Effect of porous coating and loading conditions on total hip femoral stem stability

The Journal of Arthroplasty Vol. 10 No. 6 1995

Effect of Porous Coating and Loading C o n d i t i o n s on Total Hip F e m o r a l S t e m Stability

Timothy

F r a n c e s B. B i e g l e r , M S , * J e f f r e y D. R e u b e n , M D , P h D , * J R H a r r i g a n , ScD,*J- F u J. H o u , P h D , t a n d J o h n E. A k i n , P h D * t

Abstract: An examination of femoral bone-prosthesis interface behavior under different load types is undertaken using finite-element analysis. Three-dimensional finite-element models are made of two designs of hip prostheses after implantation in a femur. Femoral geometry was determined by computed tomography scans. The models were loaded in one-legged stance and stairclimbing configurations. The implants were modeled as both smooth surfaced and porous coated. The amount of contact and the relative motion between bone and implant were calculated. It is shown that torsional loads such as occur during stairdimbing contribute to larger amounts of implant micromotion than does stance loading. Contact at the bone-prosthesis interface is more dependent on load type than on implant geometry or surface coating type. Key words: total hip arthroplasty, prosthesis, stability, contact, finite-element analysis.

Initial stability of the femoral c o m p o n e n t is one of the most important factors in ensuring the longevity of cementless total hip arthroplasty (THA). Prior to b o n y ingrowth, the cementless femoral c o m p o n e n t must rely on its contact with the cortical e n d o s t e u m to achieve stability. Recent studies have s h o w n that the stability of a cementless implant is jeopardized most w h e n the patient arises from a seated position or during stairclimbing.' During these activities, a large torsional load is generated about the long axis of the prosthetic stem, which m a y produce excessive implant micromotion. Most studies have concentrated on determining the proximal femoral stress distribution and implant m i c r o m o t i o n u n d e r axial loading. Only a few investigators have evaluated the effect of torsional loads on implant stability. 2-7

Because of the complex interactions at the bone-prosthesis interface and the influence of m a n y mechanical and geometric variables, the finiteelement analytical technique is an ideal m e t h o d to study implant stability. In addition, the finite-element m e t h o d may be used in the comparative analysis of a wide variety of design modifications to determine which designs are most meritorious; however, the finite-element technique has rarely been used to investigate prosthetic stability. A better understanding of bone-prosthesis interface behavior especially under torsional loads would be beneficial in the design of n e w prosthetic components and to improve the stability of current prosthetic designs. The objective of this study is to determine the effect of joint loading and type of surface coating on prosthetic stability and cancellous bone contact. The specific aims of this investigation are to study the effects of a smooth and a porous prosthetic coating on prosthetic stability and cancellous bone contact u n d e r single-leg stance and torsional loading during stairclimbing using three-dimensional finite-element analysis. Three-dimensional finiteelement models were constructed of a single femur

From the *Department of Orthopaedic Surgery, University of Texas Medical School at Houston, and ~Department of Mechanical Engineering and Materials Science, Rice University, Houston, Texas. Reprint requests: J. Reuben, MD, Department of Orthopaedics, MSB 6.156, 6431 Fannin, Houston, TX 77030.

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implanted with two different THA femoral components. In each case, the geometry of the model was determined directly from c o m p u t e d t o m o g r a p h y (CT) scan data of the f e m u r implanted with a radiopaque replica of the prosthetic component. The f i n i t e - e l e m e n t models i n c l u d e d n o n l i n e a r interface elements to simulate the mechanical b e h a v i o r of the u n c e m e n t e d i m p l a n t at the bone-prosthesis interface. In this way, the a m o u n t of contact and the magnitude of the relative m o t i o n b e t w e e n the implant and the trabecular bone can be determined.

Literature Review Crowninshield et al. p e r f o r m e d a biomechanical investigation of the h u m a n hip during level walking, ascending and descending stairs, and rising from a chair? During stairclimbing and rising from a chair, the peak forces in the direction of progression were roughly two to three times the corresponding forces during level walking. They suggested that this force produces a torque that is transmitted to the b o n e - c e m e n t interface and m a y cause prosthetic loosening. Andriacchi et al. experimentally analyzed the motions, forces, and m o m e n t s at the major joints of the lower limbs in 10 m e n ascending and descending stairs? They f o u n d that the m e a n m a x i m u m net flexion-extension m o m e n t s at the hip were 123.9 N-m during stair ascent with a m e a n patient body weight of 71 kg. The magnitudes of these m o m e n t s were considerably higher t h a n those produced during level walking. They suggested that the forces and m o m e n t s g e n e r a t e d during stairclimbing should be considered w h e n establishing design criteria for prosthetic devices. Phillips and Messieh analyzed the clinical results of cementless hip arthroplasty using a Moore type prosthesis in 41 patients. ~ Most patients felt pain during the first few steps after rising from a sitting position and more severe pain w h e n climbing stairs. Clinically, stem loosening was most clearly demonstrated by rotational loading. Hodge et al. m e a s u r e d the pressure distribution at the hip joint using an instrumented Moore type prosthesis. ~° The highest pressure was recorded while the patient was rising from a chair and it occurred in the superior and posterior aspects of the acetabulum. Sugiyama et al. experimentally investigated the role played by torsional loads in loosening of cementless femoral c o m p o n e n t s and in cemented components prepared by three different cementing

techniques. 1~ They found that cemented components had better stability t h a n cementless components. Even the best cement technique is prone to failure in torsion u n d e r loads associated with normal activities of daily living. Schneider et al. developed an in vitro dynamic m e t h o d for m e a s u r e m e n t of the relative m o t i o n of the THA femoral c o m p o n e n t . Both axial and torsional loads were applied in their model. 6 They confirmed that the use of cement markedly reduced subsidence and rotational micromotion. Certain cementless implants achieved stability comparable to that of c e m e n t e d implants depending on the direction of loading. Several researchers have used an instrumented hip prosthesis to measure joint reaction force during daily activities. Davy, Kotzar, and co-workers have m e a s u r e d the peak joint contact force during single-leg stance to be 2.1 times body weight (BW); during chair rise, 1.23 BW; and during stairclimbing, 2.6 BW. 12,~3More recently, B e r g m a n n et al. has m e a s u r e d peak resultant joint forces of 2.8 BW during slow walking. TM Phillips et al. carried out a series of experiments on femoral stems by placing the femurs horizontally and loading vertically (pointing posteriorly) to study the failure criteria and loosening load on the stems2 They determined that the m e a n loosening torque was 23 N-re. In all specimens, failure began with an internal rotation of the stem within the f e m u r and was followed by a combination of both internal rotation and posterior subsidence. Muscle forces were neglected in their experiments. Burke et al. experimentally evaluated the initial stability of cemented and cementless femoral components during simulated single-leg stance and stairclimbing. 2 They f o u n d that both types of prostheses were very stable in simulated single-leg stance. In simulated stairclimbing, however, the c e m e n t e d components were m u c h more stable than the cementless components. Walker et al. 7 and Hua and Walker 4 measured the micromotions of various femoral stems under axial and torsional loading. Among four different types of implants, Walker et al. found axial micromotion to range from 5.3 to 15 g m under a vertical load. Total resultant micromotion in the transverse plane ranged from 42 to 57 g m for the same implants under a vertical load with anteroposterior bending. Hua and Walker measured the axial relative motion at a location 2 cm distal to the lesser trochanter as ranging from 0 to 32 p m after cyclical loading for various implants. Neither study concentrated on the differences in the responses to the different load cases.

Femoral Stem Stability

Materials and Methods The right f e m u r of a 55-year-old female cadaver was harvested and stripped of all soft tissue. To study the relative m o t i o n and b o n e - i m p l a n t contact patterns of two different implants in a single bone, two fiberglass replicas of the f e m u r were made from a cast of the original bone using the technique described by McKellop et al. 15 To make the mold, the f e m u r was sectioned along the mediolateral plane to produce an anterior and posterior half. The cancellous bone in each half was r e m o v e d d o w n to the e n d o s t e u m and the resulting cortical shell was cast in Silastic latex rubber, producing a mold for the synthetic femurs. The two halves were cast'in barium-impregnated fiberglass and b o n d e d together to produce a f e m u r replica. The inner canal ot the fiberglass f e m u r was filled with a p o l y u r e t h a n e foam to represent cancellous bone. Using standard surgical technique, an experienced surgeon implanted an appropriate-sized prosthesis into each of the acrylic femurs. Two different implants were used: the first (model A) consisted of a Mallory-Head prosthesis (Biomet, Warsaw, IN) a n d the s e c o n d (model B) consisted of a Harris-Galante prosthesis (Zimmer, Warsaw, IN). An acrylic replica of each prosthesis was t h e n cast from barium-impregnated fiberglass resin and reinserted into the respective f e m u r model. The fiberglass f e m u r - i m p l a n t replicas were t h e n CT scanned in a General Electric 9800 CT scanner. Images were obtained at 3 - m m intervals from the femoral head d o w n to the level of the lesser trochanter followed by 5 - m m intervals, covering at least the proximal 18 cm of the femur. The CT data were t h e n used to create f i n i t e - e l e m e n t models of the two f e m u r - i m p l a n t composites.

Geometric Model For each CT slice, the endosteal and periosteal contours of the cortical bone and the outer b o u n d ary of the femoral stem were extracted using contouring software. I6 The contouring routine used a stepwise series of density threshold and gradient criteria to produce the optimal contours. These contours were used to construct the geometric and finite-element models of the femur and implant using the computer-aided engineering software PATRAN (PDA Engineering, Costa Mesa, CA). It was assumed t h a t regions b e t w e e n the stem and the cortical bone were entirely filled with trabecular bone. The most distal CT slice taken was copied and translated 75 m m distally to represent the femoral diaphysis.

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Finite-element Mesh The finite-element models were constructed with the cortical bone, trabecular bone, and the implant represented by 27-node quadratic hexahedral elements. The greater trochanter was modeled as entirely composed of trabecular bone. Quadratic interpolation interface elements were generated b e t w e e n the trabecular bone and the implant to model the nonlinear interface conditions existing b e t w e e n the bone and the implant stem. The interface elements consist of nine pairs of nodes; in each pair, one node is connected to the trabecular bone, and the other is connected to the implant. The interface elements match with the 27-noded, 9-node-per-face solid elements of the trabecular bone and the implant so that w h e n a tensile n o r m a l force is present in a pair of interface nodes, separation of the nodes occurs, indicating a lack of contact b e t w e e n the trabecular bone and the implant stem. In these cases, the tensile force is not transmitted b e t w e e n the two bodies. W h e n a compressive normal force is present across a pair of interface nodes, contact occurs and compressive forces are transmitted to the trabecular bone. A friction coefficient can be assigned to the interface element to simulate the frictional effect in the interface due to different prosthetic surfaces, such as porous coatings and smooth surface treatments. The areas of contact b e t w e e n the stem and the cancellous bone can be calculated after loading by determining w h e r e separation has and has not occurred within the interface. In both models, the interface elements covered the whole length of the stem and a perfect initial fit was assumed b e t w e e n the cancellous bone and the prosthesis. The degree of r e f i n e m e n t of the finite-element meshes was chosen so that all elements were undistorted and so that versions of the models in which the stem was considered to be fully b o n d e d to the bone passed a convergence test. The convergence test consisted of creating successive versions of the model with increasing geometric refinement. The models were considered sufficiently refined w h e n displacements at nodes on the medial surface of the cortical bone differed by less than 1% from the displacements found with a m o r e refined model. The finite-element model of the implanted MalloryHead prosthesis (model A) contains 15,000 nodes, with 1,344 27-node elements representing the bone and the implant and 328 interface elements. The finite-element model of the femur with the H a r r i s - G a l a n t e prosthesis (model B) contains 22,035 nodes with 2,007 solid elements represent-

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ing the b o n e and the i m p l a n t a n d 450 interface elem e n t s (Fig. 1).

Loads, Boundary Conditions, and Material Properties Two loading conditions were simulated in this study. The first case represented single-leg stance, the second case simulated stairclimbing. The loading condition representing single-leg stance was calculated f r o m the equilibrium equations with three forces in the midfrontal plane of the f e m u r using the simplified free b o d y t e c h n i q u e ? 7 The three forces consisted of the g r o u n d reaction force (5/6 BW), the abductor muscle force (1.754 BW inclined medially by 20 ° f r o m verticaPS), and the hip joint reaction force (2.1 BW, 16.6 ° f r o m verticaP2). Assuming a b o d y weight of 71 kg, the x, y, and z c o m p o n e n t s of the joint reaction force (F1) are (-417.87, 0, -1,401.71 N), w h e r e the z axis is aligned w i t h the longitudinal axis of the f e m u r and the x and g axes are parallel to the mediolateral and anteroposterior axes, respectively. The c o m p o n e n t s of the force g e n e r a t e d by the abductor muscle (F2) are (417.84, 0, 1,148 N). To simulate the stairclimbing load case, the magnitude of the c o m p o n e n t s of the joint contact force and m o m e n t w e r e adopted f r o m published experim e n t a l data. As the force direction and m a g n i t u d e of the individual abductor muscles are not k n o w n during stairclimbing, idealized resultant muscle forces are a s s u m e d to act on the f e m u r ? 9 F r o m

/FI

F3

gIy

Fig. 1. (A) Model A loaded in one-leg stance. (B) Model B loaded in stairclimbing.

Andriacchi et al., the p e a k flexion m o m e n t generated by a 71-kg m a n during stairclimbing is 123.9 N - m in the sagittal p l a n e a n d the a d d u c t i o n m o m e n t in the frontal plane is 25 N-re. 9 There is no m o m e n t in the transverse plane. F r o m Kotzar et al., the p e a k joint contact force during stairdimbing is 2.6 g w . ~3The load c o m p o n e n t s of the force applied at the h e a d of the f e m u r (F3) w i t h respect to the femoral coordinate system are t h e n (-709.2, 600.8, -1,553.2 N), w i t h applied m o m e n t s of M x = -123.9 N-mandM g =25N-m. The coordinate system used in the finite-element models is the same as the CT scan reference system, with the origin located at the CT scan origin of the m o s t distal slice. As the cadaver femurs w e r e placed into the scanner at an angle to the scanner axis system, the a n a t o m i c coordinate system of each f e m u r h a d to be identified a n d related to the model's global coordinate system so that the loading conditions could be applied properly. The femoral coordinate system was defined in the following m a n n e r : the longitudinal axis of the f e m u r was created by passing a best-fit line t h r o u g h the centroids of all the cross-sections b e t w e e n the level of the inferior border of the lesser t r o c h a n t e r a n d the isthmus (section with smallest cross-sectional area). A coronal plane was defined as that plane that contains the center of the femoral h e a d and the longitudinal axis. The anteroposterior axis was formed by passing a line perpendicular to the coronal plane and passing anteriorly. Once defined, loads described above were transformed f r o m this coordinate system to the coordinate system of the finite-element models. The nodes on the m o s t distal cross-section of the m o d e l w e r e constrained in all three directions. As a simplifying assumption, cortical b o n e and trabecular b o n e w e r e assumed to be separate, h o m o g e n o u s , isotropic materials with constant Young's moduli of 14 and 0.5 GPa, respectively. The titanium implant was assigned a Young's m o d u l u s of 110 GPa. 2° Poisson's ratio was t a k e n to be 0.3 for all materials. Two kinds of prosthetic surface finishes w e r e simulated in this study: a s m o o t h surface implant and identically shaped p o r o u s - c o a t e d implant. In each case, it was a s s u m e d that the implant h a d the same surface t r e a t m e n t along the entire length of the stem. The surface t r e a t m e n t was included in the finite-element models by assigning a coefficient of friction to the elements representing the interface b e t w e e n the stem and the trabecular bone. The values used for the friction coefficients of s m o o t h and p o r o u s - c o a t e d implant surfaces in contact with b o n e are 0.42 and 0.61, respectively. 2' The finitee l e m e n t models w e r e solved using ABAQUS (tIibbitt, Karlsson a n d Sorensen, Pawtucket, RI) on a Silicon Graphic (Mountain View, CA) workstation.

Femoral Stem Stability

Cancellous Bone-Prosthesis Contact As the bone-prosthesis interface will separate at the nodes in which net tensile n o r m a l stress occurs, cancellous bone contact with the prosthetic stem can be determined by analyzing the n u m b e r of nodes for each interface element that remains in contact following loading. Contact was defined as the existence of a compressive n o r m a l stress b e t w e e n the paired nodes. In cases of contact, the implant and the trabecular b o n e were said to be stuck at the interface if the resultant shear stress perpendicular to the interface n o r m a l was less t h a n the critical shear stress (coefficient of friction multiplied by the n o r m a l compressive force). The two bodies slid with respect to each other w h e n the resultant shear stress and the critical shear stress were equal. Contact within a given element was defined using the m e t h o d set forth by K e a v e n y and Barrel22; that is, complete contact was assumed to occur w h e n seven or m o r e of the nine pairs of nodes in one interface element were in contact after loading; transitional contact was defined as four to six pairs in contact; a n d separation was defined as less t h a n four pairs in contact.

Relative Motion The relative m o t i o n s b e t w e e n the prosthesis a n d b o n e w e r e calculated at four locations on the medial side of the implant: the stem tip, the lower one quarter, the u p p e r one quarter, and the calcar. These four locations are illustrated in Figure 2.

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Results Regions of Contact Figures 3 to 6 show the regions of contact for the s m o o t h a n d p o r o u s - c o a t e d models of the two implants subjected to the single-leg stance and the stairclimbing loading conditions. The black regions represent full contact, the dark gray regions represent transitional contact, and the light gray regions represent separation. In comparing the s m o o t h with the p o r o u s - c o a t e d stems, the regions of contact are similar regardless of surface finish for the same implant a n d loading condition. There are significant differences, h o w ever, b e t w e e n the two different loading conditions. In stance, b o t h implants s h o w contact occurring laterally n e a r the middle of the stem, as well as on the medial side closer to the p r o x i m a l a n d distal regions of the stem. Anterior and posterior contact is p a t c h y in one-leg stance, and m o r e likely to occur in the m i d s t e m t h a n in the p r o x i m a l regions (Figs. 3, 4). In the stairclimbing load case, b o t h implants s h o w contact primarily in the middle of the stem on the anterior side, w i t h some posterior p r o x i m a l and distal contact. Contact on the lateral side occurred in the mid to distal regions of the stem, a n d occurs to a greater extent t h a n does contact on the medial side. The regions of contact in the stairclimbing load case are less t h a n those in the single-leg stance load

! L

B POSTERIOR

Fig. 2. The prostheses of model A (left) and model B (right). The labeled points indicate the locations where relative motion was calculated.

LATERAL

ANTERIOR

MEDIAL

Fig. 3. Regions of bone-implant contact in one-legged stance for model A (top) and model B (bottom). Both implants are smooth surfaced.

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A

A

[] POSTERIOR

[] LATERAL

ANTERIOR

MEDIAL

POSTERIOR

LATERAL

ANTERIOR

MEDIAL

Fig. 4. Regions of bone-implant contact in one leg stance for model A (top) and model B (bottom). Both implants are porous coated.

Fig. 6. Regions of contact in stairclimbing for model A (top) and model B (bottom). Both implants are porous coated.

case, especially o n the posterior and medial surfaces (Figs. 5, 6). Comparing the two implants, slightly m o r e contact is achieved with model B than with model A in both the single-leg stance and the stairclimbing

load cases. The two stems, however, demonstrate similar patterns of contact for a given load case.

B POSTERIOR

LATERAL

ANTERIOR

MEDIAL

Fig. 5. Regions of contact in stairclimbing for model A (top) and model B (bottom). Both implants are smooth surfaced.

Relative Motions

The relative m o t i o n b e t w e e n the implant and the trabecular bone at the four sample nodes u n d e r the two loading cases are s h o w n in Figures 7 and 8, respectively. Axial relative m o t i o n refers to m o t i o n in the direction parallel to the longitudinal axis of the femur. Transverse m o t i o n refers to m o t i o n within the cross-sectional plane perpendicular to the longitudinal axis. The total m o t i o n is the magnitude of the sum of the axial and transverse m o t i o n vectors. In one-leg stance, high axial and transverse relative motions occur distally. The relative m o t i o n in the axial direction tends to be greater t h a n the m o t i o n in the transverse direction. The greatest transverse relative m o t i o n is 18 ~m; the greatest axial relative m o t i o n is calculated to be 32 btm. Higher values of relative m o t i o n occur in stairclimbing t h a n in one-leg stance loading. In stairclimbing, the highest relative m o t i o n calculated is 58 btm and is seen proximally in the transverse direction; the greatest axial m o t i o n is 26 btm and occurs at the stem tip. The total relative motions were below 45 btm for all cases except in the proximal region of model A in stairclimbing, w h e r e the relative m o t i o n was 58 g m for the s m o o t h stem and 56 g m for the porous-

Femoral Stem Stability

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coated stem. This is due to the large transverse relative motions seen u n d e r the torsional loads generated during stairclimbing. •

MODELASMOOTH

[ ] MODELAPOROUS

~

MODELBSMOOTH

[ ] MODELBPOROUS

Discussion

D1STALTIP(D) LOWERI/4(C) UPPERI/4 (B) TOP(A)

DISTALT1P(D) LOWERI/4(C) UPPERI/4 (B) TOP(A)

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DISTALTIP(D) LOWER1/4(C) UPPERI/4 (B) TOP(A)

Fig. 7. Relative motion for all stems in one-leg stance.

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MODI~LA SMOOTH

[]

MODELA POROUS

[]

MODELBSMOOTH

[]

MODELBPOROUS

DISTALTIP(D) LOWERI/4(C) UPPERI/4 (B) TOP(A) q

so

DISTALTIP(D) LOWEHI/4(C) UPPERI/4 (B) TOP(A)

~o

DISTAL'rIP(D) LOWERI/4(C) UPPERII4(B) TOP(A)

Fig. 8. Relative motion for all stems in stairclimbing.

The complex three-dimensional g e o m e t r y and material properties of the proximal femur make it difficult to analyze the mechanical behavior of THA devices. Experimental studies involving the comparative analysis of a wide variety of design parameters are time consuming and difficult to perform. On the other hand, analytical studies can be performed easily but suffer from the difficulty in modeling all the potential biologic factors. With the use of CT scan data, the geometry of the bone and implant can be accurately d e t e r m i n e d and incorporated into a three-dimensional finite-element analysis. Such an analysis is t h e n helpful in assessing a wide variety of design possibilities and determining which designs are most promising for further experimental analysis or clinical use. W h e n using finite-element analysis, it is imperative that the model be verified. This can be achieved by performing displacement convergence tests on the m e s h and comparing the results of the finiteelement analysis with previous studies. The convergence tests conducted have s h o w n the finiteelement meshes used to be sufficiently refined. The results of the relative m o t i o n study showed that total m o t i o n is greatest in the proximal and distal regions of the stems and least in the midstem. The total m o t i o n values range from a m a x i m u m of 58 btm to less than 1 btm u n d e r the applied loading conditions. The m o t i o n was slightly greater in the smooth stem than the porous stem. The most severe loading condition in terms of relative m o t i o n was stairclimbing. These data agree closely with other experimental studies, 3,6,',23,24 which serves to validate the accuracy of the finite-element model. The major difference between the relative motions of the two implants modeled is that model B consistently shows higher axial and transverse relative motion than model A at the distal tip of the stem. Model A experiences m u c h more transverse relative motion at the proximal node in stairclimbing than does model B. Otherwise, the relative motions calculated for the two implants are fairly similar. Certain trends can be seen for both f e m u r implant models. With only one exception, the relative motions for the s m o o t h stem and the porouscoated stem are within 5 btm of each other. The average difference in total relative m o t i o n at corresponding points on smooth and porous-coated stems is 2 btm. Additionally, the middle nodes g e n -

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The Journal of Arthroplasty Vol. 10 No. 6 December 1995

erally e x p e r i e n c e smaller a m o u n t s of relative m o t i o n t h a n do the m o r e p r o x i m a l a n d distal nodes. In one-leg stance, the average total m o t i o n for all stems at point C is I8 ~ m lower t h a n the average total m o t i o n at the stem tip. In stairclimbing, the difference b e t w e e n the two is 14 ~m. For a particular loading situation, the patterns of contact w e r e similar for the s m o o t h and porouscoated versions of an implant, indicating that the region of contact is not d e p e n d e n t on the type of surface coating. F u r t h e r m o r e , the two different geometries of i m p l a n t e x a m i n e d in this study s h o w e d similar patterns of contact for the respective load cases. This suggests that the regions of contact b e t w e e n the implant and trabecular b o n e in one-leg stance or stairclimbing conditions largely reflect the loading rather t h a n the shape of the individual stem or the type of surface coating. For the stairclimbing loading situation, the joint contact force c o m p o n e n t in the posterior direction was 601 N. This c o m p o n e n t generates a 2 1 - N - m torque in the transverse plane a b o u t the long axis of the implant, w h i c h approaches the 2 3 - N - m torque that resulted in loosening in the study perf o r m e d b y Phillips et al. ~ The large regions of separation in the interface u n d e r stairclimbing loads also support the observations f r o m other experim e n t a l studies that torsional loading is m o r e deleterious to the stability of the implant. ~,2.5,~-u The deleterious effects of torsional loading are further d e m o n s t r a t e d in the relative m o t i o n study. Although the relative motions are greater for the stairclimbing loading situation t h e n the stance loading, t h e y are still in the threshold range for allowing i n g r o w t h of b o n e into the p o r o u s - c o a t e d surface25; however, the m a x i m u m total relative m o t i o n (58 ~m) that occurs at the m o s t p r o x i m a l node u n d e r the stairclimbing loads m a y limit the extent of b o n y ingrowth. This result corroborates the findings of other authors using e x p e r i m e n t a l m e t h o d s that physiologic torsional loads should be applied to d e m o n s t r a t e differences in m o t i o n at the b o n e - i m p l a n t interface a n d to produce m o t i o n that m a y be clinically i m p o r t a n t Y In addition, the findings suggest that torsional loading of cementless p o r o u s - c o a t e d devices should be avoided in the early postoperative period, because relative m o t i o n that occurs at the b o n e - i m p l a n t interface u n d e r high physiologic loads such as during stairclimbing m a y preclude i n g r o w t h of bone. It is probably best to avoid stairclimbing until b o n y i n g r o w t h or prosthetic stabilization is suspected clinically. This also suggests that it is i m p o r t a n t to consider the torsional stability of a device if it is to be used in a cementless application. It m a y be beneficial to test

the m o t i o n at the interface or selectively screen prosthetic implants in the design stage using a finite-element m e t h o d as described in this study. The prostheses used in this study did not h a v e substantial collars. The use of a collar can decrease axial m o t i o n but m a y not affect the transverse comp o n e n t of the interface m o t i o n due to torsional loading. 23,2~ The i m p l a n t s w e r e m o d e l e d as if entirely coated with a porous i n g r o w t h surface; this is not the case w i t h c o n t e m p o r a r y implants. We h a v e shown, however, that the type of surface coating has very little effect on the regions of contact b e t w e e n the b o n e and the implant. The two implants studied illustrate the deleterious effects of torsion on cementless implants. Although the precise a m o u n t s of contact a n d relative m o t i o n b e t w e e n b o n e and i m p l a n t differ with prosthetic and femoral geometry, the prostheses m o d e l e d are representative of the effects of one-leg stance and stairclimbing load cases.

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Femoral Stem Stability 12. Davy DT, Kotzar GM, Brown RH: Telemetric force measurements across the hip after total arthroplasty. J Bone Joint Surg 70A:45, 1988 13. Kotzar GM, Davy DT, Goldberg VM et ah Telemeterized in vivo hip joint force data: a report on two patients after total hip surgery. J Orthop Res 9:621, 1991 14. Bergmann G, Graichen E Rohlmann A: Hip joint loading during walking and running measured in two patients. J Biomech 26:969, 1993 15. McKellop H, Ebramzadeh E, Niederer PG, Sarmiento A: Comparison of the stability of press-fit femoral stems using a synthetic model femur. J Orthop Res 9:297, 1991 16. Chang CH: Computer aided design of femoral stem prostheses. PhD dissertation, D e p a r t m e n t of Mechanical Engineering and Materials Science, Rice University, Houston, TX, 1990 17. Frankel VII, Nordin M: Basic biomechanics of the skeletal system. Lea & Febiger, Philadelphia, 1980 18. I n m a n VT: Functional aspects of the abductor muscles of the hip. J Bone Joint Surg 29A:607, 1947

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