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Effect of the Weight of Prosthetic Components on the Gait of Transtibial Amputees

Bateni, Hamid PhD, CPO; Olney, Sandra J. PhD

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JPO Journal of Prosthetics and Orthotics: October 2004 - Volume 16 - Issue 4 - p 113-120
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There are more than 110,000 lower limb amputations performed annually in the United States with approximately 30,000 of these being transtibial amputees (TTA).1 As a greater variety of prosthetic parts appear on the market, selection of the appropriate components is becoming a more difficult task for clinical team members. Many researchers have reported that unilateral TTAs walk asymmetrically and in a different way from nonamputee individuals.2–7 The type of prosthetic foot,3,4,8–11 suspension system,12 the amount of relative axial movement of the residual limb inside the prosthetic socket,12,13 the level of amputation,14–17 the length of the residual limb,18,19 and the prosthetic alignment17,20 are all known to affect the characteristics of an amputee’s gait.

A factor that has become the core of competition among prosthetic component designers and manufacturers in the last decade has been the weight of the components. Although the effect of prosthetic weight on the amputees’ gait and energy expenditure has been studied previously,21–24 a comparison of available components that affect the prosthetic weight has never been made. Such comparison can reveal more practical information on the possible advantage of lighter components for amputees.

There are different measurements that help in understanding the effect of prosthetic weight on energy expenditure of TTA gait. In addition to the inverse dynamic approach that is suggested by many to be a useful tool to evaluate gait,25–31 it has been suggested that measures of energy expenditure can be used to determine the efficiency of walking.32 It is well documented that amputees spend more energy for walking compared with nonamputees.18,33–35 Studies indicate that TTAs spend 9%,33 20%,36 25%,18 and even as high as 41%37 or 42%34 more energy for walking when compared with the able-bodied. Two commonly used methods of assessing metabolic energy cost are measurement of the rate of oxygen uptake from inspired air and calculating the physiological cost index (PCI). PCI measurement, first introduced by MacGregor,38 is a computation based on heart rate and walking speed. Although the oxygen uptake method has been shown to be a reliable method and is used by many,39–42 the instruments are cumbersome, expensive, and not available in many clinics.43 In contrast, heart rate measurement is an easy and more flexible method but can be affected by various factors such as disease and diet. There is controversy over validity and reliability of the PCI as an outcome measure for the assessment of energy expenditure.44 However, despite certain limitations, many reports have shown that heart rate monitoring is an accurate, reliable, and convenient method of determining energy expenditure.43,45–47

In this study, a set of steel lower limb modular parts, including socket adaptor, upper modulus adaptor, modulus, lower modulus adaptor, and foot adaptor, were replaced by lighter ones (titanium-based). This changes the weight of the prosthesis by approximately 181 g. The purpose of this study was to determine whether a change in prosthetic weight as a result of using steel components versus titanium components would affect the kinetics and kinematics of gait as well as the energy expenditure (PCI measurement) of TTAs. We hypothesized that use of a lighter prosthesis would cause dramatic changes in stride length. We also hypothesized that the pattern of the knee’s relative angles during the swing phase of the affected side would change significantly because the heavier prosthesis demands more muscle activity to decelerate during the last part of the swing.

METHODS

With approval from the ethics committee at Queen’s University, five male unilateral traumatic TTAs, aged 32 to 77 with an average tibial length of 11 cm (± 5 cm) on the amputated side, were recruited. Participants’ mean height, body mass, and prosthetic mass were 179.18 cm, 83.37 kg, and 1.59 kg, respectively. The participants had all been using a modular prosthesis with steel components and a stationary attachment flexible endoskeletal (SAFE; Campbell-Childs, Inc., White City, OR) foot for at least 1 year before the test. Participants were included in the study if they had no lesion on the residual limb or any problems with their knee joint and walked with their prosthesis acceptably (as confirmed by a prosthetist) without the need of assistive devices at the time of this study. Participants were exempt if they were under any medication that could affect heart rate. Participants were asked not to drink alcohol 24 hours before testing and not to smoke or drink coffee 2 hours before the time of data collection.

INSTRUMENTATION

A two-dimensional peak motion analysis system was used with an AMTI force plate (Advanced Mechanical Technology Inc., Watertown, MA) to yield spatial, temporal, kinematic, and kinetic data. To ensure consistent marker placement, an ink spot was made on the participants’ skin. These points were used for reapplying the markers for a second visit. Force plate data were collected at 600 Hz frequency. The video cameras (60 Hz) were synchronized with the data acquisition from the force plate through an externally triggered pulse. To disassemble the prosthesis, change the components, and reassemble it while maintaining alignment, a device to facilitate this was designed and made at Queen’s University (Figure 1). A Polar Vantage XL heart rate monitor (Polar Electro Inc., Lake Success, NY) was used to measure the heart rate during PCI tests.

F1-4
Figure 1.:
Adjustable foot holder (A) and a positioning ring with four reference pins (B). Base “A” was used to clamp the device to a bench. Foot holders “R”, “Q,” and “P” could be adjusted by rotating the adjustment screw “R” counterclockwise. The washer “Q” and a 2.5-cm location plug “P” were used on the end of the thread “R.” This allowed positioning and clamping of the foot as well as maintaining the accessibility of the prosthetic foot adaptor screw using an Allen wrench. The upper support foot holder “C” could be adjusted by a series of vertical holes in support area “E.” This part was locked into position using two 1/4-inch cap head screws. After stabilizing this part, the foot was firmly clamped between the washer “Q,” plug “P,” and the upper support yoke “C” using the vertical adjustment screw “R.” The pin plate “O” was positioned on the posterior pin of the foot adaptor to make it more stable and prevent any rotation in foot adaptor when screwing the foot screw. The pin plate “O” was locked in place using two 1/4-inch cap screws. After the foot was firmly fastened, the height stop “F” was placed on the vertical column “E.” Then the reference ring units “G” and “L” were placed on the column (“L” ring connected to base “G”). A 3/8-inch “J” adjusting cap screw was used to adjust the ring on the socket. After tightening the rings, the foot 1/8-inch reference pins “H” and “K” were pushed toward the center until it hit the socket wall and a trace was made on it. The four sliding plastic sleeves “M,” which were on the reference pins outside the reference ring, were moved to the reference ring. This maintained the distance of each reference point from the ring when putting the prosthesis in the jig for the second time. The socket and the components were removed by loosening the foot screw, whereas the foot was still firmly held by the jig. When replacing the components, they were put back again and realigned by adjusting the socket and foot adaptor until the four reference pins hit the same place on the socket that they had contacted before.

TESTING PROCEDURE

Reflective markers were taped on the fifth metatarsal joint, the ankle lateral malleolus, the lateral epicondyle of the femur, the greater trochanter at the hip joint level, and the base of the neck at C7–T1.25 During the tests, participants were wearing a pair of shoes with an appropriate heel height as suggested by the manufacturer. Foot markers were placed over the shoes. On the prosthetic side, markers were placed on the shoe at similar positions as on the unaffected side. Participants were tested under two different conditions with two randomly assigned trials during a single session; the first condition used a steel pylon, the second a titanium pylon. In each trial, they were asked to walk on a walkway and pass over the hidden force plate three times for each side of the body. Then the sensor of the heart rate monitor was attached to each participant, over the heart region, as instructed in the device manual. The heart rates were recorded during both a 5-minute rest and the following 5 minutes of walking at the participant’s self-selected speed. Heart rate monitoring trials were repeated twice. Average heart rate was recorded for each 15-second interval over the last 2 minutes of each trial. The distance walked by each participant was measured for each 5-minute trial. After the first session, the socket adaptor (with pyramid), tube clamp adaptors on both sides, and the foot adaptor were replaced by either steel or titanium components. By replacing prosthetic parts, the prosthetic weight was changed by 181 g. After the prosthetic parts were changed, participants were allowed 2 hours to become familiar with the new components. After the tests, participants were asked, in the form of a self-administered questionnaire, to compare and record their general impression of any differences in prosthetic performance when using steel or titanium components.

PROCESSING AND ANALYSIS OF DATA

Analysis of kinetic and kinematic data have been reported in detail elsewhere.2 Stride length, speed of walking, percent of double support, stance percent, range of foot segment angle, range of knee and hip relative angles during stride, range of knee relative angles during swing phase, peak knee power at K3 and K4, maximum hip power at H1 and H3, hip positive work at H1 and H3, hip positive work during stride, total positive and negative work in lower limb during a gait cycle, PCI, and participants’ perceptions of change were measured (for more details on variable definitions, refer to Win-ter,28–30,48 Winter and Sienko,31 and Bateni and Olney2). PCI was calculated per MacGregor49 using resting heart rate, walking heart rate, and velocity as follows:

where PCI = physiological cost index (beats/m); HW = heart rate while walking (beats/min), steady-state; HR = heart rate at rest (beats/min), steady-state; and S = average speed (m/min).

Kinematic, kinetic, and PCI variables were analyzed using repeated-measures analysis of variance and differences were judged to be significant at p <.05.

RESULTS

Tables 1, 2, and 3 display selected temporal, kinetic, and kinematic variables of the amputees’ gaits. The mean of selected values of gait parameters during the gait laboratory test shows that the stride length (p = .940), stride time (p = .609), speed of walking (p = .868), and overall double support (p = .551) are nearly the same for titanium and steel trials during laboratory tests. Although the stance percent did not change substantially when components were replaced (p = .674), it was significantly different between affected and unaffected sides (p = .006). In contrast to the results obtained in the gait tests, the speed of walking during PCI tests tended to be greater when the steel components were replaced with titanium (from 1.06 to 1.18 m/s, p = .026). It is to be noted that both the distance and time of walking was greater (5-minute walking) during PCI tests than during gait laboratory trials.

T1-4
Table 1:
Mean values of all trials for selected gait variables on the affected and unaffected sides (n = 5)
T2-4
Table 2:
Mean values of all trials for selected gait kinematic variables on the affected and unaffected sides (n = 5)
T3-4
Table 3:
Average values of all trials for selected gait kinetic variables on the affected and unaffected sides (n = 5)

Figures 2, 3, and 4 show typical relative angle, net moment, and power generation/absorption profiles for the hip, knee, and ankle for affected and unaffected sides during walking with steel and titanium components. All figures also indicate mean normal values according to Winter.30,48 No substantial differences in relative angles were noted between steel and titanium trials. The affected knee relative angle during the whole stride and during swing was larger than that of the unaffected side for both trials, but the difference was not significant (p = .089, p = .090, respectively). The range of affected hip relative angles during the stride was smaller than that of the unaffected side for both steel and titanium trials (p = .067). A lower than normal hip flexion angle at midswing of the unaffected side was evident across all participants when using steel components, but it was close to mean normal in titanium tests. During steel trials, total knee and hip range of relative angles during stride did not differ significantly from titanium trials (p >.1).

F2-4
Figure 2.:
Typical relative angles from one of the participants in the affected and unaffected sides. A lower-than-normal hip flexion angle at midswing of the unaffected side was evident across all participants when using steel components, but it was close to mean normal in titanium tests.
F3-4
Figure 3.:
Typical extension and flexion moments from one of the participants in the affected and unaffected sides. The hip moment on the affected side was closer to mean normal during stance phase when using titanium components but closer to mean normal in swing phase when using steel components. The affected knee extensor moment in early stance was lower than mean normal when using steel and titanium. In addition, the knee extensor moment on the affected side, when using titanium, was higher than both normal and steel trials.
F4-4
Figure 4.:
Typical power generation and absorption from one of the participants in the affected and unaffected sides. The peak power absorption at K4 on the unaffected side was larger than the affected side for both steel and titanium trials. When using steel components, the affected side showed a higher peak power than the unaffected side at H1. It also showed a lower peak power when using titanium components. The maximum hip power at H3 was slightly greater on the unaffected side than the affected side in steel trials.

The hip moment on the affected side was closer to mean normal during stance phase when using titanium components, but closer to mean normal in swing phase when using steel components. The affected knee extensor moment in early stance was lower than mean normal when using steel and titanium. In addition, the knee extensor moment on the affected side, when using titanium, was higher than both normal and steel trials. Although the peak power absorption at K4 on the unaffected side was larger than the affected side for both steel and titanium trials, no significant difference was observed between these two variables (p = .393). When using steel components, the affected side showed a higher peak power than the unaffected side at H1. On the other hand, it showed a lower peak power when using titanium components. However, none of these differences was significant (p >.1). The maximum hip power at H3 was slightly greater on the unaffected side than the affected in steel trials but the difference was not significant (p = .315).

The positive work of the hip tended to be greater when titanium components were used relative to steel. The total positive–negative work done by the lower limb was greater on the unaffected side than the affected side when using either titanium or steel components. It was also greater for both the affected and unaffected sides when using titanium components than steel components. However, none of these variables was significantly different (p >.05).

PCI data for each participant are shown in Table 4. The PCI was higher in steel trials when compared with titanium trials. The difference in PCI was not statistically significant when comparing all participants (p = .481). However, when subject H005 was excluded from our data, the difference in PCI in steel and titanium trials was significant (p = .002).

T4-4
Table 4:
Mean values of rest and work heart rates (RSHR and WSHR) and physiological cost index (PCIS) for steel trials and rest and work heart rates (RTHR and WTHR) and physiological cost index (PCIT) for titanium trials

Most of our participants (n = 4) reported differences when the prosthetic components were changed. One reported a small difference in how the limb felt but could not decide which one was definitely better. Another participant felt the prosthesis to be lighter when titanium components were used. Another participant was surprisingly impressed by the titanium components. He mentioned a quite noticeable difference when on stairs. He reported that ascending stairs was much easier and movements were much faster. He also noticed a quicker balance adjustment in a tripping or falling situation.

DISCUSSION

The results of our study show no consistent differences in the kinematic and kinetic factors of the amputee gait when the steel prosthetic components were replaced by titanium. It was expected that if titanium components were to make any changes in the gait, the stride length would have changed considerably, because it is known to be a sensitive variable for evaluating an amputee’s gait.50 The results also suggest that small changes in the weight of a prosthetic limb do not affect the stance time of the affected side. It was expected that a lighter prosthesis would cause changes in the pattern of the knee relative angles during the swing phase of the affected side. It was hypothesized that a heavier prosthesis demands more muscle activity to decelerate during the last part of swing. However, our results did not support any statistically significant difference between steel and titanium trials regarding knee relative angles.

We expected that even if no changes in hip and knee kinematic factors were observed, the change of the prosthetic weight might be reflected in the kinetic variables of the hip and knee. It has been reported that change of transfemoral prosthetic weight increases hip flexor muscle concentric work in the acceleration phase of the swing and helps transfemoral amputees recover the lost mechanical energy during the deceleration phase.22 However, when the prosthetic components were changed in our study, the hip and knee moment for either the affected or unaffected side was not changed significantly among our participants. The K1 and K2 bursts of power on the affected side were absent or near zero in all participants. K3, an absorption phase, was higher than mean normal in all participants but close to mean normal in two. These findings do not agree with those of Gitter and colleagues51 who reported a K3 within the normal range for TTAs. A large K3 is usually associated with a need for increased stability and is often seen in the elderly48 and in the presence of many pathologies.25,52 The large K3 also seen on the unaffected side can likely be attributed to the same cause. Because K4 is known to be a power absorption response to a swinging of the leg, it was expected to show a lower power peak when the lighter components were used. However, no changes were seen when the components were replaced. If the prosthetic components were to make any differences in gait, a smaller H3 would be expected when using lighter components. However, no statistically significant change was seen at H3 of the affected side between titanium and steel trials, likely because differences in weight were not large enough.

Our PCI findings are very close to those of Gailey et al.19 who reported that TTAs with relatively light prostheses (≤2.27 kg) showed mean resting heart rate, walking heart rate, speed of walking, and PCI to be 85.9 beats/min, 103.8 beats/min, 67 m/min, and 0.267 beats/min, respectively (Table 4). In contrast to the findings of others,23,24,53 our results indicate a significant difference in energy expenditure when steel components were replaced with titanium. The study by Rose and coworkers54 showed that changes in walking speed directly affected PCI. In fact, the observed changes in PCI in our study could have occurred as a result of faster walking speed during the PCI test when using titanium components as compared with steel components. The discrepancy between our findings and others could arise from the major difference between our test setup and those from other studies. In other studies, a heavier prosthesis is usually simulated through the addition of external weights to an existing prosthesis. However, external weight might not be representative of a realistic situation (i.e., use of lighter components) like we had in our study. Unfortunately, previous studies do not characterize change of prosthetic weight as a result of change of components rather than the addition of external weight. Indeed, our study is the only study that looks at a change of weight that TTAs could experience in real life.

The low PCI when using titanium components, as indicated by our results, suggests that participants were performing a more economic gait when using the light materials. One might ask why these clear changes in PCI were not similarly reflected in laboratory gait tests. Many previous studies have shown low correlations between the mechanical and metabolic findings,50,55 frequently reporting differences in metabolic measures but no differences in mechanical measures.55,56 However, there are several possible explanations for this difference in our results. First, the amputees were less constrained and more able to adopt their free walking speed during the lengthy 5-minute PCI trials and could have walked in a manner that differed between steel and titanium trials in the PCI test but were not discernible in the short gait laboratory test. Second, kinetic and kinematic changes could have been present, but the small sample of strides was insufficient for such differences to be seen. In fact, only approximately 18 seconds of gait laboratory data were analyzed, whereas PCI calculations were based on 5 minutes of data. Third, there could have been no substantial changes in the kinematics and kinetics, but the lighter components resulted in a lower expenditure of energy. In other words, TTAs preferred to keep the pattern of their walking similar to what they had used in past years, even with the lighter components. One cannot exclude the possibility of further changes to the gait that would result from biomechanical or metabolic economies if use was extended over a longer period.

Some of our participants reported more satisfaction when the prosthetic components were changed from steel to titanium. In a recent study, Hafner et al.57 reported that a participant’s perception of changes in prosthetic feet could be a valuable measurement in which experimental setup has limited clinical or statistical significance. However, it is to be noted that generalizability of our findings is limited as a result of the small number of participants.

CONCLUSION

It appears that prosthetic weight at the level of approximately 180 g does not systematically influence the kinetics and kinematics of TTA gait over a short time. However, the PCI was shown to be significantly reduced with lighter components within a 5-minute walk as demonstrated in this study. Our study suggests that TTAs do not change their walking pattern, in the short-term, with a change in prosthetic weight. Our results do not support the choice of the more expensive titanium over steel components; the results of PCI analysis warrant further study.

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Keywords:

amputee; biomechanics; gait; kinematics; kinesiology; kinetics; physiological cost index; prosthesis; transtibial amputation

© 2004 American Academy of Orthotists & Prosthetists