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

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

ABSTRACT

The purpose of this study was to determine the effect of different weights of modular prosthetic components (steel and titanium) on the gait of transtibial amputees. Gait parameters of five male unilateral transtibial amputees between the ages of 32 and 77 years were tested in a laboratory setting using both steel and titanium components. The kinetic and kinematic gait parameters were studied using a four-segment link segment inverse dynamic model. The physiological cost index (PCI) was determined from measures of walking speed and heart rate obtained from a monitor. Participants were also asked for their perception of changes. The following variables were compared for the two situations: linear velocity, relative angles, net moment, power and work at the hip, knee and ankle joints, and PCI. Kinematics and kinetic differences from mean normal were not consistent. Differences were seen between trials with different components. The PCI was greater when using steel components than when using titanium for all participants (p = 0.002). 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 = 0.026). Participants' perceptions were inconsistent. The results of the kinematic, kinetic, and subjective reports do not support the choice of the more expensive titanium over steel components; the results of PCI analysis warrant further study. ( J Prosthet Orthot. 2004;16:113–120.)

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 outcomemeasure 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.

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 Winter, 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:

PCI = ( HW - HR ) / S

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 < 0.05.

RESULTS

Table 1 , Table 2 , and Table 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 = 0.940), stride time ( p = 0.609), speed of walking ( p = 0.868), and overall double support ( p = 0.551) 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 = 0.674), it was significantly different between affected and unaffected sides ( p = 0.006). 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.

Figure 2 , Figure 3 , and Figure 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 = 0.089, p = 0.090, 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 = 0.067). 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 > 0.1).

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 = 0.393). 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 > 0.1). 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 = 0.315).

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 > 0.05).

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 = 0.481). However, when subject H005 was excluded from our data, the difference in PCI in steel and titanium trials was significant ( p = 0.002).

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 (less than or equal to 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.

Correspondence to: Hamid Bateni, PhD, CPO, Tennessee State University, Box 1104, 3500 John A. Merritt Blvd., Nashville, TN 37209-1561; e-mail: hbateni@tnstate.edu .


HAMID BATENI, PhD, CPO, is an assistant professor, Department of Physical Therapy, Tennessee State University, Nashville, TN. Hamid Bateni was supported by Queen's Graduate Award during this study.

SANDRA J. OLNEY, PhD, is Professor and Director, School of Rehabilitation Therapy, Queen's University, Kingston, Ontario, Canada.

References:

  1. Bodily KC, Burgess EM. Contralateral limb and patient survival after leg amputation. Am J Surg 1983;146:280–282.
  2. Bateni H, Olney SJ. Kinematic and kinetic variations of belowknee amputee gait. J Prosthet Orthot 2002;14:1–8.
  3. Doane NE, Holt LE. A comparison of the SACH and single axis foot in the gait of unilateral below-knee amputees. Prosthet Orthot Int 1983;7:33–36.
  4. Culham EG, Peat M, Newell E. Below-knee amputation: a comparison of the effect of the SACH foot and single axis foot on electromyographic patterns during locomotion. Prosthet Orthot Int 1986;10:15–22.
  5. Engsberg JR, Lee AG, Patterson JL, Harder JA. External loading comparisons between able-bodied and below-knee-amputee children during walking. Arch Phys Med Rehabil 1991;72: 657–661.
  6. Prince F, Allard P, Therrien RG, McFayden BJ. Running gait impulse asymmetries in below-knee amputees. Prosthet Orthot Int 1992;16:19–24.
  7. Robinson JL, Smidt GL, Arora JS. Accelerographic, temporal, and distance gait factors in below-knee amputees. Phys Ther 1977;57:898–904.
  8. Goh JC, Solomonidis SE, Spence WD, Paul JP. Biomechanical evaluation of SACH and uniaxial feet. Prosthet Orthot Int 1984; 8:147–154.
  9. Wing DC, Hittenberger DA. Energy-storing prosthetic feet. Arch Phys Med Rehabil 1989;70:330–335.
  10. Wirta RW, Mason R, Calvo K, Golbranson FL. Effect on gait using various prosthetic ankle-foot devices. J Rehabil Res Dev 1991;28:13–24.
  11. Mizuno N, Aoyama T, Nakajima A, et al. Functional evaluation by gait analysis of various ankle-foot assemblies used by belowknee amputees. Prosthet Orthot Int 1992;16:174–182.
  12. Wirta RW, Golbranson FL, Mason R, Calvo K. Analysis of belowknee suspension systems: effect on gait. J Rehabil Res Dev 1990;27:385–396.
  13. Erikson U, Lemperg R. Roentgenological study of movements of the amputation stump within the prosthesis socket in belowknee amputees fitted with a PTB prosthesis. Acta Orthop Scand 1969;40:520–529.
  14. Sanders GT, May BJ. Lower Limb Amputations: A Guide to Rehabilitation . Philadelphia: FA Davis; 1986.
  15. Waters RL, Perry J, Antonelli D, Hislop H. Energy cost of walking of amputees: the influence of level of amputation. J Bone Joint Surg [Am] 1976;58:42–46.
  16. Pinzur MS. New concepts in lower-limb amputation and prosthetic management. Instruct Course Lect 1990;39:361–366.
  17. Smidt GL. Gait in Rehabilitation. Clinics in Physical Therapy . New York: Churchill Livingstone; 1990.
  18. Gonzalez EG, Corcoran PJ, Reyes RL. Energy expenditure in below-knee amputees: correlation with stump length. Arch Phys Med Rehabil 1974;55:111–119.
  19. Gailey RS, Wenger MA, Raya M, et al. Energy expenditure of trans-tibial amputees during ambulation at self-selected pace. Prosthet Orthot Int 1994;18:84–91.
  20. Hannah RE, Morrison JB, Chapman AE. Prostheses alignment: effect on gait of persons with below-knee amputations. Arch Phys Med Rehabil 1984;65:159–162.
  21. Godfrey CM, Brett R, Jousse AT. Foot mass effect on gait in the prosthetic limb. Arch Phys Med Rehabil 1977;58:268–269.
  22. Gitter A, Czerniecki J, Meinders M. Effect of prosthetic mass on swing phase work during above-knee amputee ambulation. Am J Phys Med Rehabil 1997;76:114–121.
  23. Czerniecki JM, Gitter A, Weaver K. Effect of alterations in prosthetic shank mass on the metabolic costs of ambulation in above-knee amputees. Am J Phys Med Rehabil 1994;73: 348–352.
  24. Lehmann JF, Price R, Okumura R, et al. Mass and mass distribution of below-knee prostheses: effect on gait efficacy and self-selected walking speed. Arch Phys Med Rehabil 1998;79: 162–168.
  25. Olney SJ, Griffin MP, Monga TN, McBride ID. Work and power in gait of stroke patients. Arch Phys Med Rehabil 1991;72: 309–314.
  26. Olney SJ, Griffin MP, McBride ID. Temporal, kinematic, and kinetic variables related to gait speed in subjects with hemiplegia: a regression approach. Phys Ther 1994;74:872–885.
  27. Olney SJ, Bateni H, Griffin MP, McBride ID. Gait adaptations in below knee amputees and effect of weight change in prosthetic components: a pilot study. Paper presented at XVI Congress of the International Society of Biomechanics; Tokyo, Japan; August 25–29, 1997.
  28. Winter DA. Energy generation and absorption at the ankle and knee during fast, natural, and slow cadences. Clin Orthop 1983; 175:147–154.
  29. Winter DA. Biomechanical motor patterns in normal walking. J Mot Behav 1983;15:302–330.
  30. Winter DA, ed. The Biomechanics and Motor Control of Human Gait . Waterloo: University of Waterloo Press; 1987.
  31. Winter DA, Sienko SE. Biomechanics of below-knee amputee gait. J Biomech 1988;21:361–367.
  32. Rose J, Medeiros JM, Parker R. Energy cost index as an estimate of energy expenditure of cerebral-palsied children during assisted ambulation. Dev Med Child Neurol 1985;27:485–490.
  33. Huang CT, Jackson JR, Moore NB, et al. Amputation: energy cost of ambulation. Arch Phys Med Rehabil 1979;60:18–24.
  34. Ganguli S, Datta SR. Prediction of energy cost from peak heart rate in lower extremity amputees. Biomed Eng 1975;10:52–55.
  35. Hunter D, Smith Cole E, Murray JM, Murray TD. Energy expenditure of below-knee amputees during harness-supported treadmill ambulation. J Orthop Sports Phys Ther 1995;21: 268–276.
  36. Molen NH. Energy-speed relation of below-knee amputees walking on a motor-driven treadmill. Internationale Zeitschrift fur Angewandte Physiologie Einschliesslich Arbeitsphysiologie 1973;31:173–185.
  37. Fisher SV, Gullickson G. Energy cost of ambulation in health and disability: a literature review. Arch Phys Med Rehabil 1978; 59:124–133.
  38. MacGregor J. The objective measurement of physical performance with long term ambulatory physiological surveillance equipment (LAPSE). In: Stott FD, Reftery EB, Goulding L, eds. Proceedings of 3rd International Symposium on Ambulatory Monitoring . London: Academic Printing; 1979:29–39.
  39. Waters RL, Lunsford BR, Perry J, Byrd R. Energy-speed relationship of walking: standard tables. J Orthop Res 1988;6: 215–222.
  40. Ijzerman MJ, Baardman G, van 't Hof MA, et al. Validity and reproducibility of crutch force and heart rate measurements to assess energy expenditure of paraplegic gait. Arch Phys Med Rehabil 1999;80:1017–1023.
  41. Bowen TR, Lennon N, Castagno P, et al. Variability of energyconsumption measures in children with cerebral palsy. J Pediatr Orthop 1998;18:738–742.
  42. Boyd R, Fatone S, Rodda J, et al. High- or low-technology measurements of energy expenditure in clinical gait analysis? Dev Med Child Neurol 1999;41:676–682.
  43. Rose J, Gamble JG, Medeiros J, et al. Energy cost of walking in normal children and in those with cerebral palsy: comparison of heart rate and oxygen uptake. J Pediatr Orthop 1989;9:276–279.
  44. Ijzerman MJ, Nene AV. Feasibility of the physiological cost index as an outcome measure for the assessment of energy expenditure during walking. Arch Phys Med Rehabil 2002;83: 1777–1782.
  45. Hood VL, Granat MH, Maxwell DJ, Hasler JP. A new method of using heart rate to represent energy expenditure: the Total Heart Beat Index. Arch Phys Med Rehabil 2002;83:1266–1273.
  46. Ghosh AK, Tibarewala DN, Chakraborty S, et al. An improved approach for performance evaluation in lower extremity involvement. J Biomed Eng 1980;2:121–125.
  47. Butler P, Engelbrecht M, Major RE, et al. Physiological cost index of walking for normal children and its use as an indicator of physical handicap. Dev Med Child Neurol 1984;26:607–612.
  48. Winter DA, ed. Biomechanics and Motor Control of Human Movement , 2nd ed. New York: Wiley; 1990.
  49. MacGregor J. Rehabilitation ambulatory monitoring. In: Kenedi RM, Paul JP, Hughes JP, eds. Disability: Proceedings of a Seminar on Rehabilitation of the Disabled in Relation to Clinical and Biomechanical Aspects, Costs, and Effectiveness . London: Macmillan; 1979:159–172.
  50. Skinner HB, Effeney EJ. Gait analysis in amputees. Am J Phys Med 1985;64:82–89.
  51. Gitter A, Czerniecki JM, DeGroot DM. Biomechanical analysis of the influence of prosthetic feet on below-knee amputee walking. Am J Phys Med Rehabil 1991;70:142–148.
  52. Olney SJ, MacPhail HE, Hedden DM, Boyce WF. Work and power in hemiplegic cerebral palsy gait. Phys Ther 1990;70: 431–438.
  53. Foerster SA, Bagley AM, Mote CD Jr, Skinner HB. The prediction of metabolic energy expenditure during gait from mechanical energy of the limb: a preliminary study. J Rehabil Res Dev 1995;32:128–134.
  54. Rose J, Gamble JG, Burgos A, et al. Energy expenditure index of walking for normal children and for children with cerebral palsy. Dev Med Child Neurol 1990;32:333–340.
  55. Colborne GR. Analysis of mechanical and metabolic factors in the gait of congenital below knee amputees. A comparison of the SACH and Seattle feet. Am J Phys Med Rehabil 1992;71: 272–278.
  56. Bach TM, Barnes LJ, Evans OM, Robinson IGA. Optimization of internal characteristic of transfemoral limb prostheses: tests of predictions of a computer simulation. In: Proceedings of the Eighth Biennial Conference and Human Locomotion Symposium of the Canadian Society for Biomechanics (CSB/SCB) . Calgary: Spodym; 1994:124–125.
  57. Hafner BJ, Sanders JE, Czerniecki J, Fergason J. Energy storage and return prostheses: does patient perception correlate with biomechanical analysis? Clin Biomech 2002;17:325–344.