Biomech of Below Knee Amp

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BIOMECH ANICS OF BELOW-KNE E AMPU TEE GAITDAVID A. WINTER

Department of Kinesiology, University of Waterloo, Waterloo, Canada, N2L 3Gland

SUSAN E. SIENKOMotion Analysis Laboratory, Department of Surgery, Southern Illinois University, Springfield,IL 62708, U.S.A.

Abstract-Sagittal plane biomechanical and EMG analyses from eight below knee (B/K) amputee trialsdemonstrate considerably modified motor patterns from the residual muscles at the hip and knee. Five SACHfittings, two Uniaxial and one Gressinger prostheses were analysed. Moments of force and mechanical powerwere analysed on all eight trials and EMG profiles are reported for three of the amputees fitted with SACHprostheses. The findings can be summarized as follows:1. All eight trials had similar internal moment of force patterns at the ankle. A dorsitlexor momentcommenced at heel contact and continued for the first third of stance. The prostheses generated aplantarflexor moment for the balance of stance which increased in late stance to about 2/3 that seen innormals.2. The two Uniaxial prostheses showed a 20 I,; ecovery of stored energy which was returned at push-off.The recovery by the Gressinger fitting was 30%.3. For all but the Gressinger prosthesis the knee moment of force was negligible during early stance (whennormals have an extensor moment), below normal in late stance and fairly normal during swing. The amputeewearing the Gressinger prosthesis had a normal but slightly reduced pattern of moments of force over theentire stride.4. All eight trials had hyperactive hip extensors during early and mid-stance which resulted in above-normal energy generation by these concentrically contracting muscles. This compensation makes up for theloss of the major energy generation by the plantarflexors at push-off.5. The moment of force and power patterns at the hip for all eight trials during late stance and swing werefairly normal.6. Because of hyperactivity of the hamstrings during early stance there is an excessive knee flexor momentwhich is cancelled out by co-contracting knee extensors at that time.

INTRODUCTIONThe number of below knee amputees in our populationis increasing because of ageing and surgery related toperipheral vascular disease. Bilateral B/K amputees arealso on the increase for the same reason. Such am-putees are fitted with prostheses that have anklemechanisms with limited or no rotation which providessome ankle stability. However, they have lost theirplantarflexors which play two major roles in normalgait. These muscles, early in stance, assist in controllingthe forward rotation of the leg over the supporting footand at push-off are responsible for over 805: of themechanical power generated during the gait cycle(Winter, 1983). It is, therefore, important to analyse thegait of such amputees to see how they have adapted atthe knee and hip to compensate for this loss.

A wide variety of gait studies of amputees has beenreported and they vary drastically as to objectives andtechniques. Many studies have presented descriptions,based mainly on functional, temporal, cadence andbasic kinematic measures (Enoka et al., 1982; Hannahand Morrison. 1984; James and Oberg, 1973; Kegel et

Received December 1985; in revised form April 1987

al., 1981; Olney et al., 1979; Robinson et a/., 1977).While these outcome measures have been useful tomonitor the progress of an amputee they have notyielded new information that could lead to improve-ment of gait. One exception comes from assessmentswhere joint angle measurements aid in the alignment ofthe prosthesis. Few kinetic studies have been reportedand they report stump/socket or pylon forces duringstance (Symington et al., 1979). Studies involving themetabolic cost of amputee walking have yielded globalmeasures that reflect differences from normals atvarious walking speeds, or differences attributed tostump length (Gonzalez et al., 1974; Waters et al., 1976)or physical training (Urban, 1973). Again these de-scriptive studies have not resulted in information thatwould explain the increased metabolic cost ot would beuseful in altering the amputees gait pattern. Only threestudies have actually reported kinetic variables. Onestudy of an above knee (A/K) amputee shows thecompensating roles of the remaining ipsilateral muscu-lature and the contralateral limb (Cappozzo et al.,1976). Six A/K and six B/K amputees were analyzed(Lewallen et al., 1984) during walking and comparisonsof moments of force at the ankle, knee and hip weremade with those analysed from a control group of 17normals. One recent abstract (Miller and Munro,

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362 D. A. WINTER and S. E. SIENKO1985) summarizes some aspects of the moment of forcepatterns from both the non-involved limbs and pros-thetic limbs of running B/K amputees.

Several texts have appeared (Inman er al., 1981;Klopstegand Wilson, 1968) that address the kinetics ofgait and the implications related to the forces acting atthe stump/prosthesis interface and those muscle forcesnecessary to stabilize the prosthetic limb. However,there are no reports that have analysed in detail themotor patterns evident in either below or above kneeamputees. Such motor patterns are those related to thecause of the amputees new gait pattern and arerepresented by variables such as moments of force (atall joints), mechanical power (at all joints) and electro-myographic patterns (in the residual functionalmuscles).

Throughout all the amputee-related literature, con-tinuous references are made to variables that establishgait asymmetry. Subsequently, attempts are beingmade, without scientific justification, to force theamputee to walk more symmetrically. At the outset, theauthor would be cautious about gait retraining proto-cols which are aimed at improved symmetry based onnothing more than an idea that it would automaticallybe an improvement. It is safe to say that any humansystem with major structural asymmetries in theneuromuscular skeletal system cannot be optimalwhen the gait is symmetrical. Rather, a new non-symmetrical optimal is probably being sought by theamputee within the constraints of his residual systemand the mechanics of his prosthesis. It is the goal of thispaper to demonstrate how a number of B/K amputeeshave re-altered the motor patterns of their amputatedlimb and are walking with a considerable degree ofmotor asymmetries.

METHODOLOGY

The case studies reported here were taken from anumber of investigations that were conducted in theGait Laboratory at the University of Waterloo over thepast three years. The biomechanical analyses usedsynchronized tine and force plate data along with a sixchannel EMG telemetry system. Each of the B/Kamputees was asked to walk along a 10 m walkway inhis own footwear at his natural cadence. A total of fiveamputees were assessed and for three of them acomplete EMG profile was obtained from their re-sidual muscles. All five had SACH prostheses, two hadretrials wearing a Uniaxial foot and one of these twohad a third assessment wearing a Gressinger fitting.The complete details of data collection, processingand biomechanical analysis have been reported indetail elsewhere (Winter, 1980, 1983, 1984b), but aresummarized here. Reflective markers were located atthe following anatomical landmarks: heel, toe, fifthmetatarsal, ankle, knee (lateral epicondyle), hip(greater trochanter), and midline at the mid trunk level.Cine film data taken at 50 frames s - were digitized,calibrated, corrected for parallax, then digitally filtered

using a zero lag, fourth order Butterworth filter cuttingoff at 6 HZ (Winter et al., 1974). Subsequent kinematicanalyses yielded all necessary linear and angular dis-placements, velocities and accelerations in the plane ofprogression (Winter et al., 1974). The errors in theresultant accelerations have been shown to be in-significant when compared with directly recordedaccelerations (Pezzack et al., 1977). The r.m.s. skinmovement over the stride period of two markers de-fining any segment has been calculated to be 0.5 cm orless. Using standard link segment analyses the am-putees lower limb and prosthesis were modelled toyield the joint moments of force. Using the formula,Pi = M ,-aj, the joint muscle power, Pj, was calculated.This power is positive if the moment of force, Mj, andangular velocity, wj have the same polarity and negat-ive if they have opposite polarities. Positive power thenrepresents the rate of energy generation in a concentriccontraction and negative power the rate of energyabsorption in an eccentric contraction. In the case of anamputee, negative power would indicate the rate ofenergy absorption by any mechanism (damper, spring)built into the prosthesis and positive power wouldreflect a return of stored energy from any springmechanism.

For the three amputees on which EMG data werecollected additional trials took place in which theamputee walked back and forth on the walkway and atotal of 32 s of data were collected. This was ac-complished while he traversed the 5 m long straight-away section of the walkway. The details of the EMGbiotelemetry, processing and computer analysis havebeen reported in detail (Winter, 1984a) but are sum-marized here. Five residual muscles (gluteus maximus,biceps femoris, semitendinosis, rectus femoris andvastus lateralis) were monitored using surface elec-trodes, and the raw EMG and footswitch signals weretransmitted. Upon demodulation, the five EMGs wereprocessed in a linear envelope (LE) detector using asecond-order critically damped 3 Hz low-pass filter(Winter, 1984a). Using an interactive computer pro-gram, the operator identified heel contact and toe-offfor up to 20 strides. Computer analysis yielded anensemble average which gave the mean and standarddeviation of each of the muscles LE pattern over thestride period.

RESULTS AND DISCU SSION

All individual patient trials were compared withsuitably normalized profiles from normals walking atroughly the same cadence. It is, therefore, imperativethat these baseline curves be understood prior tocomparison with the amputee population. Figure 1presents the sagittal plane moment of force profiles forthe normal population, walking at their natural ca-dence. For the 19 subjects that formed this average, thecadence was 105 steps min- (SD. = 7.5). Stride timewas normalized to 100% and stance time to 60 %.Because of differences in body mass, each moment of


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Biomechanics of below-knee amputee gait 363

IMENT OF FORCE-NRTURRL CADENCE (N= 19 I

Fig. 1. Moment of force inter-subject averages for normalsubjects walking their natural cadence. The stride period is set100 6 and stance to 60 7;. The moments of force for eachsubject were normalized by dividing by body mass prior toensemble averaging. Solid line is the average curve, dotted lineis one standard deviation either side of the mean. The supportmoment is the algebraic sum of all three joint moments withextensor moments being positive.

force curve was normalized by dividing by body massprior to ensemble averaging of each curve. Eachensemble average curve (solid line) was calculated at2 y,; intervals over the stride period and the standarddeviation was calculated and plotted (dotted line). Theconvention is such that extensor moments are positiveand flexor negative. The support moment as definedpreviously (Winter, 1980), is the algebraic sum of allthree joint moments, and has been shown to be positiveduring stance for all cadences. Subsequent analysis ofthese moment of force patterns has shown that thereason for low variability in the positive supportmoment pattern was a consistent net moment acting onthe thigh at the proximal and distal ends plus a veryconsistent ankle moment pattern (Winter, 1984b). Suchconsistency was shown to be matched by extremely lowvariability kinematics jn the thigh, leg and foot angles.

For these same 19 subjects, ensemble averages of thejoint mechanical power profiles were calculated and arepresented in Fig. 2. The power curves are normalized tobody mass and are reported in W kg- I. Specific burstsof absorption and generation of energy are labelled andwill be summarized as follows:Al-absorption by plantarflexors as the leg rotatesforward over flat foot.

P O WE R E N/ R BS - NR T I I R RLRDE NCEN = 1 3 i

Fig. 2. Mechanical power inter-subject averages for samesubjects as in Fig. 1. Powers for each subject were normalizedby dividing by body mass prior to ensemble averaging. See

text for detailed discussion.

AZ-generation by plantarflexors (push-off) as thefoot plantarflexes prior to toe-off.

K l-absorption by knee extensors as the knee flexesduring weight acceptance.

KZ-generation by knee extensors as the kneeextends during mid stance to raise the center ofgravityof the body.

K3-absorption by knee extensors during push-offas the knee flexes prior to and after toe-off.

K4-absorption by knee flexors at end of swing totake out energy of swinging leg and foot.Hl-brief generation by hip extensors at weightacceptance as the hip extends (as knee flexes).

HZ&absorption by hip flexors to decelerate back-ward rotating thigh.

H3-generation by hip flexors as hip flexes beforetoe-off and in early swing to pull the lower limbupwards and forward; this action is now referred to aspull-off (as opposed to push-off by the plantarflexors).

Note: Hl and H2 are not well defined at naturalcadences but appear more clearly at faster cadences.General information and trends apparentfrom ull triuls

The average stride length for all eight trials was1.27 m (S.D. = 0.07 m), the average velocity was0.97 m s- 1 (S.D. = 0.08 m s ), the average cadencewas 92 steps min- (SD. = 5 steps min I), and theaverage body mass was 92 kg (SD. = 19.2 kg).Figures 3 and 4 are the moment of force and mecha-nical power profiles for the eight amputee trials. Fromthese ensemble average curves, we can see some generaltrends amongst this population group.


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364 D. A. WINTER and S. E. SIENKOM@ME N T i B O DY MA S S - B K MP U T E E S ( N = E l

0 0 0 0 0 (N 0 (D m c+% OF STRIDE

Fig. 3. Ensemble averages of moments of force of eightamputee trials. See text for detailed discussion of differences

from that seen for normals.

P OWERGC NE R R i E DND F I B S O RB E D- B Kf I P UT EE S ( Nd )I

_,! 10=60%I 0 0 0 0 c! 0N 0 ul m 0

OF STRIDEFig. 4. Ensemble averages of mechanical power of eightamputee trials. See text for detailed discussion of differences

from that seen for normals.1. Ankle moment and power patterns(a) For all amputees the ankle moment (Fig. 3) wasdorsiflexor for first 187; of stride (min = 14;/,, max= 24$, which compares with 6;/, for normals.

Normal subjects have a rapid lowering of the foot toground under control of the dorsiflexors. However,B/K amputees have rigid ankles which generate aninternal dorsiflexor moment from heel contact until theprosthetic foot is flat on the ground and it takessomewhat longer for the amputees leg to rotateforward until the foot is flat on the ground.

(b) As the center of pressure moves forward towardsthe ball of the prosthetic foot an internal plantarflexormoment is generated (Fig. 3). This was seen to rise to1.0 Nm kg- which is between 60% and 70% of thatgenerated by normals (see Fig. 1). The extra 3&40xgenerated by normals is what is required to cause activeand rapid plantarflexion of the foot that results in A2power burst.

(c) There was some energy absorption (Fig. 4) by theprosthetic foot (Al power burst) as the foot deformsslightly in dorsiflexion during mid stance. This energywas dissipated in the viscous material of the SACH footor was stored in the spring mechanism of the Uniaxialor Gressinger foot.(d) Some energy was returned during push-off bythe spring mechanism present in three of the prostheticfeet (A2 power burst). This was evident in the ensembleaveraged curve, however, the only prostheses re-sponsible for this energy conservation were theUniaxial type (which returned 20!, of their storedenergy) and the Gressinger foot (which returned 30 !,Aof its stored energy). These differences will be discussedlater in conjunction with Figs 5a and b and 6a and b.

2. Knee moment and power patterns(a) With the exception of the Gressinger fitting. the

knee moment of force was very low or near zero for thefirst half of stance, but well within the normal rangeduring the latter half of stance. At no time duringstance was there any significant extensor moment,which does not agree with that reported by Inman et al.(1981) who did not appear to have any link segmentanalyses to substantiate their statements.(b) K 1 and K2 bursts of power were missing on allamputees, but K3 was quite normal during push-offand early swing. K4 was quite small or negligible in allbut one trial, and was attributed to the low energyabsorption needed for the slowly swinging prostheticleg with a lower than normal mass.

3. Hip moment and power patterns(a) Hip moments of force (Fig. 3) for this amputee

population were highly variable, more so than the quitevariable normal population. Three of the SACH trialsshowed a completely extensor pattern during stancewhile the other two SACH and the Uniaxial trials hadabove normal flexor moments. Only the amputee withthe Gressinger fitting had a normal hip momentpattern (see Fig. 6a).

(b) The hip mechanical power patterns (Fig. 4) wereextremely variable, especially when the Hl and H2power bursts normally occur. However, the meanpattern showed a strong Hl burst as the hip extensorsshortened under tension after heel contact and gener-ated energy to propel the body forward from the rear.


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t O lEN O FORCE (SACH FOOT PROSTHESIS vs NORtl RLS)1

?O iiR GNiRBS!SACi F3OT PROSTHESIS v;. NORMt XSi

.: : - NC f WL S ( N = l 9 ): :: . . . _ . . . . . . . Qd , ~ e " ,

RNKL E

-j. OF STRIDEFig. 5. (a, b) Moment of force and power patterns for one ofthe amputees fitted with a SACH prosthesis overlaid on top ofthe inter-subject averages from normals. See text for detaileddiscussion.This energy generation appears to be one majorcompensation for the lack of energy generation by theankle plantarflexors. The H3 burst of mechanicalenergy generation by the hip flexors (pull-off power)

:NT O FORCE (GRESSIKER PROSTHESIS vs NORMALS)

0 0 0 0N 0 In m

-OF STRIDEPONER GENd@ S (CRESSIf f iER PROSTHESIS vs. NO t MlLS)

: - ~ ~ ~ ~ ~

a t ~ ~ ~ ~

: 47 _. ..:: :: :: :

% OF STRIDE -Fig. 6. (a, b) Moment of force and power patterns for theamputee fitted with a Gressinger prosthesis overlaid on top ofthe inter-subject averages for normals. See text for detaileddiscussion.

5a and b are the moment of force and mechanical

Specijc amputee assessmentsFrom the amputee assessments, representative

curves are presented for two different fittings. Figureswas quite well defined in all cases.


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366 D. A. WINTER and S. E. SIENKOpower comparisons for one of the SACH prostheses,Figs 6a and b are for the Gressinger fitting. In boththese figures the amputee profiles are superimposed ontop of those for normals, but the amputee curve timebase is adjusted by a few percent so that his toe-offcoincides with that of normals (60 y0 of stride). In Figs5a and b this resulted in reducing the stance periodfrom 68 %, and in Figs 6a and b stance was reducedfrom 64 %. Only by normalizing both stance and swingin this manner could proper comparisons be made ofrapidly changing events that occur around toe-off.

Trial WP92C represents the typical assessments ofthe SACH foot, the dashed line in Fig. 5a shows themoment of force patterns (Nm kg-) vs that seen for19 normal subjects. The dotted line either side of themean curve (solid line) represents + one S.D. Themoment of force patterns at the ankle showed aninternal dorsiflexor moment until 22 % stride, at whichtime the leg rotation allows the foot to become flat. Theinternal plantarflexor moment increases as the centerof pressure of the body moves forward of the anklejoint and reaches a maximum as body weight moves tothe ball of the foot. The knee moments were negligiblefor the first half of stance and fairly normal during thelater half of stance and into late swing. The hip momentis extensor during most of stance and falls well withinthe normal range during late push-off and well intoswing. The support moment is extensor for all of stancebut has a reduced amplitude compared with that ofnormals.Figure 5b shows the power generation/absorptionfor this SACH fitting. The ankle joint of the prosthesishad some compliance which resulted in some energyabsorption (Al) as the body weight caused a smalldorsiflexion of the ankle. A miniscule amount of energywas returned during A2 indicating the energy absorbedduring Al phase was lost as heat in the viscousmaterial. At the knee, there was insignificant energyabsorbed and generated during the Kl jK2 phases, butat toe-off the quadriceps absorbed a normal amount ofenergy (K3) as the knee flexed. K4 absorption by theknee flexors at the end of swing was quite normal. Atthe hip, the pull-off power was quite normal, but H2energy phase was missing. A small amplitude Hl phasewas seen during almost all of stance; and this was due togeneration of energy by the hip extensors contractingconcentrically. The area under this Hl phase representsenergy generated to propel the pelvis/trunk forwardfrom the rear using the residual hip extensors.Trial WP68E on the Gressinger prosthesis showedsome differences in the moment of force (Fig. 6a)compared with the SACH fitting. The ankle momenthad almost the same shape and magnitude; however, atthe knee and hip, the moments of force were almostnormal. The power profiles (Fig. 6b) show that 30 y0 ofthe energy stored by the ankle spring mechanism (Alburst) was returned at push-off (A2 burst). At the knee,Kl and K2 absorption and generation phases were notpresent, but K3 and K4 were of quite normal ampli-tude. At the hip, all four power phases were well within

the range seen on normals. In summary, it appears thatthe Gressinger fitting allowed the residual muscles ofthe knee and hip to act more normally than does theSACH foot, and it also returned some absorbed energyat the ankle making the gait stride more efficient.EMG assessments

For three of the SACH foot amputees, EMG signalswere recorded from the residual muscles: rectus fe-moris, vastus lateralis, semitendinosis, biceps femorisand gluteus maximus. The ensemble average for aminimum of 15 strides were determined for eachmuscle and plotted (dashed line) on top of similar inter-subject averages for normals (solid line + 1SD.) inFig. 7. Somewhat similar EMG profiles were evidentfor the other two amputees. The abnormal EMGpatterns can be explained by a sequence of adaptions asfollows:

(i) The hip extensors were dominant for most ofstance as evident by the hip moment of force patternand the hyperactivity in the gluteus maximus, bicepsfemoris and sensitendinosis at that time.

(ii) These hyperactive hamstrings also generatedabove normal knee flexor moments for the first 40 :/;, fstride period, and this was seen in the knee momentsbeing biased towards a flexor pattern (i.e. reduced tonear zero from a moderate extensor pattern).

(iii) The knee extensors (vastus lateralis, rectusfemoris) were also hyperactive indicating a major co-

E MG WP S O R v s N ORMF I L S_ - - - - - - - - _ _ _ _ _ _ _ _ _ECTUS FElOR-.__

4001 --__NOR,,# S(N=, 1 VRm LRTERRLIIStd. Lb.

Fig. 7. Ensemble average of EMG linear envelope signalfrom five residual muscles from one of the SACH prosthesistrials (dashed line) overlaid on top of similar inter-subjectaverages for normals (solid line for average, dotted line forone standard deoiation). See text for detailed discussion.


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contraction at the knee joint. This co-contractionagainst the hamstrings was present for first 40/, ofstride, and subsequently caused the net knee momentduring stance to be near zero. During push-off(40-60% stride), the hamstrings were less active, butthe knee extensors remained active well into swing.This resulted in a net knee extensor moment whichcontrolled the knee as it flexed at this time.

CONCLUSIONS

Based on a range of biomechanical analyses andEMG profiles on eight different B/K amputee trials,several trends appear that indicate a somewhat modi-fied motor patterns from the residual muscles at the hipand knee. Whether these motor patterns are optimal isnot known. However, it is important to recognize thatthe amputees neural system has recognized his newasymmetries and has compensated for his lost motorfunction and altered anthropometrics.

(i) All amputees, regardless of type of fitting, hadhyperactive hip extensor during early and mid-stancewhich resulted in above-normal energy generation bythese concentrically contracting muscles. This com-pensation, partially, made up for the lack of energygeneration by the plantarflexors at push-off.

(ii) This above-normal hip extensor activity wasachieved, partially, by the hamstrings which resulted inan above normal knee flexor moment which must becancelled out via co-contracting quadriceps. Duringearly and mid-stance the net knee moment was nearzero.

(iii) There appeared to be little or no change in themoment of force and power patterns at the knee andhip during late stance and swing.

(iv) All three prostheses (SACH, Uniaxial andGressinger) had similar internal moment of forcepatterns at the ankle. The dorsiflexor moment com-mencing at HC continued much longer into stance. Theplantarflexor moment reached about 2/3 that seen innormals.(v) The two Uniaxial trials showed a 20 Y; recoveryof energy at the ankle, and the one Gressinger trialshowed a 30 y0 energy recovery. Also, somewhat morenormal motor patterns were seen at the hip and kneethan that seen in the Uniaxial or SACH foot amputeeassessments.Acknowledgement-The authors acknowledge the financialsupport of the Medical Research Council of Canada (GrantMT4343) and the technical assistance of Mr Paul Guy andMr Steve Cornall.

REFERENCES

Cappozzo. A., Figura. F.. Leo, T. and Macthetti, M. (1976)Biomechanical evaluation of above-knee prostheses.Biom echanics V-A, 36&372.Enoka, R. M., Miller, D. I. and Burgess, E. M. (1982) Below-knee amputee running gait. Am. J. phys. Med. 62, 66-84.Gonzalez, E. G., Corcoran. P. J. and Rodolfo, L. R. (1974)Energy expenditure in B/K amputees: correlation withstump length. Archs. Phys. Med. Rehabil. 55, 1 I-1 19.Hannah, R. E. and Morrison, J. B. (1984) Prosthetic align-ment: effect of gait of persons with below-knee ampu-tations. Archs Phy s. Med. R ehabil. 65, 159-162.Inman, V. T., Ralston, H. J. and Todd, F. (1981) HumanWalking: Application to Lower Limb Prosthetics. Williamsand Wilkins, Baltimore.James, U. and Oberg, K. (1973) Prosthetic gait patterns inunilateral A/K amputees. Stand. J. Rehab. Med. 5, 35-50.Kegel, B., Burgess, B. M., Starr. T. W. and Daly. W. K. (1981)Effects of isometric muscle training on residual limbvolume strength, and gait of B/K amputees. Phvs. Therap.62, 1419-1426.Klopsteg, P.E. and Wilson, P. D. (1968) Human Limbs undtheir Substitutes. Hafner, New York.Lewallen, R., Quanbury. A. C., Ross, K. and Letts, R. M.(1985) A biomechanical study of normal and amputee gait.Biomechanics IX-A (Edited by Winter, D., Norman, R..Wells, R., Hayes, K. and Patla, A.). pp. 587-592. HumanKinetics, Champaign, IL.Miller, D. J. and Munro, C. F. (1985) Joint torque patterns ofB/K amputees during running stance. J. Biomechanics 18,236.Olney, S. J., Elkin, N. D., Lowe, P. J. and Symington, D. C.(1979) Ambulation profile for clinical gait evaluation.Physiother. Can. 31, l-6.Pezzack, J. C., Norman, R. W. and Winter, D. A. (1977) Anassessment of derivative determining techniques used formotion analysis. J. Biomechanics 10, 377-382.Robinson. J. L., Smidt, G. L. and Orora. J. S. (1977)Accelerographic, temporal and distance gait factors in B.IKamputees. Phys. 7her. 57, 8988904.Symington, D. C., Lowe, P. J. and Olney, S. J. (1979)Pedynograph: clinical tool for force measurement and gaitanalysis in lower extremity amputees. Arch. Phgs. Med.Rehabil. 60, S&61.Urban, J. (1973) Effect of physical training in healthy maleunilateral A,!K amputees. Stand. J. Rehab. Med. 5, 88-101.Waters, R., Perry, J., Antonelli, D. and Hislop. H. (1976)Energy cost of walking amputees: the influence of level ofamputation. 3. Bone Jt Surg. 58A, 4246.

Winter, D. A., Sidwall, H. G. and Hobson. D. A. (1974)Measurement and reduction of noise in kinematics oflocomotion. J. Biomechanics 7, 157-159.Winter, D. A. (1980) Overall principle of lower limb supportduring stance phase of gait. J. Biomechnnics 13, 923-927.Winter, D. A. (1983) Energy generation and absorption at theankle and knee during fast, natural and slow cadences. ChinOrthop. Rel. Res. 197, 147-154.Winter, D. A. (1984a) Use of computer averaged EMGprofiles in the diagnosis of pathological gait. Archs Phys.Med. Rehabil. 65, 393-400.Winter, D. A. (1984b) Kinematic and kinetic patterns inhuman gait: variability and compensating effects. Hum.Mumt Sci. 3, 51-76.

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