ck=”if (window.scroll_to_id) { scroll_to_id(event,’R3-3′); return false; }” xpath=”/CT{06b9ee1beed5941901b763cd53362c21cc9ec01f975b9454417416ca7fe6c76fa3676cc20c20abb0ac0c9091fb223267}/ID(R3-3)” title=”3″ onmouseover=”window.status=this.title; return true;” onmouseout=”window.status=”; return true;”>3). It further seals the hip joint, creating a hydrostatic fluid pressure in the intra-articular space (4), and assists in dispersing the load if impingement between the femur and acetabulum occurs. Thickness of the joint capsule has also been linked to joint stability (5). The role of the joint capsule as a stabilizer is an especially important consideration in total hip arthroplasty (THA), as research has suggested that compromise of the capsule may be an important, if not the primary, cause of dislocation after arthroplasty surgery (1). As with all synovial joints, the structure of the hip joint is intimately related to its function. Although the basic architectural design of a given joint is the same among normal individuals, there are significant enough differences in the anatomical dimensions and shapes between individuals to allow for differences in precise function. Variations in anatomical characteristics of the hip joint have been observed to affect hip joint stability, mobility, and load at the joint. Consequently, anatomical features have implications for normal hip joint function during various activities as well as dysfunction at the hip joint.
The basis for proper anatomical function of the hip joint is the congruence of the opposing surfaces. Joint congruence is controlled first by the shape of the skeletal anatomy of the opposing surfaces and secondly by the superimposed, overlying articular cartilage. The articular cartilage is contoured beyond the skeletal anatomy to allow even greater congruence. This results in cartilage of varying thickness across an articular surface. It has been shown that pressure distribution varies greatly around the acetabulum depending upon the activity. During walking and going down stairs it is the lateral roof of the acetabulum that receives peak pressure and during standing up and sitting down the posterior horn receives peak pressure (2). Thus, although some cartilage regions are experiencing load, others are less loaded or even unloaded depending on activity.
The stability and congruence of joint surfaces at the hip are affected by soft tissue structures of the joint, principally the acetabular labrum and the joint capsule. The fibrocartilaginous labrum contributes to the stability of the hip joint during extreme ranges of motion (<A onclick="if (window.scroll_to_id) { scroll_to_id(event,'R7-3'); return false; }" onmouseover="window.status=this.title; return true;" onmouseout="window.status=''; return true;" title=7 class=LK href="#R3-3" name=to-R3-3 xpath="/CT{06b9ee1beed5941901b763cd53362c21cc9ec01f975b9454417416ca7fe6c76fa3676cc20c20abb0ac0c9091fb223267}/ID(R7-3)" (7). This comprises approximately 85% to 90% of the daily activity spectrum (4), and assists in dispersing the load if impingement between the femur and acetabulum occurs. Thickness of the joint capsule has also been linked to joint stability (5). The role of the joint capsule as a stabilizer is an especially important consideration in total hip arthroplasty (THA), as research has suggested that compromise of the capsule may be an important, if not the primary, cause of dislocation after arthroplasty surgery (5). The strongest of the capsular ligaments, the iliofemoral ligament, restricts hyperextension of the hip joint. Because of the position of this ligament anterior to the hip joint, it has a special role in maintenance of erect posture, balancing the force of the body’s weight on the femur during standing.
Hip Joint Kinematics and Kinetics
The biomechanics of the hip joint is usually described in terms of kinematics and kinetics, that is, angular rotations (flexion–extension, abduction–adduction, internal–external rotations) and forces acting on the joint. Active range of motion (AROM) of the hip joint in normal healthy adults is limited by the bony architecture of the joint, soft tissue structures including the capsule and labrum as mentioned above, and to a lesser extent, the muscles that cross the hip joint. A cross-sectional investigation in over 1,600 adults aged 25 to 74 reported maximum mean AROM values in the sagittal (121 degrees of flexion and 19 degrees of extension); frontal (42 degrees of abduction); and transverse (32 degrees of internal and 32 degrees of external rotation)
planes. Interestingly, when separated into three age categories, variations between the youngest and oldest subjects were small and probably of little clinical significance with the exception of hip extension where ROM in the older subjects was over 20% less than that of the youngest (6).
planes. Interestingly, when separated into three age categories, variations between the youngest and oldest subjects were small and probably of little clinical significance with the exception of hip extension where ROM in the older subjects was over 20% less than that of the youngest (6).
A description of the relative motion of the hip during gait is an important expression of the routine motion demands at this joint because gait is the most frequent daily activity for the average adult. Daily activity levels have been observed to apps necessary to balance the external adduction moment, which acts almost throughout the entire period of stance (Fig. 3.1). The rotational moments, which act in the transverse plane, are considerably smaller than the flexion–extension and abduction–adduction moments; however, it is usually these rotational moments which give subtle hints regarding a developing patho-mechanism. Typically, during the first half of stance there is an external rotation moment, followed by an internal rotae of how the hip functions during walking is essential for therapeutic purposes. Relative motion of the hip joint during gait involves all three angular degrees of freedom. At heel strike, the hip is in a flexed position and extends throughout stance phase. During the time period from initial swing through mid-swing the hip moves from extension into flexion and stays flexed until heel strike. For a healthy adult the flexion–extension range of motion during stance spans from +30 degrees (flexion) to -10 degrees (extension) accompanied by an arc movement of approximately 10 degrees in the frontal plane and approximately 15 degrees of motion in the transverse plane (9). It should be noted that subjects with pathology often show an abnormal gait pattern with decreased ranges of motion, as has been shown for subjects with joint space narrowing (cartilage loss) (10).
Ground reaction forces, arising during human activity, impose external forces and external moments at the hip joint, which must be balanced by a set of internal forces. These internal forces consist of contact forces, muscle forces, and soft tissue constraints. The muscles are in the best position to resist the external moments, because they have sufficient lever arms, defined from their lines of action to the point of contact at the joint (Fig. 3.1). The external moment patterns during the stance phase of gait are shown in Figure 3.2. Typically, at heel strike, there is an external moment tending to flex the hip joint. To balance this moment internally, the extensor muscles have to become active. As the hip moves into mid-stance, the external moment becomes zero and reverses direction, demanding the action of the flexor muscles during terminal stance. In the frontal plane, the activity of abductors is necessary to balance the external adduction moment, which acts almost throughout the entire period of stance (Fig. 3.1). The rotational moments, which act in the transverse plane, are considerably smaller than the flexion–extension and abduction–adduction moments; however, it is usually these rotational moments which give subtle hints regarding a developing patho-mechanism. Typically, during the first half of stance there is an external rotation moment, followed by an internal rotation moment in the second half. Characteristic peak moment values during gait for healthy adults with a mean age of 55 are shown in Table 3.1. Considering an average American male with 190 lb (86.2 kg) and 5′10″ (1.78 m) tall (BMI = 27.3 kg/m2), the flexion moment amounts to 102 Nm, the adduction moment to 74 Nm and the internal rotation moment to 13 Nm. To put these values into context, the flexion moment is comparable to the maximum torque of a 74 hp engine in a SMART car. This analog makes clear that the muscles need sufficient lever arms to counteract the external moments.
 Figure 3.1. The ground reaction force (GRF) acts to adduct the hip joint. The muscles work to internally counteract the externally generated moment and keep the joint stable. Because of the smaller lever arm of the muscles compared to that of the ground reaction force, the muscles need to pull in excess of the GRF. The sum of both GRF and muscle force is balanced by the joint contact force, which typically reaches up to two and a half times the body weight during gait. |
Table 3.1 Motion and Peak Moments During Gait of Healthy Adultsa | ||||||||
|---|---|---|---|---|---|---|---|---|
Forces at the Hip Joint The external moments measured in a gait laboratory are surrogate markers of the load experienced at the hip joint during ambulation. Actual forces at the hip must be either measured directly from implants instrumented with force transducers or estimated using analytical models with external moments and/or EMG as input. The measurement of hip forces is limited to subjects receiving implants (e.g., total hip or hemi-arthroplasties). Because it requires complex modifications of the implant to incorporate strain gauges, amplifiers, and telemetry (to avoid extruding cables), only data from a very limited number of arthroplasty patients is available. Rydell (19) published the first such work in 1966. The most complete data set is available from Bergmann et al., (20,21) who published data on five subjects. During level walking the force at the hip joint generally reaches an initial peak in early stance and a second peak in late stance (17,18). However, EMG can be easily used to determine which muscle is active during a particular period of the gait cycle and to what extent.
Forces at the Hip Joint The external moments measured in a gait laboratory are surrogate markers of the load experienced at the hip joint during ambulation. Actual forces at the hip must be either measured directly from implants instrumented with force transducers or estimated using analytical models with external moments and/or EMG as input. The measurement of hip forces is limited to subjects receiving implants (e.g., total hip or hemi-arthroplasties). Because it requires complex modifications of the implant to incorporate strain gauges, amplifiers, and telemetry (to avoid extruding cables), only data from a very limited number of arthroplasty patients is available. Rydell (19) published the first such work in 1966. The most complete data set is available from Bergmann et al., (20,21) who published data on five subjects. During level walking the force at the hip joint generally reaches an initial peak in early stance and a second peak in late stance (Fig. 3.4). These peaks are usually similar in magnitude; however, different muscles are active during these two phases of the gait cycle. The average peak load at the hip while walking at a “normal” speed is approximately 240% body weight (BW). This is slightly more than standing in single leg stance. When going upstairs the joint contact force is 250% BW and downstairs 260% BW. The peak contact forces during all other common daily activities are comparatively small, except for stumbling. Peak forces during unanticipated stumbling have been found as high as 870% BW (<A onclick="if (window.scroll_to_id) { scroll_to_id(event,'R22-3'); return false; }" onmouseover="window.s to handle (and thus more common) because it does not need an a priori knowledge of structure properties. All it needs is knowledge of the geometry of the involved structures and their displacement vectors over time.
Following the inverse dynamics approach and knowing the three-dimensional position of the limb segments and the ground reaction force, it is then assumed that the external forces and moments are balanced by a set of forces and moments acting internally, which are primarily generated by muscle contraction, other soft tissue tension and articular reaction forces. Because of the redundancy of the internal structures, this approach, however, induces more unknowns than can be solved with the number of equations available. In general, two attempts have been made to solve this indeterminate problem. The first reduterns as compared to age-matched control subjects (24,25
Related posts: Stay updated, free articles. Join our Telegram channel
Full access? Get Clinical Tree
|
