Adult Hip, The
2nd Edition

5
Biomechanics of the Hip
James D. Johnston
Philip C. Noble
Debra E. Hurwitz
Thomas P. Andriacchi
An understanding of the biomechanics of the hip is vital to advancing the diagnosis and treatment of many pathologic conditions. Some areas that have benefited from advances in hip biomechanics include the evaluation of joint function, the development of therapeutic programs for treatment of joint problems, procedures for planning reconstructive surgeries, and the design and development of total hip prostheses. Biomechanical principles also provide a valuable perspective to our understanding of the mechanism of injury to the hip and the contributions of the capsule, labrum, and femoroacetabular impingement (FAI) to the etiology of degenerative hip disease.
The biomechanics of hip function may be described through reference to the kinematics or the kinetics of the hip joint or its prosthetic replacement. Joint kinematics is the description of the angular or translational motion of the joint in response to applied forces; kinetics refers to the forces and moments acting on the joint during motion, whether they arise from muscle activity, inertia, ligamentous restraints, or contact between the femur and pelvis and adjacent structures. Different approaches have been applied to study hip biomechanics. The kinematics of the joint may be quantified using motion analysis, especially in conjunction with analytical models of the musculoskeletal system. On the other hand, joint forces may be estimated from data derived from gait and force platform measurements, in combination with analytical models simulating the force of contraction and line of action of each of the hip muscles.
The Biomechanics of Injuries to the Hip Joint
The Mechanical Role of the Soft Tissues of the Hip
Over the last decade, there has been increasing interest in the diagnosis and treatment of acute injuries of the hip joint, particularly those involving the labrum and the articular surfaces. This interest has arisen from the recognition that many events in the pathomechanics of degenerative joint disease occur early, and that soft tissue disruption is probably involved more frequently than had been recognized previously. In addition, advances in hip arthroscopy have allowed clinicians to examine the articular surfaces of the hip joint in patients with debilitating symptoms despite a normal radiographic appearance. In these patients, a frequent finding has been labral pathology, most commonly chondrolabral separation in the anterior aspect of the joint, with a disturbing prevalence of full-thickness articular lesions, often in communication with chondrolabral defects (53).
Given the significance of periarticular structures in the pathomechanics of hip disease, it is instructive to review the basic anatomy of the passive stabilizers of the hip joint, including the capsular ligaments and the acetabular labrum. The hip capsule (capsular ligament) is critical to the stability and proper function of the hip joint (78) and serves as a “check rein” preventing dislocation at the extremes of motion (38,46). This ligament is actually a complex structure consisting of three discrete ligaments: (a) the anteriorly located iliofemoral ligament, which restricts extension of the joint and limits internal rotation, (b) the femoral arcuate ligament, which limits abduction and external rotation and is also anteriorly located, and (c) the posteriorly located ischiofemoral ligament, which limits internal rotation and adduction when the hip is flexed (27,46). Mechanical testing of these
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ligaments has shown that the posterior ligament is much weaker than its anterior counterparts (37,78), which may explain why posterior dislocation occurs so much more frequently than anterior dislocation (62).
The acetabular labrum is a relatively stiff fibrocartilaginous tissue that forms an extension of the acetabular rim (25). The labrum increases the effective depth of the socket and the coverage of the femoral head and, so, has the potential to enhance the mechanical stability of the hip joint. Although the load-bearing function of the labrum was questioned by Konrath et al. (49), recent experimental studies have demonstrated that, due to its low permeability, the labrum acts as a seal, preventing fluid from flowing in and out of the intra-articular space (24,80). This increases the stability of the joint, as the seal must be broken to dislodge the femoral head from the acetabulum. Moreover the presence of the labrum elevates the intra-articular pressure during weight bearing, which is expected to enhance joint lubrication and minimize friction (23,24). Additionally, the removal of the labrum has been shown to increase strains within the cartilage matrix during weight bearing (23,24).
Clinical evidence has shown that labrum removal or pathology is closely linked to joint changes consistent with early osteoarthritis and joint disease (21,35,52,53,54), with the relative risk of significant cartilage erosion doubling in the presence of labral tears (53). It is hypothesized that a labral tear (occurring predominately anteriorly) disrupts the stability of the hip, especially at the extremes of joint motion. This leads to abnormal sliding of the articular surfaces under the dynamic torsional conditions often associated with sporting activities involving strenuous repetitive twisting and pivoting motions (such as ballet, football, soccer, basketball, and placekicking) (53). These motions accelerate degenerative changes and the progression of chondral involvement with time, leading to joint disease (44,53).
The Role of Mechanical Factors in the Etiology of Coxarthrosis
Osteoarthrosis is a disease involving the symptomatic loss of articular cartilage in a normal load-bearing area of a joint and is frequently associated with subchondral sclerosis and osteophyte formation (67). It is generally thought that the primary etiology is mechanical. Different authors have implicated excessive impulsive loading of the articular surface, leading to acute cartilage injury (64,68) and/or accumulated microtrauma of cartilage and subchondral bone in response to repetitive impulse loads (68). Articular cartilage is a viscoelastic material; its interstitial fluid component contributes to its bulk mechanical properties. When cartilage is loaded rapidly, the fluid does not have time to flow, resulting in increased tissue stiffness and high internal stresses (68). As the mechanical response of articular cartilage is mediated by both the permeability and the elastic modulus of the tissue, the pathologic response of joints to load bearing is highly dependent upon loading rate, in addition to load magnitude (68).
Previous experimental studies have demonstrated that osteoarthritic changes in weight-bearing joints occur in direct response to repetitive loading above threshold levels (69). An interesting study by Hadley et al. (33) showed that, provided that mean pressures at the articular surface are kept below 2 MPa, articular cartilage is capable of tolerating repetitive loads almost indefinitely. Once pressures rose above 2 MPa, due to reduced or incomplete head coverage, degenerative changes were observed, depending upon the duration of exposure. Unsatisfactory outcomes were observed in 90% of the patients exposed to >10 MPa-yrs of articular loading, while in 81% of hips experiencing <10 MPa-yrs, satisfactory outcomes were observed.
Although joint loading or overloading can lead to cartilage degeneration, the precise mechanism remains controversial. Opinions differ concerning the relative contributions of direct mechanical trauma to articular cartilage versus elevation of articular stresses secondary to stiffening of the subchondral plate. Certainly, finite element studies have shown that localized increases in subchondral stiffness can lead to marked elevations of stresses in the overlying cartilage (12). However, there are no studies that allow us to directly separate the contributions of hydrostatic (i.e., direct compressive) and deviatoric (i.e., shearing) stresses generated by joint loading. Consequently, it is not known whether degenerative changes are primarily due to excessive loads applied during normal activities, or shearing of the cartilage during abnormal motions involving local instability. It is known that individuals performing heavy lifting or participating in elite sporting activities are at greater risk of osteoarthritis (39). It is hypothesized that acute injuries from trauma experienced during athletic activities, in addition to the strenuous twisting and pivoting motions, leads to recurrent microtrauma and eventual cartilage deterioration.
Femoro-Acetabular Impingement (FAI)
Clinical implications of Femoro-Acetabular Impingement (FAI) include acute hip pain (22,48), loss of motion (66), and chronic leveraging of the head in the acetabulum, possibly resulting in labral, cartilage, and chondral injury (51,66). Repeated contact between the femoral head–neck junction and the labrum can lead to progressive cartilage damage and labral tearing (22,45), thus influencing mechanical stability and the weight-bearing role of the joint (24,53). This in turn would accelerate degenerative changes and possibly lead to joint diseases such as osteoarthritis (44,54). The pathomechanics of FAI have been elucidated by Ganz and coworkers in many published studies. Impingement between the femoral neck and the acetabular rim results from reduced joint clearance (30,45,61) and possibly from forced articulation beyond the mechanical limits of joint motion (54). Joint clearance is compromised in the presence of abnormal morphologies of the acetabulum and/or proximal femur, including widening of the femoral neck, reduction in the head–neck offset, and overcoverage of the acetabulum, all of which reduce joint clearance, increasing the probability of impingement (30,45,58,61,66,84). Additional proximal femoral morphologies that potentiate impingement include abnormal head/neck configurations (e.g., a pistol-grip deformity) (57,66,75,81), reduction in femoral anteversion (45,81), reduced concavity at the femoral head–neck junction (61), subclinical displacement of the femoral epiphysis (head tilt or post slip) (31,45,57,66,75), malpositioned proximal fragments after neck fracture (22), and residual
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effects from childhood diseases such as Legg–Calvé–Perthes (30). On the acetabular side, the risk of impingement increases with acetabular retroversion, which results in a prominent anterolateral overcoverage, creating an obstacle for flexion and internal rotation (30). Other predisposing conditions are deepening of the acetabular socket (coxa profunda and protrusio acetabuli) and some posttraumatic deformities (30,48,73).
Two distinctive types of FAI have been recognized (Fig. 5-1). The first type is due to linear contact between the acetabular rim and the femoral head–neck junction, thus limiting range of motion (66). The first structure to fail is often the labrum, with continued impact resulting in deterioration of the labrum and ganglion formation, or ossification of the rim leading to additional acetabular deepening and over coverage. This type of impingement is seen more frequently in middle-aged women who participate in athletic activities involving hip motion (30,51). The second type or “cam” FAI is due to an abnormal femoral head with increasing radius in the anterolateral region (22,45,61). Motion of the abnormal head in the acetabulum results in high shear forces, producing abrasion of the acetabular cartilage and/or its avulsion from the labrum and the subchondral bone, often leading to tearing or detachment of the labrum. This type of impingement is most common in young and athletic males (30,51).
Figure 5-1 Mechanism of impingement caused by (A) direct linear contact between the labrum–acetabular rim and the head–neck junction, or (B) an increased head radius abutting against the acetabular rim (cam). (From
Beck M, Leunig M, Parvizi J, et al. Anterior femoroacetabular impingement: part II. Midterm results of surgical treatment. Clin Orthop. 2004;418:67–73
.)
Forces Transmitted by the Hip Joint
In Vivo Measurements of Forces at the Hip Joint
The long-term integrity of both the normal and prosthetic hip is strongly influenced by the direction and magnitude of the force developed between the femoral head and the acetabulum during daily activities. As there is no method of directly measuring forces acting on the normal hip joint, the only in vivo measurements of forces or pressures at the joint surface have been provide by prostheses and endoprostheses instrumented with transducers (strain gauges) (Table 5-1). The earliest experiments utilizing a strain-gauged prosthesis were performed by Rydell, who measured peak forces of 3 body weights (BW) during gait (70). More extensive studies have been recently conducted by Bergmann et al. (3,4,5,6), who recorded peak forces that varied from 2.1 to 4.3 BW during gait (3,5,6), 2.3 to 5.5 BW during stair-climbing (3,6), to greater than 8 BW during accidental incidents of stumbling (4,5).
The magnitude of the peak resultant force generated during gait can vary from 1.6 to 4.3 BW (Table 5-1) and is affected by numerous factors, especially stride length and the speed of ambulation. During the stance phase of gait, the orientation of the load in the frontal plane is relatively constant and directed medially and inferiorly, while in the sagittal plane the orientation is more variable and is toward the posterior during the first part of the stance phase and anterior during the later part of the stance phase (Fig. 5-2). Analysis of data from several investigators indicates that the lateral, posterior, and inferior peak components during gait range from 0.4 to 1.7 BW, 0.2 to 1.0 BW, and 1.4 to 4.1 BW (3,5,6,16,70). In some cases, the determination of these force components was based on estimated anteversion angles and varus/valgus stem angles (4).
The magnitudes of the out-of-plane loads during daily activities can be substantial, with the anterior–posterior component of the hip reaction force reaching 20% to 25% of the force in the frontal plane during stair-climbing (6). In vivo force transducer data demonstrate that, during climbing up stairs, the torques generated about the implant longitudinal axis are 23% greater than in normal walking. Conversely, the axial torques recorded during walking and descending stairs were of similar magnitude (3,6). Moreover, the peak contact force and resulting torsional moment, and the posteriorly directed force component increase with walking speed. Out-of-plane loads, and the torques they generate about the femoral axis, are also affected by the anteversion angle (5). The largest torsional moments measured in vivo during activities of daily living reach the average experimental strength of implant fixation (33.1 N m), as determined from in vitro tests (65) (Fig. 5-3). Consequently, the moments generated during stair-climbing may be detrimental to the stability of implant fixation, especially in uncemented stems (8,28,34,65).
Contact pressures from a Moore-type endoprosthesis have been studied during numerous activities including walking, jogging, stair-climbing, and chair-rising (40,41,50,79). Peak pressures during gait occur between heel strike and early midstance and relate to increases in both ground reaction forces and
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abductor muscle activity. During gait, the maximum pressures (mean: 5.6 MPa) occur on the superior anterior femoral surface, which corresponds to the superior acetabular dome (79). Chair-rising triples pressures up to 9 to 15 MPa on the apex of the femoral head or superior posterior aspects of the acetabulum. The sites of high acetabular pressure on the superior posterior region of the acetabulum correspond to sites of frequent degenerative changes observed in cadaver specimens (40).
TABLE 5-1 HIP CONTACT FORCES MEASURED IN VIVO IN PATIENTS WITH INSTRUMENTED IMPLANTS
Activity Typical Peak Force (BW) Total Number of Patients Time Since Surgery (Months) References
Walking, slow 1.6–4.1 9 1–30 (3,5,6,70)
Walking, normal 2.1–3.3 6 1–31 (3)
Walking, fast 1.8–4.3 7 2–30 (3,5,6,70)
Jogging/running 4.3–5.0 2 6–30 (5,6)
Ascending stairs 1.5–5.5 8 6–33 (3,6,70)
Descending stairs 1.6–5.1 7 6–30 (3,6,70)
Standing up 1.8–2.2 4 11–31 (3)
Sitting down 1.5–2.0 4 11–31 (3)
Standing/2-1-2 legs 2.2–3.7 3 11–14 (3)
Knee bend 1.2–1.8 3 11–14 (3)
Stumbling 7.2–8.7 2 4–18 (4,5)
Figure 5-2 Force vectors and points of load transfer on the prosthetic head during level walking for a patient. Positive x is directed medially, positive y is in the direction of progression (anterior), and positive z is superior. (From
Bergmann G, ed. HIP98: Loading of the Hip Joint. Berlin: Free University of Berlin; 2001
, with permission. Compact disc, ISBN 3980784800.) See Color Plate.
Although it is tempting to extrapolate data from instrumented prostheses to the physiology of the normal hip, some limitations of these experimental studies must be remembered. Firstly, because of their complexity, studies performed using instrumented prostheses have been limited to one or two subjects at most institutions. Early studies were frequently limited by equipment failure so that data were only collected during the early postoperative period, when the acute effect of surgical trauma may still have played a role. In some instances, assistive devices were routinely used for ambulation or subjects walked extremely slowly. Nonetheless, the consistency of peak loads measured by several different institutions seems to indicate that stems with similar features experience similar peak forces. In many instances, differences between the measured forces can be attributed to variations in the hip position during gait. For example, alterations in anteversion have a large effect on the out-of-plane loads in the transverse plane.
Analytical Estimates of Forces at the Hip Joint
Basic analytical approaches to the balance of forces and moments about the hip joint can be useful in estimating the effects of alterations in joint anatomy or different treatment modalities on the hip joint reaction force. The static loading of the hip joint has been frequently approximated with a simplified, two-dimensional analysis performed in the frontal (coronal) plane (Fig. 5-4). Several authors have used this
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approach to model the static forces during one-legged stance, both with and without a cane (13,26,59,60). In this analysis, it is assumed that the body is at rest, and that only the abductors are active. Thus, for equilibrium at the hip, the force of contraction of the abductors must generate a moment of equal magnitude, but opposite direction, to that produced by the weight of the body supported by the lower extremity, often called the effective body weight, acting on the head of the femur. During one-legged stance, the weight of the supporting extremity is distal to the hip joint and so does not contribute to the weight that must be supported. Thus the effective body weight is generally assumed to be five sixths that of the total weight of the body. The effective body weight acts vertically, in the direction of gravity, while the abductor muscle force has both a horizontal and a vertical component and is generally assumed to be oriented at 30° with respect to a vertical axis. Using this type of analysis, joint forces for one-legged stance of 2.75 BW (26) and 3.00 BW (60) have been calculated.
Figure 5-3 Torsional moments in the transverse plane: gray bar indicates fixation strength of cementless implants. (From
Phillips TW, Nguyen LT, Munro SD. Loosening of cementless femoral stems: a biomechanical analysis of immediate fixation with loading vertical, femur horizontal. J Biomech. 1991;24:37–48
, with permission.
Bergmann G, Graichen F, Rohlmann A. Is staircase walking a risk for the fixation of hip implants? J Biomech. 1995;28:535–553
, with permission.)
Figure 5-4 Joint reaction force acting across the left hip during one-legged stance. (B) Variation of the Joint Reaction Force with changes in the ratio of the lever arms, band c. (From
Greenwald AS. Biomechanics of the hip. In: Steinberg M, ed. The Hip and Its Disorders. Philadelphia: WB Saunders; 1991:49
, with permission.)
When estimates of hip mechanics are needed for dynamic activities, analytical approaches can be useful, as they are noninvasive and easily applied to a large number of patients. To calculate the forces and moments acting across the hip joint, several quantities must be measured or estimated, including the reaction forces developed between the ground and the foot (and any support devices) (Fig. 5-5), the inertial properties of moving limb segments, and the three-dimensional position of the joint centers during dynamic activities. Limb motions are measured with various types of optoelectronic methods, while the foot–ground reaction forces are measured using a force plate.
Once the net forces and moments acting across the hip have been calculated, it is possible to estimate the contributions of muscle contraction, passive soft tissue stretch, and articular reaction forces in a manner similar to that for the two-dimensional case, as described above. However, because so many muscles contribute to the joint reaction force, it is not possible to directly calculate the contribution of each of these components. Numerous authors have proposed methods of predicting the force of contraction of each muscle, based upon
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electromyographic measurements and mechanical models of each muscle. Others have predicted the distribution of contraction force between muscles corresponding to the minimum value of force per cross-sectional area of each muscle, and other optimization criteria. These methods have been applied to estimate muscle forces and contact forces on the basis of externally measured forces during the activities of gait, stair-climbing, and chair-rising (Table 5-2) (10,15,63,72). A shortcoming of all of these methods is the fact that they are unable to account for the effect of cocontraction of muscles acting on opposite sides of the joint or the effect of capsular stiffness. These factors can be significant at the extremes of joint motion and in activities where the hip capsule contributed to joint stability. Studies involving both analytical estimates and in vivo load measurements have shown promising comparisons: Heller et al. (36) showed mean peak force differences of 12% during walking and 14% during stair-climbing, and Stansfield et al. (76) demonstrated load differences of approximately 16% during activities such as walking and sit-to-stand. A recent parametric model (Hurwitz et al., 2003) estimates the potential range in contact forces resulting from physiologically feasible muscle force distributions and thus allows for muscle co-contraction while not requiring an optimization criterion. For a representative subject, the peak contact force for a reprepresentative subject varied from 2.7 to 3.2 Body weights in the absence of antagonistic activity and increased approximately 0.2 Body weights for every 10% increase in antagonistic activity.
TABLE 5-2 ANALYTICAL METHODS OF ESTIMATING PEAK HIP CONTACT FORCE
Activity Magnitude (BW) Method Reference
Statically determinate methods
Walking 4.2 after heel strike
4.8 before toe off
Reduction method (63)
Stair-ascending 7.2 Reduction method (63)
Stair-descending 7.1
Statically indeterminate methods
Walking slow with cane 2.2 Optimization: maximize endurance (10)
Walking slow without cane 3.4    
Walking 5.0   (15)
Stair-climbing 7.4
Chair-rising 3.3
Walking 5.5 Quasi static: minimize muscle forces (72)
Figure 5-5 Hip joint reaction force during walking: shaded area represents subject variability. A: Normal men. B: Normal women. (Adapted from
Paul JP. Forces at the Human Hip Joint [doctoral thesis]. Chicago: University of Chicago; 1967
, with permission.)
Clinical Applications of Hip Biomechanics
The Impact of Walking Aids on the Hip Joint Force
Both analytical and in vivo studies have clearly shown that walking with a cane in the contralateral hand reduces the force
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acting across the hip joint (10,16). This is evident from the two-dimensional static analysis discussed earlier. The moment produced from both the cane and abductor muscles together produce a moment equal and opposite to that produced by the effective body weight. The two-dimensional static analysis indicates that the joint reaction force can be reduced by 50% (from 3.3 BW to 1.7 BW) when approximately 15% BW is applied to the cane (13). The substantial reduction in the joint reaction force, predicted when a cane is used for support, arises because the cane–ground reaction force acts at a much larger distance from the center of the hip than the abductor muscles. Thus, even when a relatively small load is applied to the cane, the contribution it makes to the moment opposing body weight is large enough to lead to significantly decrease the demand placed on the abductor muscles. Using the kinematics and kinetics of preoperative total hip replacement (THR) patients who routinely walked either with or without a cane, the three-dimensional analytical optimization models showed that those who walked with a cane had a contact force of only 2.2 BW—65% of those who walked without a cane (3.4 BW) (Table 5-2) (10).
The Effect of Hip Joint Geometry on the Forces at the Hip Joint
Alterations in joint anatomy, whether due to surgical intervention or a disease process, can dramatically affect the force acting across the joint and the stresses developed within the articular surfaces. These changes occur through alterations to the moment arms of the hip muscles and the area of contact between the femur and the acetabulum. A decreased head–neck angle (varus hip) increases the mechanical advantage of the abductors (Fig. 5-6). For a given neck length, joint contact forces decrease as the neck–shaft angle is reduced (i.e., as the neck becomes more horizontal) because of the corresponding increase in medial head offset (47). Lower neck–shaft angles also improve joint stability through increased coverage of the femoral head by the acetabulum. In addition, the mechanical advantage of the abductors may be increased by moving the greater trochanter laterally, and deepening the acetabulum. Clinically, increased abductor–adductor strength has been associated with increased neck length and a more distal position of the greater trochanter with respect to the joint center (32).
The length and inclination of the femoral neck also influence the bending moments generated within the proximal femur. A varus hip and an increase in neck length will both increase the bending moment within the proximal femur by increasing the moment arm of the joint reaction force. After hip replacement, these bending moments will generate stresses within both the femoral stem and its interfaces, which, if excessive, can lead to loosening. Conversely, a shorter or more vertically inclined (valgus) femoral neck reduces the bending moment in the stem. However, the reduction in head offset means that larger abductor forces are needed to balance the weight of the body, leading to an increase in the joint reaction force. In practice, this leads to a significant increase in the wear rate of the artificial joint and a greater incidence of implant failure secondary to wear and osteolysis.
TABLE 5-3 THE EFFECT OF A 2-CM DISPLACEMENT IN THE HIP CENTER POSITION ON THE MOMENT GENERATING CAPACITY OF THE MUSCLES
  Superior Inferior Anterior Posterior Lateral Medial
Abductors –49% 26% 2% 1% –3% –8%
Adductors –18% 11% 2% 0% 40% –40%
Extensors –7% 0% 36% –36% 2% –2%
Flexors –22% 12% –34% 22% –3% –3%
From Doehring TC, Rubash HE, Shelley FJ, et al. Effect of superior and superolateral relocations of the hip center on hip joint forces: an experimental and analytical analysis. J Arthroplasty, 1996;11:693–703, with permission.
Figure 5-6 The abductor mechanism changes with head–neck angle or neck length. A valgus neck angle decreases the moment arm, while a varus neck angle or an increased neck length increases the moment arm. (Art by Judy Weik.)
Previous investigators have developed mathematical models to calculate the effect of changes in the anatomic position of the hip center on the moment generating capacity of muscles crossing the joint (17,18,47). These analyses predict that the joint reaction force will be minimized when the joint center is moved medially, inferiorly, and anteriorly (Table 5-3). This position maximizes the moment generating capacity of the abductors and brings the joint center closer to the line of action of the foot–floor reaction force, thereby reducing the external moment that must be balanced by the muscle forces acting at the hip (18,47). This analysis also predicts that superior displacement of the hip center will reduce the moment generating capacity of the abductors, adductors, flexors, and extensors due to alterations in the resting lengths and moment arm of each muscle (18). Increasing the length of the neck of the prosthesis or advancing the greater trochanter partially compensates for these losses in muscle moment generating capacities (17). An increase in the hip joint forces with a
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superolateral joint center has also been identified using an experimental setup in which a loading fixture simulated the hip abductors, adductors, and extensors during one-legged stance and stair-climbing. In this experimental setup, purely superior displacement of the joint center did not substantially increase the hip joint force (20). The clinical application of these theoretical and experimental simulations assume that alterations in joint and femoral geometry do not alter the manner in which subjects perform the activities being simulated, and that antagonistic muscle contractions are not significant.
In general, the analytical and experimental results on the effect of joint geometry on hip joint forces are consistent with clinical studies performed on patients after THR. In these studies inferior functional outcomes have been associated with superior displacement of the joint center (9) and have associated decreases in abductor strength and loss of passive hip flexion motion with superior movements of the joint centers unless the superior movements are compensated for with increased neck length (32). In addition, higher rates of femoral loosening have been associated with neck lengthening (19) and superior and lateral displacement of the joint center with respect to the anatomical position (19,85), while increased wear of polyethylene cups has been correlated with a reduction in the medial offset of the femoral head and the abductor moment arm (71).
Gait and Functional Adaptations
Individuals with degenerative or artificial joints often alter how they accomplish activities of daily living. This change in function can be considered an adaptation to a stimulus such as pain, muscle weakness, or instability. For example, during the single-support phase of gait, the contraction of the abductors prevents the contralateral side of the pelvis from dropping as the weight of the swinging leg and upper body are supported by the weight-bearing extremity. If the abductors are compromised by reductions in either muscle strength or the length of the abductor moment arm, subjects may adopt a Trendelenburg gait pattern to lessen the demand on the hip abductor muscles.
Candidates for hip replacement typically display slower walking speeds with decreased cadence and shortened step length (56). Some of the reported abnormalities of the arthritic gait pattern may be a direct consequence of the reduction in walking speed (1,2). The customary increase in the duration of stance phase on the unaffected side corresponds to an increase in swing time on the affected side, which may be needed to bring the painful hip through its arc of motion (7). Decreased step lengths on the affected side occur from both a loss of hip extension during late stance phase, as well as a decrease in maximum knee flexion during swing phase and ankle extension at toe off (56).
Preoperatively, THR patients commonly demonstrate a loss of hip motion in flexion–extension and hesitation or reversal of motion as the hip goes into extension during stance phase (42,43,56,82). This hesitation or reversal of hip motion occurs in the presence of a flexion contracture and may be reflective of compensation for a lack of hip extension through increased lumbar lordosis. It may also serve as a pain avoidance mechanism by decreasing the hip joint force (56). In the coronal plane, subjects rotate the trunk laterally over the affected hip, increasing joint stability, while also reducing the demand on the abductors and the force on the hip joint. In the transverse plane, increased external rotation at heel strike may further increase lateral stability (82).
For the vast majority of patients, hip replacement eliminates pain and restores hip function; however, despite these achievements, normal function in performing activities of daily living is not always completely achieved (11,57,74). In general, following THR, gait velocity increases due to improved cadence, step length, and hip motion. Stance times become more symmetric with reductions in lateral lurching. Postoperatively, hip motion in the sagittal plane increases but is still less than normal (42,55,77,83). This decrease in hip motion may minimize the anterior and posterior components of the hip joint force, increasing implant stability by reducing the rotational moments about the implant stem (20).
Clinical experience suggests that a valgus stem position provides better function than a varus stem position, with fewer gait abnormalities (1,14,29,40). This may be indicative of biomechanical adaptation of gait in response to the increased bending moments or the increased implant micromotion associated with loading of varus stems.
References
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8. Berzins A, Sumner DR, Andriacchi TP, et al. Stem curvature and load angle influence the initial relative bone-implant motion of cementless femoral stems. J Orthop Res. 1993;11:758–769.
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10. Brand RA, Crowninshield RD. The effect of cane use on hip contact force. Clin Orthop. 1980;147:181–184.
11. Brown M, Bahrke MS, Balke B. Walking efficiency before and after total hip replacement. Phys Ther. 1980;60:1259–1263.
12. Brown TD, DiGioia AM 3rd. A contact-coupled finite element analysis of the natural adult hip. J Biomech. 1984;17:437–448.
13. Cochran GVB. In: A Primer of Biomechanics. New York: Churchill Livingstone; 1982:240–250.
14. Collis DK. Femoral stem failure in total hip replacement. J Bone Joint Surg Am. 1977;59:1033–1041.
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