Hip

During initial contact / weight accceptance:

• The hip moves toward extension immediately after heelstrike(7, 12), although the range of extension is less than normal and may even stay in flexion(11), particularly with shorter stumps(7).
• Peak hip extension occurs later than normal (~65% stride c/w 55%), suggested to be an attempt to increase intact limb step length in absence of prosthetic push off(7).
• Concentric hip extensor activity is present, to generate a hip extensor moment(22).
• The hip extension moment is greater than normal, and continues through the first half of stance(19, 22). This is used to extend the knee and hip, and propel the Head/Arms/Trunk (HAT) segment forwards.
• The gluteus maximus is active over a longer period through early stance, up to the first 25-40% compared with the first 15% in non-amputees(28).The same study noted that in amputations in the proximal half of the femur, the gluteus maximas may be active during the entire stance phase, particularly when the iliotibial band is not fixed.

During mid stance:

• The peak height of the hip (and centre of mass) is higher on the prosthetic limb than normals. This occurs due to reduced flexion of the prosthetic knee, and may also be influenced by the characteristics of the prosthetic foot(33).

After contralateral foot contact:

• There was an abrupt transition from hip extension to flexion(7, 12)

During late stance:

• There is larger than normal eccentric hip flexor activity, generating a hip flexor moment(22). This greater moment is required to decelerated the extending hip, and ensure the HAT segment does not lag behind as the amputee pulls themself over the prosthesis(19).

During early swing:

• The hip flexors act concentrically to initiate a "pull-off". There was a trend towards greater power generation than normal, and over a shorter period(19, 22). Thus the same power seems to be required to lift a prosthesis that may be only 30-40% of the weight of a normal limb, and that same power is generated over a shorter period.
• The hip torque required during preswing and early swing also appears to be less in knees with variable damping and stance resistance(9) These authors hypothesise that the reason is due to the alignment of the knee axis, which is placed more anteriorally in mechanically passive knees so as to increase the extension moment during heel contact. However, this anterior alignment may also make the knees more difficult to initiate flexion in preswing / early swing.

During late swing:

• The range of hip flexion was greater than normal, at all walking speeds(7), as was the amount of hip flexor activity(28). This was postulated to increase step length, but also by keeping the hip flexing helped decelerate the extending knee and reduce terminal impact.
• Some amputees with constant friction knees transitioned from hip flexion, to extension, to flexion again (Murray 1983), possibly to assist in knee extension by reversing the direction of the thigh.
• Some other authors(5) found intermediate changes in hip moments, rather than the typical flexion / extension. They found:
• A hip flexor moment to initiate swing & cause knee flexion through the weight & inertia of the shank
• A hip extensor moment to decelerate the the thigh & cause extension of the knee.
• A second hip flexor moment to further flex the hip and increase step length.
• A final hip extensor moment to decelerate the thigh and maintain the knee in extension in preparation for heel strike.

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Knee

During initial contact / weight acceptance:

• The knee is typically in full extension, forced there by a combination of hip extensor activity, body placement, prosthetic alignment, and sound limb push-off(33). With most knee types there is no knee yield / shock absorption(7, 12, 14, 25). This lack of flexion is postulated to reduce the 'rebound' at the hip as well, which contributes to the increased activity seen in all hip muscles on the amputated side(28).
• Despite the claims of the manufacturer, some authors(16, 18) found that stance phase yielding was not improved with the C-Leg after 3 months acclimation. Despite an increase in the knee flexion moment, both C-Leg and Mauch SNS remained in constant extension throughout stance. This may have been due to the previous long-term use of the Mauch SNS knee, as the amputees commented that it was difficult to break the habit of applying extensor force to brake knee flexion in stance phase. Other authors did, however, find an increase in stance phase knee yielding with microprocessor controlled knees versus the Mauch SNS(32).

During mid stance:

• There is a knee flexor moment throughout most of stance phase, due to the extensor stop in mechanical knee joints (14, 19, 22).

During late stance / preswing:

• Knee flexion is initiated later (with most knees), starting at around 58-62% stride rather than 40-45%(7). This is around 20% later than normal, most likely due to the need to maintain full extension during weight bearing in mechanical knees, and the need to load to toe to release stance resistance or locking in more advanced knees.
• The knee flexion was initiated by hip flexion.

During early swing:

• Peak heel rise is greater than normal, and greater with mechanical / constant friction knees than hydraulic or pneumatic knees(2, 10, 12). The greater the heel rise, the longer the swing phase and greater the asymmetry.
• Peak heel rise with the C-Leg is reported to be decreased compared to the Mauch SNS knee(18, 30), Rheo Knee, Adaptive 2 Knee, and Hybrid Knee(30) at various walking velocities, and is closer to heel rise of the sound limb, suggested to lead to improved gait symmetry.

During late swing:

• There is less energy absorption than normal due to the slower swing phase, and lighter prosthetic shank(22).
• It has been reported that the C-Leg provides better terminal extension damping(30) at all tested velocities.

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Ankle

Kinematics depend on the type of foot, however, most notable is a slightly earlier heel rise in late stance, and the lack of plantarflexion for pushoff.

Power profiles show absorption until midstance due to the viscous nature of the prosthetic foot, but no power generation at push off (with a SACH foot(22)), or up to 20% normal power generation (with a Seattle foot(19)).

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Lateral Displacement

During initial contact / weight acceptance:

• There is a common tendancy to shift the trunk towards the stance side(7, 12). There was no correlation between stump length or stride width and the amount of side flexion, however those with the greatest side flexion had the most abductor muscle atrophy and low maximal isometric abductor torque(7).
• The abductors may also be less effective at controlling frontal plane movement due to abduction of the femur inside the socket(12).
• Despite kinematics indicating the difficulty the abductors have in controlling lateral movement, they are active over a longer period than normals, up to 53% of stride time, or even the entire stance phase with shorter stumps(28). This compares to a peak in midstance for non-amputees.
• Prosthetic causes, such as a high medial socket wall, must also be considered when analysing frontal plane movements(12).
• There is also some suggestion(31) that amputees have a wider step to compensate for the lack of an ankle strategy, which in iintact limbs is used to correct inaccuracies of foot placement. Increased trunk sway and use of hand support may similarly compensate for this lack of lateral ankle control.

Some amputees show an increase in abductor activity at the beginning of swing phase, which may be explained by the use of circumduction (flexion & abduction) to initiate swing phase(28).

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Pelvis / Trunk

Studies are finding that pelvic and upper body motion in amputees is not normal compared to controls. One(24) found angular ranges of motion are globally increased in the upper body, despite a lower walking velocity. The coordination between pelvis and upper body was altered, being more 'in-phase' rather than the out-of-phase counter-rotation observed in controls. They also found pelvic tilt tended to be down on the prosthetic stance side, rather than down towards the swing side, caused by the inability of the hip abductors to stabilise the pelvis, and an increased trunk lateral flexion to the prosthetic side.

Other authors have found that the pelvis tilts forward on the amputated side during stance, and tilts backwards during swing(25). They have suggested the forward tilt was an attempt to keep body weight forwards and stabilise the knee, or an overactivity of iliopsoas. The backward tilt during swing was suggested to be due to a possible increased rounded lumbar spine as the abdominals were recruited to assist in lifting the prosthesis, although they did not measure abdominal EMG.

A further group found that after heel strike the pelvis retracts and the stump externally rotates, suggesting a twisting motion inside the socket. The most likely reason for this is the inability to stabilise the pelvis and residual femur. Inclusion of a torque absorber allowed the socket to rotate with the stump and pelvis, which they suggested would reduce shear forces on the skin(29).

The higher the level of transfemoral amputation, the more the muscles are affected, leading to greater insufficiency and inability to stabilise the pelvis(28).

In the trunk, both sides of the erector spinae muscles become active immediately after heelstrike of both feet, as a reaction to weightbearing, to avoid ventral flexion of the pelvis, to assist in moving the centre of mass of the trunk to the weighbearing foot, and as a reaction to lateral trunk bending. In transfemoral amputees, the activity is over a longer period compared to normals (0-20% vs 1-10%, and then 40-65% vs 50-65% of the stride)(28).

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Intact Limb

The intact limb is known to compensate for the loss of function on the amputated side. This is suspected to lead towards increased degenerative changes in the intact limb.

During initial contact / weight accceptance:

• There was a 270% increase in concentric hip extensor work compared to controls(19), suggested to compensate for the lack of prosthetic push off on the amputated side.
• The knee usually flexes although the amount may depend on walking velocity. One group of authors found the degree of flexion for shock absorption was smaller due to a slower walking speed(7), but others(25) found that knee flexion range and extensor moment were higher (almost double), and eccentric power almost 5 times as high than both amputated side and controls in a group who walked at a similar speed to the controls.

During mid stance:

• The knee did not always fully extend through midstance(7), although the quadriceps concentric power has been found to be higher than controls(25).

During late stance / preswing:

• There is generally a decrease in hip extension range compared to controls(7).
• There is also a decrease in eccentric hip flexor activity, due to greater concentric activity earlier in the cycle(19). However, the hip flexion moment and range is greater for the intact limb than the amputated side and the reference group used for comparison, despite lower peak hip flexor power generation(25).
• Intact knee flexion started later in the gait cycle compared to normal (51-58% stride, compared to 40-45% in normals)(7).
• Some amputees demonstrated increased plantarflexion range, up to 15 degrees greater than normals. This increase also occassionally occurred earlier (in midstance) as the amputee vaulted over their prosthesis(7).
• The ankle power generation was 33% greater than normal(19) during push off.

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References

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22. Winter DA, Olney SJ, Conrad J, White SC, Ounpuu S & Gage JR (1990). Adaptability of motor patterns in pathological gait. In: Winters JM, & Woo SL-Y (eds) Multiple Muscle Systems: Biomechanics and Movement Organisation. Springer-Verlag, New York. 680-693.


23. Zuniga EN, Leavitt LA, Calvert JC, Canzoneri J & Peterson CR (1972). Gait patterns in above-knee amputees. Archives of Physical Medicine & Rehabilitation, 53, 373-382.


24. Goujon-Pillet H, Sapin E, Fode P & Lavaste F (2008). Three-dimensional motions of trunk and pelvis during transfemoral amputee gait. Archives of Physical Medicine & Rehabilitation, 89, 1, 87-94.


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26. Sjödahl C, Jarnlo G-B, & Persson BM (2001). Gait improvement in unilateral trans-femoral amputees by a combined psychological and physiotherapeutic treatment. Journal of Rehabilitation Medicine, 33, 114-118.


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29. Van der Linden MI, Twiste N & Rithalia SVS (2002). The biomechanical effects of the inclusion of a torque absorber on trans-femoral amputee gait, a pilot study. Prosthetics & Orthotics International, 26, 35-43.


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33. Stein JL & Flowers WC (1987). Stance phase control of above-knee prostheses: Knee control versus SACH foot design. Journal of Biomechanics, 20, 1, 19-28.


34. Winter DA & Sienko SE (1988). Biomechanics of below-knee amputee gait. Journal of Biomechanics, 21, 361-367.

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