difficult to achieve.35 To reduce the eccentricity of loading, design consider-
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Table II Summary of contralateral displacement measurements Contralateral micromotion Loading information Direction Inferior Superior
Posterior Anterior Total
Articular thickness
Angle (°)
2 mm
4 mm
6 mm
20 10 20 30 20 20
6.71 ⫾ 12.82 4.00 ⫾ 10.27 14.27 ⫾15.48 51.89 ⫾ 69.33 10.16 ⫾ 8.81 13.35 ⫾ 19.54 16.73 ⫾ 22.71
25.73 ⫾ 19.77 3.52 ⫾ 5.94 19.14 ⫾ 17.67 42.37 ⫾ 32.09 15.66 ⫾ 8.03 11.65 ⫾ 8.78 19.68 ⫾ 15.38
30.13 ⫾ 22.16 4.75 ⫾ 8.84 23.44 ⫾ 16.30 51.06 ⫾ 39.32 21.09 ⫾ 13.65 20.84 ⫾ 11.06 25.22 ⫾ 18.56
ations, which shift the load line of action more centrally, may be beneficial from the viewpoint of implant stability. This study has shown that decreasing component thickness and glenohumeral conformity can achieve this goal. To some degree, the micromotion noted previously can be likened to a rigid body on an elastic foundation. With uniform mechanical properties, this structure, when loaded within its middle third, undergoes no tension in response to an offcenter (ie, eccentric) load.33 For each load applied to the glenoid component throughout this study, a simple geometric analysis showed that the load at the component-bone interface was located outside of the middle third of the component. As a result, tension (ie, distraction) in response to an off-center compressive load is likely. The mechanical properties of the implant are assumed to be constant, because it is constructed of a homogeneous material, but the underlying cancellous bone has been shown to be both heterogeneous and anisotropic.2,4 However, this simplification is thought to introduce minimal errors. This tension in response to an off-center compressive load generates liftoff of implants from the bone interface and is thought to contribute to implant loosening. By use of a simple geometric analysis, as the loading angle is increased from 20° to 30°, for a 4-mm articular thickness, the location of the load at the component-bone interface is moved 4.6 mm peripherally away from the center of the component. Similarly, for a constant angle of load application of 20°, as the thickness of the component is increased from 2 to 4 mm, the location of the load at the componentbone interface is moved 0.7 mm peripherally away from the center of the component. Our findings are in agreement with previous investigations. Wang et al42 confirmed that eccentric loading of conforming shoulder implants produces contact near the glenoid edge. Walch et al40 showed that increased glenohumeral conformity leads to a higher incidence of radiolucent lines. Oosterom et al31 demonstrated that decreasing glenohumeral conformity
may increase component micromotion, which is in contradiction to our findings. However, they tested a cemented polyethylene component in a bone substitute model, using loads 3 times greater in magnitude than those in our study. The effect of glenohumeral conformity on contralateral liftoff was not evident at all angles of load application. This effect of decreased contralateral liftoff was evident for an angle of load application of 10° and 20° but was not represented at a load application angle of 30°. We hypothesize that this may be because 30° is sufficiently large that edge loading occurs even with 9 mm of glenohumeral nonconformity. This may effectively negate the effect of decreased glenohumeral conformity resulting in decreased contralateral liftoff. This study contains some potential limitations. First, component micromotion was measured with respect to bone, based on displacement transducers located a short distance away from the edge of the implant. As a result, the values obtained overestimate the relative motion of the component with respect to bone. However, the location of the measurement sites was consistent within a particular specimen, therefore allowing comparisons between variables within each specimen. Second, the measured micromotion represents a combination of the desired motion of the glenoid component with respect to bone, deformation of both the component and bone, and movement of the transducers with respect to both the component and bone. These additional motions are thought to be minimized through rigid attachment of the transducers to the component and bone at distances from the implant-bone interface that were as small as possible. In addition, the micromotion was measured in a single plane, although it is recognized that this is a representation of a 3-dimensional motion. Third, each specimen may have been conditioned by applying increasing loads sequentially, thereby allowing the sequence of the testing protocol to influence the results. Although the order of fixation technique was
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Table II Continued (mean ⴞ SD) (m) Glenohumeral conformity 0 mm
3 mm
6 mm
9 mm
Total
0.29 ⫾ 0.04 5.90 ⫾ 5.16 9.23 ⫾ 8.49 14.32 ⫾ 10.00 3.96 ⫾ 6.90 7.13 ⫾ 0.11 6.81 ⫾ 5.12
0.02 ⫾ 0.13 3.37 ⫾ 4.34 9.36 ⫾ 8.37 13.18 ⫾ 9.16 2.57 ⫾ 2.47 7.88 ⫾ 3.40 6.06 ⫾ 4.65
0.05 ⫾ 0.02 5.95 ⫾ 5.03 5.28 ⫾ 4.96 11.81 ⫾ 9.11 2.99 ⫾ 0.13 9.03 ⫾ 4.37 5.85 ⫾ 3.94
0.02 ⫾ 0.02 0.35 ⫾ 2.14 5.86 ⫾ 5.29 11.71 ⫾ 9.06 0.28 ⫾ 0.61 8.54 ⫾ 3.37 4.46 ⫾ 3.42
8.99 ⫾ 7.85 3.98 ⫾ 5.96 12.37 ⫾ 10.94 28.05 ⫾ 25.44 8.10 ⫾ 5.80 11.20 ⫾ 7.23
Figure 5 Articular thickness. Contralateral displacement of the glenoid component is shown as a function of implant thickness. The mean values are shown for all directions and angles of load application.
Figure 6 Glenohumeral conformity. Contralateral displacement of the glenoid component is shown as a function of glenohumeral conformity. The mean values are shown for all directions and angles of load application.
kept constant for each specimen, the loading protocol was block-randomized for each implantation technique to minimize these conditioning effects. In addition, each specimen was preconditioned to 400 cycles after component insertion and before data collection. These precautions were thought to be effective in eliminating any effect of the order of testing, as none could be detected in our analysis. Fourth, there are likely differences between the loading used in this study and that occurring in vivo. Given that the glenohumeral joint is subjected to a large range of motion in vivo, we chose to use a loading protocol that accommodated for a wide range of load directions and a variation in the line of action. Thus, this loading protocol likely enveloped a number of common in vivo loading patterns. Fifth, our in vitro model differs from the in vivo state. By using resected cadaveric glenoids, we have eliminated the significant ligamentous and muscular contributions to stabilization.15,19 Furthermore, the quality of bone may not have been fully retained because of the storage process and the lack of a physiologic environment. This may result in the mechanical properties differing from the in vivo
state. However, by use of a repeated-measures design, these concerns should be less pronounced. Finally, each of these limitations may have contributed to the relatively high variability between specimens. This is evidenced by our SD as a measure of variability for each measurement in our study. Each SD was relatively large, comprising, on average, between 50% and 75% of the mean value. However, our statistical analysis accounted for this variance, and therefore, the statistically significant effects noted are thought to be true. In conclusion, this study has shown that an all-metal glenoid implant design with reduced articular thickness and decreased glenohumeral conformity may be efficacious from the perspective of reducing the eccentricity of applied loads and thereby improving glenoid implant stability. The possibility of metal-onmetal articulations for the shoulder may be efficacious from the viewpoint of wear; however, this awaits further study. It is important to note that, although this study used an all-metal glenoid component, with respect to the relative effects of the implant variables addressed (ie, articular thickness and glenohumeral
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conformity), the influence of these variables should be equally applicable to either a metal-backed polyethylene or cemented polyethylene implant.21,37 REFERENCES
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