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In TF amputees, the sensory and motor functions of the foot, ankle, knee and shank do not exist. These need to be compensated for by a prosthetic device and the body’s adaptation mechanisms. The artificial foot-ankle must absorb the impact force at initial stance and simulate the plantar-dorsi-flexor muscle actions in the gait cycle. In addition, the structure of the foot-ankle complex must generate enough mechanical power and produce propulsion force to accelerate the leg. The artificial knee unit must also be able to act like a normal knee by absorbing forces transmitted from the shank and mimic the normal knee kinematics in the AK prosthesis. Practically, the components of artificial limbs cannot act exactly like anatomical structures and this leads to kinematic and kinetic differences between the amputees’ prosthetic limb and the normal limb. This further leads to compensatory kinematic and kinetic changes in the intact side and the remaining part of the limb in the amputated side (the residual limb). A large number of studies related to the biomechanics of LLAs’ gait have investigated the effects of different prosthetic components (socket designs, suspension systems, prosthetic knees and ankle/foot) which is not a matter of interest in this study. Thus, only the general characteristic of unilateral LLAs’ gait, with a focus on TF amputees walking using mechanical passive prosthetic components, will be reviewed in the following.

2.5.2.1 Biomechanics of TF amputees’ walking

Spatio-temporal variables: The variables related to timing and distances in the walking

of LLAs are significantly different from able-bodied individuals, and the differences become more prominent in higher levels of amputation. Their self-selected speed of walking is lower than the age-matched group of non-amputees. In addition, they have shorter stride length, and the stance time of their intact limb is longer than their prosthetic limb, which results in an asymetrical pattern of walking, including shorter intact limb and longer prosthetic limb steps (Jaegers et al., 1995; Nolan, Lee et al., 2003; Farahmand et al., 2006; Berke et al., 2008; Uchytil et al., 2013; De Asha et al., 2014; Khiri et al., 2015; Jarvis et al., 2017). The longer loading on the intact limb, due to its longer stance, and, in contrast, the shorter time of weight bearing on the prosthetic limb are a matter of concern because these might lead to tissue pain or joint damage in the intact limb and a decrease in the bone density of the residual limb over a long period of time (Berke et al., 2008).

Kinematics: The joint motion of the intact lower limb of TF amputees during normal

walking is similar to non-amputees, but there are various differences in the kinematics of their prosthetic side (Figure 2.11). The range of motion of prosthetic limbs has limitations in comparison to natural limbs. As can be seen in Figure 2.11-AA, a prosthetic ankle joint might provide similar plantar flexion at initial contact. But, the main difference between the prosthetic ankle joint and the natural ankle is its inability to simulate plantar-flexor muscle functions to produce plantar flexion in the late stance, which is needed for active push-off. Figure 2.11-BA also shows that mechanical passive prosthetic knees do not have knee flexion during loading acceptance. In fact, amputees cannot control the knee flexion during the loading response by generating an active extension moment in the mechanical prosthetic knee. Thus, the stance phase knee flexion is restricted by the prostheses’ alignment and features (such as the increasing friction between the moving components of the prosthetic knee during weight bearing) (Gard, 2016). Both limbs of TF amputees have a hip range of motion similar to non-amputees, with a difference in timing of its maximum extension (due to the longer stance duration), which happens at the end of the stance (Figure 2.11-CA) (Seroussi, R. et al., 1996).

Lower limb amputees have a sinusoidal pattern of COM displacement similar to non- amputees, as is shown in Figure 2.9. The range of TF amputees’ COM vertical displacement during walking does not differ from non-amputees (Gitter, A. et al., 1995; Weinert-Aplin et al., 2017). But, the intact and affected limb’s gait cycles are less symmetrical. They represent a higher position of the COM at the prosthesis’s toe-off in comparison to non-amputees. This might happen as the result of a prosthetic limb’s insufficient propulsion and the absence of plantar flexion (Nolan, Lee et al., 2003). Mediolateral displacement of the COM is larger for TF amputees than for non-amputees, which is associated with a wider BOS (Weinert-Aplin et al., 2017).

Figure 2.11 Pattern of lower limb joints’ angles, moments and powers in non- amputees (solid line), intact limb of TF amputees (dashed line) and their prosthetic side (dotted line) (Modified from Seroussi, R. et al. (1996))

Kinetics: In general, the peak values of joint moment and power are larger for the

affected limb hip joint as it requires adjustment with amputation and the compensating limitations of a prosthetic device (Winter, 1987). On the other hand, the ankle and hip of TF amputees’ intact limbs also experience a change in kinetics due to their compensatory and supporting role for the amputated limb (Nolan, L. and Lees, 2000). As seen in Figure 2.11, the longer stance of the intact side has a more obvious impact on the joint powers of the ending stance in the form of observed delays in A2, K3 and H3 of the intact limb. A prosthetic ankle-foot has a small plantar-flexor moment due to the lack of plantar-flexor muscles and, consequently, an inability to push-off actively. As the passive prosthesis remains extended during the stance phase, the knee moment is negligible (Figure 2.11-BM), and no power generation or absorption are seen in the prosthetic knee during weight acceptance (K1 and K2 in Figure 2.11-Bp). However, it has a small flexor moment and large power absorption (K3) due to the damping mechanism of the prosthetic knee (Winter, 1987).

The hip extensor moment at early stance is greater for the intact limb of amputees in comparison to the affected limb and to non-amputees, which leads to greater power generation. This is an indicator of a dramatic hip extensor concentric contraction (H1 Figure 2.11-Cp), which is supposed to compensate for the contralateral prosthetic limb’s weak push-off via facilitating forward movement of the body. The second noticeable difference in the kinetics of the hip joint is seen between the hip flexor moment and the eccentric hip flexor contractions (hip power absorption, H2) of the affected limb in comparison to non-amputees and the intact limb of TF amputees (Figure 2.11-CM and Cp). In the absence of knee flexion and the consequent smaller hip flexion during initial stance, the COM is placed posterior to the hip joint. This raises the need for a larger

eccentric contraction of the hip flexors (which is seen as a hip flexor moment and hip power absorption) to control the extension of the hip and to pull the body on an extended prosthetic leg. A sudden transition from H2 to H3 is seen in the affected hip joint of TF amputees, which is needed to unlock the prosthetic knee joint before the swing and to push the body forward in the absence of ankle push-off power (Seroussi, R. et al., 1996; Sjodahl et al., 2002).

Figure 2.12 demonstrates the pattern of vertical and anteroposterior ground reaction force changes of the prosthetic and intact sides of TF amputees during walking at different speeds. As is seen, the first peak of vertical force changes with a sharp slope after initial contact (which is an indicator of fast loading) and its peak increases at higher speed. In addition, the A/P forces are greater for the intact limb (Schaarschmidt et al., 2012). Castro et al. (2014) have reported similar findings. However, they mentioned the second peak of vertical GRF for amputees (the prosthetic side lower than in the intact limb) was lower than for non-amputees. In addition, they reported larger mediolateral force for both limbs of the TF amputees compared to non-amputees, which has been related to the movement of the COM in the coronal plane. The lower A/P forces might show that the amputees’ initial contact is more vertical; they might face difficulty when trying to decelerate the prosthetic limb and it has a lower capacity for braking (Castro et al., 2014).

Figure 2.12 Pattern of Vertical and A/P GRF of TF amputees using a passive mechanical knee joint in different walking speed (Adapted from (Schaarschmidt et al., 2012)

2.5.2.2 Gait deviations of above-knee prosthesis users

A simple observation of how LLAs walk might reveal several deviations of their gait from normal walking. The pattern of walking might have alterations due to the consequences of amputation (i.e. joint contractures, stump pain, muscle weakness) or improper construction and alignment of the prosthetic limb (i.e. loose/tight socket, length discrepancy between the prosthetic limb and the intact limb, extra stiff or easy motion of the prosthetic knee or foot-ankle). The common walking deviations of TF amputees (as the amputees who will be studied in this thesis) have been depicted in Figure 2.13 (A, B, D, I, J during stance and E, F, G during swing of the prosthetic limb) and their descriptions have been presented in Table 2.2 (Berger and Fishman, 1997; Berke et al., 2008; Rajťúková et al., 2014).

Figure 2.13 Common above-knee prosthesis users’ walking deviations: (A) Foot rotation at heel strike, (B) Lateral bending, (C) Wide walking base, (D) Swing phase whips, (E) Circumduction, (F) Uneven heel rise, (G) Vaulting, (H) Terminal impact, (I) Foot slap, (J) Excessive lordosis (adapted from (Berger and Fishman, 1997).