Biomechanics of the Hip

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Original Editor - Ewa Jaraczewska

Top Contributors - Ewa Jaraczewska, Jess Bell, Rucha Gadgil, Kim Jackson and Lucinda hampton  

Static Loading

Koch’s model

Koch’s model is static, unlike the conditions that occur during gait, where the body’s center of gravity usually remains medial to the supporting foot in a dynamic equilibrium. Secondly, Koch limits the source of resistance to the body’s varus load to a single structure, the gluteus medius. These two factors have dramatic consequences for conclusions drawn from the Koch model. The most important of them is the conclusion that, below the insertion point of the gluteus medius, the entire lateral aspect of the lower extremity experiences a tensile load. Koch went so far, in his original article, as to attempting to qualify the loads experienced by the femur during unilateral support. He used positive integers to represent compressive forces and negative integers to designate areas of tensile loading (Figure 2). Intuitively, it can be understood that he labeled the entire medial aspect of the femur as experiencing compressive loading. However, curiously he does not explain how the lateral femur experiences tension along its lateral proximal 2/3’s but coverts to compression along the lateral distal 1/3 of its length. It may be assumed that Koch was attempting to keep his model consistent with the reality, and contemporary belief, that there is compression within the lateral compartment of the knee. It was believed, at that time, to be the explanation for the presence of a lateral meniscus within the knee’s lateral compartment. An element intended to buttress against compressive loading.

at birth the neck-shaft angle of the femur is similar to that of a quadruped, approximately 160-165 degrees. As one assumes, in bipedal gait, by the age of four, the neck-shaft angle reduces to its final value of 130-135 degrees and remains so throughout the remainder of a lifetime. This has been explained as being the result of the upright posture placing a varus load on the young plastic bone. However it is interesting, that in spite of significant time spent in an upright posture and with increasing body mass, the femoral neck-shaft angle remains relatively unchanged throughout the remainder of growth and development of the femur.

Bone, as all connective tissues, responds to its environment. It hypertrophies and atrophies in response to demand. Also, the quality of bone reflects the type of load it experiences: cortical bone appears in areas of compression and cancellous bone in areas of tensile loading, i.e., apophyses and points of tendon attachment. It would appear that, rather than “reconstruct” damaged hip anatomy, femoral component designs and various methods for their fixation, based on the Koch model, have created nonphysiologic patterns of loading within the femur following hip replaceme

When the weight of the body is being borne on both legs, the centre of gravity is centred between the two hips and its force is exerted equally on both hips. Under these loading conditions, the weight of the body minus the weight of both legs is supported equally on the femoral heads, and the resultant vectors are vertical.

It correctly assumed that the body’s center of gravity (COG) was located in the midline of the body, 1 cm anterior to the first sacral segment. From this point the effect of gravity on the lower extremity was represented by a vertically oriented, downward directed vector.

Joint stability is a function of two factors: articular geometry and soft tissue integrity. In terms of geometry, it is easy to describe the three major joints of the lower extremity as being either inherently geometrically stable or unstable. The ankle, although the smallest of these three joints, has an extremely stable mortise structure with very limited degrees of freedom, particularly in the dorsiflexed position during weight bearing. As such it suffers minimal frictional wear throughout a lifetime unless the integrity of the mortise is compromised by either a fracture or more subtly disruption of the syndesmosis ligament. In either event, translational movement increases across the articular surfaces and demise of the ankle is rapid. With regard to soft tissues contributing to the stability of a joint, they can be divided into static and dynamic structures. Each plane of movement is stabilized by a couple comprised of a dynamic and a static entity. The dynamic structure is a muscle-tendon unit and the static stabilizer is a ligament. Muscles can adjust their length in response to demand. Ligaments however cannot. Hence since ligaments are critically important at extremes of motion and are lax at mid-range, all ligaments are composed of two parts: one taut in extension, the reciprocal part taut in flexion.

iliotibial band (ITB) is that static stabilizer of the hip against varus loads [9].

From this observation, it was concluded that the ITB should be included in the analysis of hip biomechanics to produce a more complete, dynamic, and accurate model of hip stability. The inclusion of the ITB served to resolve many of the paradoxes raised by the previous static model. It resolved the seeming contradiction of the gluteus medius being less active at the midstance phase of gait. At that point of the gait cycle the ITB apparently serves as a tension band to relieve the metabolic demand and reduce electrical activity of the gluteus medius (Figure /www.hindawi.com/journals/aorth/2019/5804642/fig4/|4). It also provides an explanation for the poorer functioning of an AKA relative to that of a BKA, where loss of the distal attachment of the ITB in an AKA compromises the function of the ITB as a static stabilizer of the hip joint. It further gives rationale to the surgical technique of tenodesis of the ITB and lateral soft tissue structures to the distal femur in the performance of an above-the-knee amputation. This would be similar to the technique of wrapping the posterior calf musculature around the distal end of a below-the-knee amputation.

Single Leg Stance

the effective centre of gravity moves distally and away from the supporting leg since the nonsupporting leg is now calculated as part of the body mass acting upon the weight-bearing hip This downward force exerts a turning motion around the centre of the femoral head – the moment is created by the body weight, K, and its moment arm, a (distance from femur to the centre of gravity). The muscles that resist this movement are offset by the combined abductor muscles, M. This group of muscles includes the upper fibres of the gluteus maximus,the tensor fascia lata, the gluteus medius and minimus, and the piriformis and obturator internus. The force of the abductor muscles also creates a moment around the centre of the femoral head; however this moment arm is considerably shorter than the effective lever arm of body weight. Therefore the combined force of the abductors must be a multiple of body weight. The magnitude of the forces depends critically on the lever arm ratio, which is that ratio between the body weight moment arm and the abductor muscle moment arm (a:b) [20]. Typical levels for single leg stance are three times bodyweight, corresponding to a level ratio of 2.5. Thus, anything that increases the lever arm ratio also increases the abductor muscle force required for gait and consequently the force on the head of the femur as well (see Fig. 4). People with short femoral necks have higher hip forces, other things being equal. More significantly people with a wide pelvis also have larger hip forces. This tendency means that women have larger hip forces than men because their pelves must accommodate a birth canal [21].

Joint Forces at the Hip

Rydell NW. Forces acting on the femoral head-prosthesis. A study on strain gauge supplied prostheses in living persons. Acta Orthop Scand 1966; 37(Suppl 88): 1-132.

These studies have shown that although patients in the early postoperative period can execute planned activities of daily living with relatively low joint contact forces, unexpected events such as stumbling or periods of instability during single leg stance can generate resultant forces in excess of eight times body weight [25]. It is important to remember that although the data from hip prostheses have established the magnitude of the loads acting on the hip joint, the patients in these studies have undergone total hip replacement and therefore the results cannot be directly correlated to the physiology of the normal hip.

The average patient loaded his hip joint with 238% BW (percent of body weight) when walking at about 4 km/h and with slightly less when standing on one leg. This is below the levels previously reported for two other patients (Bergmann et al., Clinical Biomechanics 26 (1993) 969-990). When climbing upstairs the joint contact force is 251% BW which is less than 260% BW when going downstairs. Inwards torsion of the implant is probably critical for the stem fixation. On average it is 23% larger when going upstairs than during normal level walking. The inter- and intra-individual variations during stair climbing are large and the highest torque values are 83% larger than during normal walking. Because the hip joint loading during all other common activities of most hip patients are comparably small (except during stumbling), implants should mainly be tested with loading conditions that mimic walking and stair climbing.

Dynamic Model of Biomechanics

When Koch applied this analysis to the hip joint, he assumed the counterbalancing valgus torque would be supplied by isometric contraction of an abductor muscle (A), specifically the gluteus medius. He further assumed that the average length of the body’s lever arm (b) from the center of the hip’s rotation was approximately twice that of the length of the abductor’s insertion on the lateral aspect of the greater trochanter to the center of hip rotation (a), b = 2a.

From these assumptions, Koch wrote the equation for hip stability as B x b = A x a. And since b = 2a, he concluded that the gluteus medius must generate twice the weight of the body force, 2 x B, in order to maintain a stable equilibrium during unilateral stance. (Figure 1).

Painful Hip

Management of painful hip disorders aim to reduce the joint reaction force. Bearing in mind the basic principles outlined above, this can be achieved by reducing the body weight or its moment arm, or helping the abductor force or its moment arm. Increases in body weight will have a particularly harmful effect on the total compressive forces applied to the joint. The effective loading of the joint can be significantly reduced by bringing the centre of gravity closer to the centre of the femoral head (decrease the moment arm b). This can be accomplished by limping, however the lateral movements required take a considerable amount of energy and is a much less efficient means of ambulation. Another strategy to reduce joint reaction force involves using a cane or walking stick in the opposite hand. 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 times

 body weight to 1.5 times body weight) when approximately 15% body weight is applied to the cane [19]. 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 centre 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 significantly decrease the demand placed on the abductor muscles

Stair Climbing

Our study aimed to compare stair ascent and stair descent biomechanics between individuals with hip OA and asymptomatic healthy controls. Participants with hip OA ascended stairs with less hip range of motion in all three planes, and a lower hip peak external rotation moment compared to controls. During stair descent, participants with hip OA descended stairs with greater ipsilateral trunk lean, reduced sagittal plane range of motion, a lower external peak extension moment, a lower external peak rotation moment, a higher hip adduction moment, higher hip adduction moment impulse and internal rotation moment impulse. These findings improve our understanding of the influence of hip OA on the biomechanics of stair negotiation and may assist clinicians and researchers to optimize conservative rehabilitation.

Lower Extremity Amputation

It has been found that below knee amputees (BKA) do not usually exhibit a positive Trendelenburg gait pattern. They only lose about 10% of their metabolic efficiency and with today’s technological and material advancements can function at near normal levels of performance, while above knee amputees always exhibit a positive Trendelenburg gait pattern and lose 40-70% of their metabolic efficiency. Yet both the BKA and the AKA have intact abductor, gluteus medius, and musculature. The question is an obvious one. What is lost in an AKA that creates such a significant compromise of the gluteus medius’ ability to provide stability against the varus load of the body during gait? (Table 1)