The spatial-temporal and kinematic measurements have been useful in the analysis of gait param- eters [229]. The objective was to determine if there are any gait anomalies on unilateral amputees as a result of using passive mechanical prosthetic ankles. Therefore, the distance variables and anatomical angles were evaluated. The use of state of the art recording equipment, such as the Noraxon MyoMotion System and processing the data in MR3, minimised the need for excessive ltering. The results presented in this section are categorised as distance variables, orientation and anatomical angles. A t-test was performed on all the distance variables, anatomical and orientation angles so as to determine the signi cance of the study and the signi cant dierence between the amputees and the normal subjects.
5.3.1 Analysis of distance variables
The results presented in this section include cadence, velocity, stride time, stride length, step time and stance, as illustrated in Table 5.1. These parameters were evaluated so as to determine anomalies using a benchmark of normal ambulators and knowledge of gait. The results of the t-test are presented in Table 5.1.
Table 5.1: Average and standard deviations of temporal distance factors
Item Parameter Unit NormalValue AmputeeAverage
Value p-value 1 Cadence steps/min 110 107.23 ± 3.50 0.211 2 Velocity m−s 1.5 1.41 ± 0.07 0.088 3 Stride time s 5.6 5.60 ± 0.18 1.00 4 Stride length m 0.8 0.79 ±0.03 0.878 5 Step time s 1.12 1.13 ± 0.05 1.876
The results presented in Table 2 revealed that there is no signi cant dierence (p > 0.05) with respect to distance variables for normal subjects and amputees. Such results were also reported in previous studies [193]. However, there was evidence of unequal step length with respect to the amputees, resulting in an asymmetrical gait. The amputees were able to perform all ambulatory activities and hence achieve the same results as the normative data from the MR3 system, despite minimum variations. The step size was reduced each time the amputated leg with the prosthetic attachment was used as the leading leg. The dierence was approximately 0.15 m, therefore it was too small to be signi cant (p = 0.878). Furthermore, the stance time on the amputated limb was slightly lower than on the normal side, resulting in limited stance time and early swing. The amputees exhibited decreased cadence as compared to normal subjects as a result of the reduction in self-selected walking velocity. However, the 2.8% decrease in cadence caused no signi cant dierence in the walking velocity (p = 0.211) of the amputees.
In previous studies, it was shown that healthy subjects have the unique capability of utilising the dorsi exors and plantar exors to reduce the dorsi exor moment at heel strike so as to rapidly reach foot- at, resulting in an early stance. However, the sti ankle on mechanical prosthetic limbs used by amputees generated an extended moment which aects the transition from stance to early swing gait phase.
The reduction in stance time was as a result of reduction in the time taken during mid stance (single support) as the amputee avoids loading the amputated leg for a long period of time. The load bearing positions of the residual limb were sensitive to increased load, hence the stance phase occurs at 58% during the gait cycle.
5.3.2 Analysis of anatomical angle parameters
The anatomical angles investigated include the hip, knee and ankle for the unilateral amputees. The two-sample t-test was performed assuming unequal variances. The results prompted further investigations into the gait cycle of the amputee as there existed a signi cant dierence (since p < 0.05) between the gait of a normal subject and that of an amputee. The results presented in Table 5.2 indicated that there is signi cant dierence between the amputee and the non- amputee normative data with respect to hip exion-extension, knee exion-extension and the ankle dorsi exion-plantar exion movements. The signi cant dierence (p = 0.0060) recorded on the ankle inversion is as a result of a lack of an intuitive control of the prosthetic limb by the user. The prosthetic limb orientation was being governed by the sti ankle mechanism, hence the only possible motions were along the sagittal plane.
Table 5.2: t-test: analysis of anatomical angles
Item Movement p - value 1 Hip exion 0.0290 2 Hip abduction 0.0001 3 Hip rotation 0.0010 4 Knee exion 0.0020 5 Ankle dorsi exion 0.0300 6 Ankle inversion 0.0060 7 Ankle abduction 0.0400
The analysis involved evaluation of the hip exion-extension, pelvis tilt and dorsi exion- plantar exion movement of both the amputated and the sound leg as follows:
Hip exion-extension and the pelvis tilt
The hip exion recorded a mean peak of 360 on the amputated left leg as compared to the mean
peak of 250 on the amputee's intact right leg, as shown in Figure 5.1. The 44% additional exion
was as a result of the amputee's eort of trying to compensate for the lack of propulsion force on the passive ankle during toe-o. The plantar exors were absent on the amputated leg, hence the quadriceps assisted in achieving the toe-o position. The hip forms the basis for human gait as it generates the power for all the movements. Previous ndings also suggested that the excess hip abduction and exion will result in unstable lumbar movements leading to a decrease in gluteus medius and internal oblique activity [230].
The changes in hip exion-extension angle as the gait cycle progresses is illustrated in Figure 5.2. There was an early swing phase due to limited stance phase with regard to the amputated leg, this is illustrated in Figure 5.2 and also compliments the values of the stance phase in Table 3, although this was insigni cant (p = 0.562). The basic function of the lateral pelvic tilt (pelvic obliquity) is to control the vertical excursion of the centre of mass. The lateral pelvic tilt is important as it lowers the centre-of-mass, hence causing a decrease in energy expenditure [231]. Similar eects of pelvic tilt on acetabular retroversion were also reported in literature [232]. The poor range of motion of the hip during exion and extension as illustrated, in Figure 5.2, is as a result of poor propelling force during heel-o and toe-o events. The mechanical ankle is sti and provides limited propulsive force during these events. Therefore, the excessive hip tilt has adverse short and long term eects to the amputee's health and well-being.
The lack of hip extension during the terminal stance, as illustrated in Figure 5.2 was mainly attributed to lack of balance which is usually provided by the dorsi exors and plantar exors: these were absent on the amputated limb. The hip rotation in Figure 5.2 exhibited poor range of motion during the stance phase with a 60% deviation from the mean of the normal subjects. The poor hip rotation was due to the dragging of the aected limb resulting in the circumduction of the hip.
Figure 5.2: Hip exion-extension and pelvis during normal sagittal gait for amputee participant 1 (a) and Amputee participant 2 (b). The grey band is the normative data: mean ± 1 SD, the red line is the amputated leg and green line is the intact leg
The mean hip exions during the terminal stance for both subjects were well above 420.
This was as a result of the amputee leaning on the unaected leg so as to create su cient heel clearance. This may result in posture challenges since posture is adapted from the hip. Taking into consideration that the hip exors connect the femur to the lower back, the excessive use of the exors may result in lower back pain [231]. The limited range of motion in hip extension contributes much to the long term development of anterior pelvic tilt [233]. Thus, there is a need to have a complete range of motion for the hip exion and extension as a way of postural
correction and of limiting the causes of lower back pain [234].
The mean pelvis tilt for the amputees was 180, that is a 50% deviation from the expected
mean of 120. The expected pelvic tilt is 50 to 180 with 120 being the average and 180 being
the worst case scenario in most cases. During normal gait, the pelvis was expected to oscillate along the z-axis between −20 and 60. However, due to the compensation within the hip exion,
the oscillation shifted gradually towards the 20. The pelvis tilt is mostly used to identify the
possibility of gait anomalies and their relationship to short and long term eects on the spinal cord [235]. The lateral pelvis tilt is controlled by the hip abductors of the stance leg. As a result, there is notable change in lateral tilt, as shown in Figure 28 when the amputated leg is within the stance phase.
Pelvis tilt is an important parameter to determine the possibility of lower back pain in the future for most amputees. The most important parameter is the pelvic anterior tilt; however, during the gait, only the pelvis pitch, roll and yaw were monitored as orientation parameters. Therefore, the tilt was derived from these three parameters. The status of the hip exors and the hamstrings determine the level and type of tilt which may result. The hip exors were aected by the excessive hip exion and hip extension, the hamstrings were aected mostly by the excessive knee extension and exion. However, the extreme changes in the pelvis aect the degree of lumbar lordosis [236]. This may result in lower back pain. Similar results were recorded on trunk kinematics analysis revealing an asymmetrical gait due to compensatory mechanisms based on 54 subjects [237]. These compensatory mechanisms put the lumbar spine under immense stress [232].
The knee exion-extension analysis
The knee angle has a direct eect on velocity and other distance variables. The status of the knee angle is illustrated in Figure 5.3 during the normal gait. The knee angle achieved the expected range of motion for the amputees. However, these achievements were as a result of excessive knee exion rotation. The results presented in Figure 5.3 show that there is neither excessive extension nor excessive exion of the knee. However, during the early stance there is evidence of extension when exion is expected and also a poor range of motion with respect to knee rotation.The poor range of motion and excessive rotation has an adverse eect on the patellofemoral joint [238]. The principle role of the knee is to allow movement and stability by absorbing, transmitting and redistributing forces. However, such excessive rotations will aect the performance of the knee resulting in long term health eects related to joints.
Figure 5.3: (a) Knee extension- exion during normal gait for amputee participant 1 and (b) amputee participant 2. The grey band is the normative data: mean ±1SD, the red line is the amputated leg and green line is the normal leg
The ankle dorsi exion and plantar exion analysis
The ankle dorsi exion and plantar exion plays an important role during the toe-o period and assisted in moving the body mass forward, as shown in Figure 5.4.
Figure 5.4: (a) Ankle movements for amputee participant 1 and (b) Amputee participant 2. The grey band is the normative data: mean ±1 SD, the red line is the amputated leg and the green line is the normal leg
However, in amputees the passive ankle does not achieve much. The amputated leg does not even achieve 25% of the expected range of motion during both the dorsi exion and the plan- tar exion for both amputees, as shown in Figure 5.4. As a result, the intact limb of the subject will be used to develop compensation for moving the mass of the body forward. The plantar ex- ion movement experienced during heel-strike provides balance and smooth tibia projection. Using
the phases suggested by [229], the lack of power absorption between 5% - 40% of the stride at the ankle on the amputated leg results in the excessive hip exion on the normal leg due to con- traction of the plantar exors. The absence of the exible and adaptive ankle results in the lack of power generation by the ankle during the 40% - 60% stride cycle. When addressing long term eects associated with asymmetrical gait, foot alignment is a major contribution, hence the ankle dorsi exion and plantar exion angles should be monitored [239]. The lack of power absorption from initial contact to approximately 15% of the stride on the amputated leg results in the ab- sorption of power within the socket, thereby increasing the risk of bruises on the contact surfaces. The minimum, maximum and mean angles of the amputee have exceeded those of the non- amputated leg in all respects, as shown in Figures 5.5. The results revealed that the mechanical prosthetic ankle was exceeding expected angles as a way of compensating for poor hip exion and extension as the lack of intuitive control of the device leaves the user with no control over range of motion. The four-bar ankle linkage model suggested by, [240], shows that the calcane- o bular ligament and the tibiocalcaneal ligament provides stable and controlled dorsi exion for the proper positioning of the ankle as a result of the Tibialis Anterior muscle activation.
The subtalar movement permits the foot to change from being exible to a rigid structure during normal gait, thereby assisting as a rigid force lever for force transition. However, the passive ankle could not achieve such transformation, as illustrated in Figure 5.5, during the early swing phase at 58% gait. The double support occurred late in the gait cycle, as illustrated in Figure 5.4. The decrease in the double support time shows evidence of notable balance disorder. The step duration varied depending on the side being measured; however, comparing the amputated leg and the intact leg, the duration decreased with respect to the amputated leg.
5.4 Limitations of the study
The amputees were using prosthetic limbs manufactured from two dierent suppliers. The extent to which this has impact on the amputee's gait was not evaluated as the devices could not be interchanged between the amputees as they were custom made. However, same Prosthetist was used to t the passive prosthetic limbs and his recommendation was that they were t for purpose. Even though the amputees did not report any di culties on using the prosthetic limbs, the swelling of the stump was noted at the end of every two hour walking session. The sample of participants was small (two amputees), however, the number of gait activities was increased per amputee and the number of days were increased in an eort to increase the number of data sets for statistical analysis. A comparison was made to previously completed studies [241], [242], [243]. The small sample has an eect on diagnosis studies, but has less signi cance when a patient speci c device is being designed. In this study, the analysis of the signals were only meant for determination of signal quality in order to develop design speci cations.
5.5 Conclusion
The experimental study revealed that the use of sti mechanical prosthetic ankles causes asym- metries in the gait cycle leading to possible lower back pain injuries and bruises on the load bearing points of the socket. There is no signi cant variation in stride duration, cadence and step size between the amputee data and the normative results. This may have been attributed on the fact that the amputees were well aware that they were under investigation, resulting in proper movement within the laboratory. The amputees had a great desire to achieve walking ca- pabilities, hence they usually achieved these distance variables regardless of the type of arti cial limb being used. However, the short term eects of such eorts are bruises at the load bearing surfaces of the residual stump.
There is poor range of motion (ROM) management on the knee exion/extension and also on the hip extension. This can aect the psoas muscles which are directly connected to the ve lumbar vertebrae. Therefore, this may also result in anterior pelvic tilt. The long term eects of anterior pelvic tilt includes hip pain, lower back pain and at foot on the non-amputated leg. The fundamental principle is that the lumbar is aected by the action of the hip exors which are monitored by means of the hip exion angle. This could have an adverse eect on the anterior pelvic tilt which may lead to long term eects of lower back pain. Furthermore, the continual gait asymmetries may lead to di culties in walking, thereby causing signi cant health problems. The asymmetries in gait may be due to injury, pain or tissue damage that may result in musculoskeletal, cardiovascular, pulmonary and physiological problems for the amputees.
Therefore, there is a need to develop an active prosthetic limb which is capable of achieving the desired dorsi exion and plantar exion movements intuitively. The active ankle should have a control architecture that is capable of predicting the position of the limb in space during the gait cycle and ultimately minimising anomalies in the gait cycle. Robustness and adaptability should be the key characteristics of the control system of the active prosthetic ankle so as to adapt to changes in body weight, height and posture of the amputee, thereby minimising deviations from normal gait.
Chapter 6
Development of an activity prediction
system
Published as conference proceedings.1
6.1 Introduction
The functionality of a device is regarded as its suitability for a speci c task [244]. However, the means of determining the suitability of prosthetic limbs in developing nations has currently been a challenge. Most prosthetic lower limbs in developing countries are not tested using proper gait analysis equipment once tted. This is due to the cost associated with the analysis and the limited availability of the resources in some of the remote areas [7]. As a result, it is assumed that once the prosthetic limb ts well on the amputee, then it will perform as expected.
The leg is regarded as the lower limb of the human being [245] and by monitoring the seg- ments of the leg, one can monitor the activity of the human being. It has long been reported that active powered prosthetic limbs use sEMG signals and orientation sensors to predict the human intention [246]. However, the clinical applicability of this technology is limited due to lack of sensor fusion techniques and proper supporting hardware. At times the reported results were