Biomedical Engineering Reference
In-Depth Information
Initial swing:
Gracilis and Sartorius muscle function is partially affected by the ampu-
tation insofar as length and effectiveness are compromised. Amputees accommodate for
this using increased hip-flexor muscle activity to lift prosthetic feet off the ground and
initiate the swing phase. In addition, amputees need to rotate the pelvis and lower trunk
while applying load on the toes of the prosthetic. Because there is no muscle activity at
the level of the knee or foot to initiate toe off, ground clearance is compromised. The loss
of the popliteus and the short head of the biceps and the pretibial muscles eliminate active
knee flexion and toe lift.
Mid-swing:
Limitation of knee flexion becomes more critical and results in the prosthesis
being lifted by using the hip (hip hiking), vaulting, or circumduction to compensate for
reduced ground clearance. These gait deviations result in asymmetrical loading of the body,
which puts strain on the lower back and requires more energy. Some of these limitations
are due to knee design that often requires toe load to initiate flexion.
Terminal swing:
Prosthetic knees all suffer from friction, which reduces angular acceler-
ation of the knee during the advancement of the lower leg in the swing phase. This reduces
the pendulum motion, and users must overcome this by means of additional muscle activity
and abnormal motion patterns to complete the swing. Typically, users kick the knee into
extension by contracting the gluteus muscles. This is abnormal and reduces the efficiency
of the gait. The reaction also shifts the center of mass backward, counter to the walking
direction, which affects balance and walking effectiveness. This passive swing extension
also limits the possibility of stumble recovery because if the knee does not reach full
extension in time it cannot accept the load of the next step and users generally fall.
Initial contact and loading response:
This phase is responsible for stance stability and
shock absorption. A passive prosthesis cannot produce the knee-flexion motion for shock
absorption because the knee needs to be fully extended or even hyperextended to prevent
it from buckling into flexion. Because the amputee kicks the prosthesis into extension, the
initiation of flexion after initial contact is slower and more difficult with the result that
there is a larger vertical displacement of the center of mass and therefore lower efficiency.
10.8.3 Knee Prosthetics
Knee kinematics is complex and difficult to quantify exactly due to variations between
subjects and because the joint does not operate as a simple hinge. However, follow-
ing motion studies undertaken by Walker (Walker, Kurosawa et al., 1985), it is possible
to describe the instantaneous position of the hinge axis in the anterior-posterior axis,
z
dis
(mm), and the proximal-distal translation,
y
dis
(mm), in terms of the flexion angle,
θ
f
(deg):
2
f
y
dis
=−
0
.
05125
θ
f
+
0
.
000308
θ
(10.11)
f
z
dis
=−
0
.
0602
θ
f
+
0
.
0000178
θ
A guide or cam can be used to produce this motion curve, but these present drawbacks
in terms of high manufacturing cost and resistance and pinching between the parts. An
alternative is the crossed four-bar linkage described in Chapter 3, which has the advantages
of simplicity, robustness, and ease of design.
