The forces applied in the hip joints are the sum of body weight and muscle forces, expressed as a resultant force (R). The magnitude and distribution of the compressive stresses depend on three factors: 1) the magnitude of R which is affected by the body weight and balance of the muscle forces, 2) the area of the weight bearing surface which is basically determined by the size of the femoral head, 3) the orientation of R. The weight bearing areas of the contact surfaces in the hip joint were
found to range from 22.19 to 33.68 cm^ with an average of 26.77 cm2 under load, measured by Greenwald and Haynes (1972).
During the "single-legged stance" of gait, the centre of gravity moves towards the opposite side. The moment of the abductor (MA) balances the moment of partial body weight (MW) (full body weight excluding the swinging leg). The resultant force R intersects at the centre of the femoral head and diverges distally, divided into compressive, tensile, and shear forces (Fig.3.1). Because of the offset of the femoral head related to the long axis of the femur, it creates a bending moment which is function of the offset of the femoral head and the loading force. Since the force R acts to compress as well as to bend the femoral neck, the maximum compressive stresses on the medial side are always greater than the tensile stresses on the lateral side. The magnitude of the shearing component 8 depends on the inclination of the force R to the axis of the femoral neck. The torsion about the long axis of the femur increases in stair climbing and rising from a low chair, and is also affected by the anteversion of the femoral neck.
It has been observed that the force acting on the hip joint is greater than body weight due to the muscle forces surrounding the hip. However, because of the complexity and integrity of the hip joint, in the past, it was difficult to measure in vivo the forces transmitted by the hip joint and associated periarticular structures. Therefore a force plate has traditionally been used to mathematically determine the intersegmental force and moment resultants at the hip joint by modelling the body or parts as a rigid system. Because of the large number of load-carrying elements (muscles, ligaments) that act across the joint, it is unlikely to calculate the joint contact forces accurately and realistically. With the
development of advanced telemetry technique, It has been possible to measure in vivo the contact forces of the hip joint by using instrumented prostheses (Rydell 1965, English and Kilvington 1979, Davy et al 1988, Bergmann et al 1993). However, the hip forces which were measured directly in vivo were usually in the situation of post-operation, where the muscles surrounding the joint may have been weakened. Consequently the forces which were measured across the femoral head have not entirely agreed with the forces which were calculated.
Generally the predicted force by using analytical models is higher than that experimentally measured. These differences may be explained in part by the fact that mone of the experiments used subjects who had a normal pattern of gait. Paul (1965, 1976) predicted that the peak contact force of the hip joint is in the range of 3.5 to 6 times body weight. This was close to that calculated by Crowninshield (1978), whose result ranged from 3.3 to 5 times body weight. These results were also agreed by Hardt (1978), which were 5.7 times body weight. In 1965, Rydell used a strain gauged Austin Moore prostheses to directly measure the forces across the hip joint. The results showed three times body weight. Similar results were obtained by English and Kilvington (1979) with a telemeterized hip prosthesis. Davy et al (1988) recently used a telemeterized total hip prosthesis for a patient. They found that the peak resultant force during single-limb stance was 2.1 times body weight with the direction of force at 32 degrees respect to the axis of the neck of the prosthesis and 15 degrees towards posterior from the midplane of the prosthesis in the sagittal view. For ascending stairs, the peak force increased to 2.6 times body weight, and the orientation of the force changed to 70 degrees posteriorly. Bergmann et al (1992) used instrumented prostheses in two patients to measure the three
dimensional hip forces acting on the head, and to calculate the torsional moments around the stem axis. The results varied between the two patients. Overall, the torsional moments for up-stairs were 50 % higher than that for the level walking, but there were no significant differences between down-stairs and level walking. The force also depended upon two other factors: time after operation and speed of walking. Bergmann et al (1993) measured two patients using a telemetrised total hip prostheses. The results showed that the loads stayed nearly constant for the first several months postoperatively, followed by an increase for another several months, then a succeeding decrease and remained constant. The force also had a linear relationship with the walking speed, the faster speed producing higher forces (English and Kilvington 1979, Bergmann et al 1993). However, Kotzar et al (1991) pointed out that in the range of walking speeds that are normal for a given individual, the peak joint contact forces are not strongly affected. Only when the walking speed exceeds the normal range does a pronounced speed effect become apparent.