Performance assessment of the Terry Fox jogging prosthesis for above-knee amputees

J. Biomechonm Printed

in Great

Vol. 22, No. 617, pp. 543-558,

1989

Britain

G

0021-9290/89 S3.W + .M) 1989 Pergamon Press plc

PERFORMANCE ASSESSMENT OF THE TERRY FOX JOGGING PROSTHESIS FOR ABOVE-KNEE AMPUTEES DENIS J. DIANGELO*, DAVID A. WINTER?, DHANJOON. GHISTA~:and W. ROY NEWCOMBE* *Department of Mechanical Engineering, McMaster University, Hamilton, Ontarib, Canada L8S 4L7; TDepartment of Kinesiology, University of Waterloo, Waterloo, Ontario, Canada; IDepartment of Medicine, McMaster University, Hamilton, Ontario, Canada L8S 4L7 Abstract-The Terry Fox jogging (TFJ) prosthesis was developed at Chedoke-McMaster Hospital to alleviate the asymmetric jogging pattern experienced by above-knee amputees when attempting to jog with conventional walking prostheses. This prosthesis features a spring-loaded, telescoping shank designed to eliminate any vaulting action and control the trunk motion during stance. The spring is intended to attenuate the impact forces and release its stored energy at push-off to provide momentum transfer to the jogger. This prosthesis was comprehensively assessed in the gait laboratory, by evaluating the kinematics, energy and power flow patterns of an above-knee amputee jogger wearing the TFJ prosthesis. Included in the assessment is the ability of the prosthesis to satisfy a set of relevant design criteria that have been established from non-amputee jogging patterns. An increased swing phase time for the prosthetic limb and the need to have the knee hyperextended throughout the stance phase contributed to an asymmetric jogging style. The telescoping action did lower the amputee’s centre of mass, thereby reducing the vaulting effect. However, the spring only imparted a lifting action to the jogger and the ground reaction forces were double those of a non-amputee jogger. These findings clearly indicate a need to redesign the TFJ prosthesis and are being incorporated in the design of a new physiological jogging prosthesis.

INTRODUCTION

jog, (b) provide a natural, balanced, prosthetic-tonatural leg stride and (c) enable an amputee to participate more actively and naturally in sports involving rapid lower limb motion. Hence, there is a growing market for a prosthesis that can fulfill this need.

Need A rapidly growing interest in the need for sports oriented prosthetics was demonstrated by the publicity generated by the Terry Fox and Steve Fonyo runs akross Canada. Because of bone cancer, the number of young above-knee amputees is growing, and according to a very rough estimate there are about 10,000 in Canada and the United States. The majority of these are young people who wish to continue an active sports life. This is supported by a study by Kegal et al. (1980), which showed that 61% of 134 lower extremity amputees participated in one or more physical recreations. Many of the amputees surveyed were active in sports prior to losing a limb, and expressed a desire to remain active despite their disability. Enoka et al. (1982) indicated that the ability to run at least short distances provides for a more active life-style and improved quality of life for lower extremity amputees. Jogging, which is the basis for many sports, was indicated as the most difficult activity to achieve and that which causes the most discomfort. Nevertheless, many still expressed an interest in being able to run. The study also included a list of suggestions from amputees on how to improve amputee sports and recreational prostheses, and the most popular were to build better prostheses and impact absorbing legs. Thus, there is a need to provide above-knee amputees with a prosthesis that will (a) allow an amputee to

Received in final form June 1988.

Criteria

The development of prosthetic limbs for above-knee amputees requires a detailed understanding of the natural functioning of the lost extremity to enable the definition of design objectives and specifications. Artificial limbs must be capable of restoring as many of the lost functions as possible, be comfortable and safe, without inflicting undue mental or physical effects on the user and have a normal appearance. Furthermore, the prosthesis has to respond to the varying demands of the amputee, as it becomes an integral part of the amputee’s daily life. A substantial amount of research has been conducted on analyzing the walking patterns of aboveknee amputees; this has resulted in the development of conventional above-knee prostheses which facilitate a reasonably natural walking gait. The conventional prosthesis works on the principle that the amputee must keep the line of action of the ground reaction force vector in front of the hinge knee joint when it is hyperextended. To accomplish this he must create a hip extensor moment during stance to keep the knee locked against an extensor stop. However, when above-knee amputees attempt to jog with the conventional single or double-hinge knee prosthesis, they encounter the common problem that they cannot produce a sufficiently rapid extension of the lower limb to the hyperextended locked position

543

544

D. J. DIANGELOet al.

ready for heel strike and weight bearing. Because of this requirement, the jogger must wait for the prosthetic knee to lock, and then vault over the straightened leg and this produces the unorthodox jogging pattern of double support periods on the natural limb. Also, the ground reaction impact forces rise more rapidly and are much higher than during normal running, which is a low velocity heel contact. Thus, premature degeneration of the residual hip joint and a shortened life for the prosthesis results. These drawbacks clearly indicate the need to define a complete set of design criteria for a prosthesis which can enable an above-knee amputee to jog in a more natural fashion. Based on the biomechanical analysis of non-amputee joggers (Winter, 1983), the main design principles to be incorporated in the prosthesis are defined according to stance and swing phase requirements. The stance phase requirements are to: (1) provide some controlled knee flexion to (a) absorb the large ground reaction (impact type) forces at heel contact, as well as (b) lower the centre of gravity of the body, which eliminates the energy expended to vault over the prosthesis, and corrects the abnormal pelvic rotation that would otherwise be present; (2) provide a self-stabilizing motion at the knee joint and (3) generate some vertical but predominantly forward motion. The prosthesis should also be designed to eliminate the need for utilization of the hip muscles to maintain stability during stance by locking and unlocking the knee joint, which is required with conventiona knee units. Swing phase requirements are to (1) regulate the motion of the prosthesis in a manner similar to that effected by the neuromusculature structure surrounding the intact knee-the prosthetic leg must first be accelerated and then undergo controlled deceleration prior to heel strike; (2) allow for free rotation of the lower limbs during non-support periods and (3) provide some amount of knee flexion by the end of the swing, necessary for stability reasons at heel contact.

The Terry Fox jogging prosthesis In an attempt to satisfy these criteria, the Terry Fox jogging (TFJ) prosthesis was developed at Chedoke Hospital, Hamilton, Ontario and is shown in Fig. 1. The new concept here is the incorporation of a large compression spring in the shank which will be compressed during weight acceptance, thus attenuating the peak of the impact force. The energy stored in the spring would then be released at push-off to aid in propelling the runner forward and upward. The original idea was proposed by the late Terry Fox who had

the metal shank of a conventional prosthesis replaced by a pogo stick. The total prosthesis consists of an open-end suction socket, commercial four-bar linkage knee unit with pneumatic swing control, a Griessinger multi-axis foot and a telescoping shank. The shank is referred to as the shin-tube spring unit and consists of two telescoping metal tubes surrounded by and working in parallel with a coil spring. A two-way air damper is also provided here to further reduce the impact force, and prevent the coil spring from collapsing completely. The damper consists of a plastic plug working inside the tubes. The rationale for the telescoping shank is to allow the residual hip joint to be lowered in a manner similar to that displayed by non-amputee joggers who accomplish this by flexing their knee during weight acceptance. Such a mechanism serves to control pelvic rotation and effectively eliminate vaulting over the prosthetic leg, thereby providing a more natural jogging appearance. At toe-off the amputee’s hip flexors act normally to pull the limb upwards and forwards and the swing control device should instinctively decelerate the prosthetic knee motion in the same way the hamstring muscle groups govern the anatomical knee during late swing. Objectives and scope

This paper presents a set of relevant design principles for a jogging prosthesis and provides a comprehensive analysis of the kinematics, energy and power flow patterns of an above-knee amputee jogger wearing the Terry Fox jogging prosthesis. This has enabled a complete assessment of the jogging patterns of an above-knee amputee as well as a quantitative comparison with non-amputee jogging patterns. Although this prosthesis was not found to fulfill the criteria of a jogging prosthesis, the conclusions drawn from this study have been most useful for the design of a new prosthesis that will permit a more physiological jogging pattern. METHOD

A complete assessment of the jogging stride of an above-knee amputee wearing the TFJ prosthesis was carried out at the Gait Laboratory of the Department of Kinesiology at the University of Waterloo. The subject was a 21 yr-old male weighing 72 kg and 184 cm in height who had his right leg amputated above the knee. The subject wore his own running shoes. Reflective markers were placed on the normal leg at the following anatomical landmarks: toe; fifth metatarsal joint; heel; ankle (lateral malleous); lateral femoral epicondyle and the greater trochanter. Markers were placed on the prosthesis at the closest estimated location of the same landmarks, except that markers were placed at the bottom of the spring tube assembly

Fig. 1. The Terry Fox jogging prosthesis consists of an open-end suction$ocket, commercial four-bar knee unit with pneumatic swing control, a Griessinger multi-axis foot and a specially designed shin-tube spring unit.

545

Terry Fox jogging prosthesis (lateral malleous), top of spring tube assembly (midshank) and at the low and upper pin connections of the 4-bar knee mechanism. The exact location of these is shown in Fig. 2. An estimate of the mid-trunk region was made, but was not included in the running analysis. The subject was instructed to jog at a slow and comfortable speed, while a tracking cart containing a TV and cinecamera (50 frame s-l) ran on a track parallel to the walkway at a distance of 4 m. Background markers were placed on the wall beside the walkway to provide a reference system for transferring the body-coordinates to an absolute coordinate system. Simultaneously, force plate data were recorded from an AMTI force platform along with a 50 Hz

All

dimensions o indicates

in

547

synchronizing pulse generated by the cinecamera, and the vertical and horizontal components were combined with the centre of pressure to give the force vector under the foot. Coordinates of the body and background markers were extracted from the film using a Numonics digitizer interfaced with a Sorcerer microprocessor. The raw coordinates were scaled, and then corrected for parallax error between the planes of the jogger’s motion and background wall. The absolute coordinates of the limb markers were calculated and processed on an IBM 3431 computer for the kinematic analysis. Previous analyses conducted by Winter (1983) revealed that the noise in the data, mainly due to the digitizing process, had a root mean square of

(cm)

&16.9 R:17.7

L111.7 kl0.2

marker

Fig. 2.Marker locations for the amputated side, R, and non-amputated side, L, employed with the gait laboratory study.

548

D. J. DIANGELOer al.

2 mm or less for all markers. In the process, the coordinates were digitally filtered using a fourthorder, zero-lag, low-pass Butterworth filter cutting off at 6 Hz (Winter et al., 1974). Anthropometric constants for the normal leg were obtained using Dempsters tables (Winter, 1979) based on the subject’s height and weight. The necessary anthropometric constants for the TFJ prosthesis were determined using the procedure outlined by Martin (1982) and are provided in Table 1. ANALYSIS Link segment model-forces and moments

l

The biomechanics analysis uses a standard link segment program, which enables an equilibrium balance to be performed on each segment for the calculation of the vertical and horizontal forces plus the net joint moments, at the ankle, knee and hip, for one complete stride commencing with heel contact on the force plate (Quanbury et al., 1975; Winter and Robertson, 1978). The foot segment was modelled by assuming a rigid link between the ankle and the fifth metatarsal marker. The toe marker was used to record the toe trajectory during swing and the heel to fifth metatarsal was used to define the angle of the foot segment for purposes of calculating foot angular velocities and ankle joint angles. Note that the muscle forces represent the net activity developed at a joint and are expressed as net muscle moments. If muscle co-contractions take place at a common joint, the analysis yields only the net effect of both agonist and antagonist muscles. Therefore, both flexors and extensors can be active simul-

taneously at a common joint, but the resultant net muscle moment will be the summation of the two and have the same polarity as the larger of the two. Furthermore, the net muscle moments at the hip, knee and ankle joints will be addressed as either a flexor or extensor moment or by the muscle group responsible for creating a flexor or extensor moment about the respective joint. Power calculations

The instantaneous mechanical energy of any segment, s, having one of its ends at joint, j, at time, ti, is given as the sum of the potential and translational and rotational kinetic energies (Robertson and Winter, 1980); E(s, ti) = m(s)g Y(s, 4)++m(S)[ VS,ti)12

+3rz(S)C~z(S~ ti)12

(1)

where m(s) is the segment mass, I,(s) is its moment of inertia about the medial-lateral z-axis through its center of gravity, Y(s, ti) is the height of the segment mass at time, ti, V(s, tr) and w(s, ti) are its linear and angular velocities at time, ti. For the (X, Y, 2) coordinate system, X refers to the anterior-posterior axis, Y the vertical axis and 2 the medial-lateral axis. The segment’s rate of change of energy with respect to time is calculated by a finite-difference technique: fi(s, ti)=[E(s,

t,+,)--E(s,

ti-l)lPAT

where ti, ti+l, ti-l represent time increments separated by an interval AT=0.02 s. A positive result indicates-a rate of inflow of energy to the segment whereas a negative sign indicates a rate of outflow. The power flow across any joint is termed joint

Table 1. Segment parameters for Terry Fox jogging prosthesis Segment parameters Mass (kg) % Total mass Segment length (m) Segment length from centre of gravity to proximal end (m) x 10m2 Proportion of length to centre of gravity Mass moment of inertia about centre of gravity (kgm’)

Griessinger foot

Spring unit

Lower knee unit

Upper knee unit

0.859

0.112

1.108

0.046

5.656

25.156

1.19

1.07

1.54

0.06

7.84

35.70

0.160

0.218

0.156

0.054

0.390

1.63

8.20

3.70

0.27

0.102

0.372

0.23

0.500

0.353

0.626

0.004421

0.005017

0.009101

0.0001

0.1087

1.8395

*HAT refers to head, arms and trunk.

(2)

Stump and socket

13.76

@AT*

0.540

33.80

Terry Fox jogging prosthesis force power (JFP)

and is calculated by:

JFf’j(j,s)=F(j,s)* v(j)=F,(j,s)V,(j)+F,(j, s)v,(j) (3) where JFP,(j, s) is the power delivered to or taken from a segment, s, at joint, j, due to work done by the joint reactions, F(j, s) moving with the joint’s linear

velocity, V(j) (Winter et al., 1976). The joint force power indicates the rate of transfer of energy across a common joint, j. Both segments experience the same joint velocity, but the joint reaction forces will be equal in magnitude and opposite in direction. Muscle power (MP) is the mechanical power delivered to or taken from segment, s, at its joint, j, due to work done on or by muscles and is given by: MPj(j,

(4)

s)=M(j, Sk44

where M(j, s) is the joint moment acting on segment, s, at joint, j, and w(s) is the absolute angular velocity of the segment, s. If M (j, s) and w(s) are of the same polarity, then MPj(j, s) is positive; otherwise, it is negative. A positive MPj(j, s) represents the rate of mechanical work done by the muscle on segment, s, and involves the muscle acting concentrically as it generates or transfers energy to the segment. A negative rate implies that the segment does work on the muscle which absorbs energy as it lengthens and contracts eccentrically, or acts to transfer energy from the segment. Two segments meeting at a common joint do not necessarily have the same angular velocity, and therefore are capable of generating or absorbing power at joint, j. The total power generated or absorbed by the muscles will be the summation of the individual muscle powers affecting both segments and is given by:

where M, and M, are the respective muscle moments on the adjacent segments at a common joint, j. Note that M, and M, are equal in magnitude but opposite in direction. By convention, the moment at the distal segment defines the net muscle moment, M,, at the joint. This is exemplified in Fig. 3 where M, of the

549

distal segment is positive (counterclockwise) resulting in a positive net muscle moment, Mj. The net, or relative, angular velocity of the joint is oj=02 -w,; therefore, the total muscle power absorbed or generated at joint, j becomes: MPj, gen/abs= Mjmj.

P-5)

If both segments rotate in the same direction, energy can also be transferred through the muscles from one segment to the other, always flowing to the segment with a positive MPJ(j, s), i.e. that segment to which energy is being generated. The amount of power transferred through the muscles will be either equal to that segment’s MPj(j, s) which has a positive value when MPj, gsn/abs is negative, or equal to that of the segment with a negative MPj(j, s) when MPj,gsn,abs is positive. If the angular velocities of the two segments are the same, then only a transfer of energy from one segment to the other will occur, as would hold true for an isometric contraction. A list of all possible muscle power functions (i.e. generation, absorption and transfers) that can occur between a proximal and distal segment connected by an active muscle experiencing a positive joint muscle moment, is given in Robertson and Winter (1980). Note that the muscles referred to are not the actual anatomical muscles, but are equivalent joint muscle groups which can either flex or extend the joint. Moments and angular velocities used in equations (4H6) are taken as positive when acting in a counterclockwise direction. The total instantaneous power for a segment at time, ti is given by

TP(S,ti)=C [JFPj(j, S),G)+ MPj(j, S),ti)I sj

(7)

where CSj implies the summation, of all the segment joints, and when the segment’s total power is positive it is gaining energy. Conversely, a negative total instantaneous power implies the rate of loss in mechanical energy. To analyze the jogging stride, the difference in values of the total rate of change of energy, E(s, ti) was

PROXIMAL

/ 1

Aa DISTAL Fig. 3. Representation of the net joint moment, Mj, between two segments which is equivalent to the net moment at the proximal end of the distal segment.

D.J. DIANGELO etal.

550

calculated by applying equation (2) to each prosthesis segment, and the total power flow, TP(s, ti) was calculated by applying equation (7). This was done at every time interval, ti in the stride. Some minor discrepancies in the power flow values resulted, which were generally due to small movements of the centre of rotation of the joint under consideration. The following assumptions, defined by Winter (1979) were utilized in the power flow calculations: (1) each link segment has a lumped mass located at its centre of mass; (2) the location of the centre of mass remains fixed during movement; (3) the mass moment of inertia of each link segment about its centre of mass remains constant; (4) pure rotation occurs about the ideal hinge or pin joint; (5) no dissipative forces such as friction exist and (6) joint reaction forces act through the centre of rotation of the joint. The results of this analysis are given in the next section.

RESULTS

The subject’s jogging characteristics are summarized in Table 2. Two trials were performed on each side. However, because the two results were very close, only the result of one trial from each limb is discussed, namely trial WP89B for the amputated leg and jogging prosthesis, and trial WP89C for the non-amputated leg. A stick-diagram representing the amputated side of the body is shown in Fig. 4 for the complete jogging stride. The peak ground reaction loads for both trials are given in Table 3, and include the average peak forces recorded for normal joggers. The peak forces produced on both sides of the amputee jogger are greater than those observed for normal joggers, and the peak reaction loads on the prosthetic leg are almost double those associated with normal joggers. The operation of the spring tube assembly will

Table 2. Characteristics of jogging trials Characteristics of jogging trials

Event*

Amputated leg WP89A WP89B

Non-amputated leg WP89C WP89D

Stride rate (1 s-l)

0.86

0.84

0.84

0.88

Stride length (m) Stride velocity (m s-t)

2.546

2.613

2.650

2.168

2.96

3.11

3.15

3.14

0

0 0

0

26

26 33

34

56

57

81

82

Ipsilateral heel contact (% of stride) Ipsilateral toe-off (% of stride) Contralateral heel contact (% of stride)

HCA HCNA TOA TONA HCNA HCA

42

44

Contralateral toe-off (% of stride)

TONA TOA

72

74

*The event represents the time of occurrence of heel contact (HC) and toe-off (TO) for both the amputated (A) and non-amputated (NA) legs during the jogging stride.

Fig. 4. Stick figure representation of jogging stride for the amputated side.

551

Terry Fox jogging prosthesis Table 3. Ground reaction forces Peak resultant forces Horizontal Vertical (NJ (N)

Trials

4.3 4.2 3.0 2.8

119 161 233 268

3070 2993 2091 1981

WP89A* WP89B* WP89Ct WP89Dt

Peak vertical forces as a function of body weight

$ Average from normal joggers

2.4

*Amputated side. t Non-amputated side. $ Average values recorded for nine normal joggers with similar jogging speeds as those of the amputee jogger.

RELRTIVE ANGLES-WP89B vs NORMALS

.* . -

-.y.*-

RELflTIVE ANGLES-WP89C vs NORMALS

4

.,......

.........

I

Qd,h”,

0

0

0

0

N

% OF STRIDE Fig. 5. Relative angular displacement patterns at the hip, knee and ankle joints for the amputated leg, WP89B, vs those for a selected group of normal joggers. The artificial knee joint remains hyperextended throughout stance and undergoes excessive rotation up to 120” during swing.

be discussed with regard to its ability to store and return energy back to the jogger. Plots of the relative angle patterns of the ankle, knee and hip are shown in Figs 5 and 6. The artificial knee joint remained hyperextended throughout the stance phase and exceeded the normal level of knee flexion during swing. The intact knee joint also displayed excessive knee flexion during the swing phase. The ankle, knee and hip moment of force patterns are plotted in Figs 7 and 8, with extensor moments shown

% ;F STR;BE

0

m

I *

Fig. 6. Relative angular displacement patterns at the hip, knee and ankle joints for the non-amputated side, WP89C, vs those for a selected group of normaljoggers. Increased ankle planar flexion occurred during late stance and excessive hip and knee flexion was present during the swing phase.

as positive and flexor moments negative. The summation of the three moments, called the support moment (Winter, 1980) is also included in the moment curves. The net support moment prevents collapse of the lower limbs during stance, and will occur if flexion of the hip, knee and ankle joints cannot be resisted. Figure 10 shows that support for the intact limbs is maintained by a net extensor pattern at all three joints. However, the net support moment of the prosthetic leg, shown in Fig. 9, is minimal. This is due to the

552

D. J. DIANGELOet at.

MOMENT OFFORCEWPfl!B vsNORMALS

MOMENT OF FORCEWPEl!C vsNORMALS SLOW JOG 101

’

wI

-MMnStM) ..........c&l&“.

K5 m- ’

-

--IpBsc fNKLE

I

I._

0

0

0

N

t

0

u)

0

m

% OFSTRIDE

0 0

I

I._

0

.I

Fig. 7. Moment of force patterns calculated at the hip, knee and ankle joints for the amputated leg vs those for a selected group of normal joggers. Support moment is the summation of all three joint moments, with positive polarity for extensor, negative for flexor.

flexor moment developed by the extensor stop about the artificial knee joint, since the knee has to be hyperextended and locked to provide positive support throughout stance. The muscle power absorbed and generated at each joint is shown in Figs 9 and 10, and is equal to the angular velocity of the joint times the moment about it. Average results for the mechanical power developed by normal joggers, determined from previous studies by Winter (1983), are included in all graphs, along with the major phases of power generation and absorption. The passive artificial ankle displayed two minor power phases, labelled AlA and A2A, whereas the intact ankle had two distinct phases of power generation and absorption. Figure 10 shows seven phases of power occurring at the intact knee joint, which are similar to the knee power patterns for a non-amputee jogger, with the exception of phases K2A and K2B. The hip joint of the amputee’s normal limb displayed three well defined phases of power (labelled Hl, H2 and H3). Figure 11 is a graph of the power flow delivered to the spring unit, and relates the exchange of muscle power and joint force power between the passive artificial foot and the residual limb. The power flow pattern is also affected by the operation of the mechanical spring; this is discussed in the next section. It

0

0

0

N

*

In

0

m

’ ’

% OFSTRIDE Fig. 8. Moment of force patterns calculated at the hip, knee and ankle joints for the non-amputated leg vs those for a selected group of normal joggers. Support moment is the summation of all three joint moments, with positive polarity for extensor, negative for flexor.

should be noted that for the joint force power flow analysis, the artificial knee joint was idealized as a single hinge by considering the upper pin of the fourbar mechanism to be the centre of rotation.

DISCUSSION Reaction forces

The ground reaction forces are summarized in Table 3, showing peak vertical loads of 3070 N for the prosthetic leg and 2091 N for the normal leg. The peak loads produced on the prosthetic leg are almost twice as high as those observed on normal joggers, and occur because the artificial knee joint is completely extended during heel strike and throughout the stance phase. Thus, the ground reaction force is transmitted straight up the rigid prosthesis and through to the hip joint, whereas for the normal jogger, the knee is flexed 15-20” at heel contact and flexes to 45” during weight acceptance to allow the knee muscles to absorb most of the shock. The peak joint reaction force occurring at the hip joint of the TFJ prosthesis (assuming a rigid thigh segment) is given in Table 4, and reached 2878 N for a ground reaction force of 3070 N, the difference being the inertia forces of all components distal to the hip joint.

Terry Fox jogging prosthesis

POWER GEWABS -WPOSB vs WtifUS

1900.0~

553 POWER fLOW TO SPRING UNIT

,;

-UN=31 .........5td.h”. -

-ml98

STRIDE

SLOW JOG

% OF STRIDE

-

Fig. 9. Composite of muscle power patterns at all three joints for the amputated leg vs that for a selected group of normal joggers. Two minor phases of power absorption and generation, AlA and A2A, occurred at the artificial ankle. Power levels at the conventional knee and residual hip are minimal.

Table 4. Resultant joint forces

Joint

Resultant reaction force (N) WP89B WP89A

TIME

(SEC)

Fig. 10. Composite of muscle power patterns at all three joints for the non-amputated leg vs that for a selected group of normal joggers. In all phases of power generation and absorption, the non-amputated leg exceeds that for normal joggers.

Shin-tube spring unit

When calculating the instantaneous power of the spring unit, I$, ti) in equation (2), the changing link length is taken into account, whereas the total power calculation, TP(s, ti) in equation (7), did not account for this varying length. The difference between these power calculations provides a good estimate of the power flow into, or out from, the shin-tube unit AP, = @s, ti) - TP(s, ti).

Lower knee pinned connection

3012.7

2933.7

Upper knee pinned connection

3011.5

2932.5

Hip joint

2878.0

2804.7

To accommodate

these high loads, the prosthesis would need to have heavier and stronger components. Existing prostheses are designed for walking, during which the ground reaction forces are of the order of body weight. During jogging, these forces increase to over four times body weight, and the life of a standard commercial prosthesis will be severely shortened. These high forces will also cause premature degeneration of the hip joint and problems at the stumpsocket interface.

(8)

A negative value for AP, indicates an inflow of power into, whereas a positive value indicates an outflow of power from the segment. An indication of the energy stored/recovered by the shin-tube unit can be calculated by summing up all the respective values for AP,, and multiplying by the time interval between frames, At =i +AP,;At. AErot,a~,s

(91

Note that the values for f AP,, are considered for the compression and extension states of the spring during the stance phase only, and, n represents the number of frames for any state. The spring was being compressed for 42% of the stance phase, and was extending during the remaining 58%. The shin-tube spring unit absorbed 55.76 J and returned 28.12 J during trial WP89B. Thus, only about 50% of the energy absorbed was returned to the

D. J. DIANGELOet al.

554

jogger, the remainder being lost to friction. The air damper was only active during the latter part of the stance phase when the spring was extending. This accounted for most of the losses, and this item is not necessary because the extension of the spring can be self-controlled by the weight of the runner and the ability of the spring to push against him. When the spring unit was designed at Chedoke Hospital it was conceived to absorb and dampen the impact force. However, the exact way in which this would be accomplished was not clearly understood. As a result, 94% of the peak reaction force was transmitted to the residual hip joint and indicates the inability of the spring unit to provide adequate damping. The results from the biomechanical study, Table 5, indicate that the spring undergoes complete collapse, but the unit is not attenuating as much of the impact force as is theoretically possible. The spring unit provides a resistive force that acts against the impact load, thereby reducing the rate of change of the vertical momentum of the jogger. The optimum design depends on the spring characteristics, such as the working length and spring rate, and are parameters to be selected by the designer. Summary of support moment

Support of the body is maintained during stance by a net extensor pattern at the hip, knee and ankle joints, and is defined as: M,=M,-MA-M,

(10)

where subscripts K, A and H represent knee, ankle and hip, respectively. The support moment for a normal jogger is primarily extensor during stance, flexor during early swing, and extensor during late swing. The hip has an extensor peak at 20%, the knee at 40%, and the ankle at 60% of stance. In addition, a significant decrease in the variability of the support moment pattern of normal jogging, compared with the same pattern for normal walking, has been reported by Winter (1983). Amputated side. Here, the net support moment for the prosthetic leg was similar neither in magnitude nor polarity to that of the non-amputated jogger; Fig. 7 shows a peak of 0.8 N-m kg- ’ at 10% stride for the amputee and 5.4 N-m kg- ’ at 16% stride for the nonamputee jogger. The ground reaction force vector passed in front of the hyperextended artificial knee joint, thereby developing an external extensor moment about the knee. This is calculated as an internal flexor moment in Fig. 7 and was generated by the extensor stop of the conventional knee unit. Thus, knee collapse in this atypical situation is prevented by a mechanically generated internal flexor moment which reduces the net total limb extensor pattern seen in M,. However, extensor moments at the hip and ankle will act to hold the knee in hyperextension and therefore prevent knee collapse. If the reaction force were to pass behind the knee centre, then the prosthesis would require sufficient resistance in the knee hinge to overcome the collapsing moment. This would be similar to knee flexion on the non-amputated side,

Table 5. Vertical reaction forces and segment length differences of the spring unit

Trial

Event HCA

WP89A

TOA HCA

WP89B

TOA

Frame number

Time (s)

Segment length difference (cm)

8 9 10 11 12 13 14 15 16 17 18 19

0.00

0.0

0.02 0.04 0.06 0.08 0.10 0.12 0.14 0.16 0.18 0.20 0.22

0.5 2.5 4.0 4.4 4.5 4.3 3.8 2.6 1.3 0.9 0.0

4 5 6 7 8 9 10 11 12 13 14 15

0.00

0.0

0.02 0.04 0.06 0.08 0.10 0.12 0.14 0.16 0.18 0.20 0.22

0.5 2.0 3.7 3.9 4.3 4.2 3.1 1.6 0.7 0.4 0.0

Vertical reaction force (N) 27 615 1328 2175 3070 2531 1436 923 701 395 122 20 38 703 1344 2290 2993 2249 1273 913 680 360 92 11

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Terry Fox jogging prosthesis and the need for a large extensor pattern at all three joints to control the amount of collapse. Information such as this is necessary for prosthesis design. Non-amputated side. The net support moment of the intact leg peaked sooner, and reached a higher maximum compared with that of the normal jogger. Figure 8 reveals a peak of 6.9 N-m kg- ’ at 13% stride for the amputee. The increase in the extensor moment was largely due to the higher ground reaction forces resulting from his greater vertical displacement during his flight phase and higher velocity at heel contact of his non-amputated limb. This results in greater energy absorption by the knee extensors as shown in Fig. 10. During the remainder of the stride it was similar to that of a normal jogger. To maintain symmetry between the orientation of the amputee’s amputated and non-amputated sides during the stance phase, it is desirable to have the support moment data as representative as possible of the normal situation. Muscle power analysis

The muscle power at each joint was calculated using equation (6), and it is equal to the angular velocity at the joint times the moment about it. The work done by the muscles at each joint is equal to the area under the corresponding portion of the power curves and will be discussed for both the legs. Amputated side. Ankle. Two power events were present in the muscle

power curves for the ankle, as shown in Fig. 9 and have been labelled Al A and A2A. At heel contact, the soft foam insert of the Greissinger foot compressed immediately at the heel, causing the ankle to take on a dorsiflexion orientation. A plantarflexor moment was developed at the ankle, which allowed for approximately 12.5 J (AlA) to be absorbed from the foot. The ankle then reversed direction and allowed the passive ankle to return a small amount of power, 3.6 J (A2A). This represents a 29% recovery of stored energy and is similar to that reported for the Greissinger foot in walking (Winter and Sienko, 1988). No energy generation or absorption was present at the ankle throughout swing because the artificial foot is completely passive. Knee. The quantity of power absorption/return occurring at the artificial knee joint during the stance phase was negligible compared to that observed by Winter (1983) for a normal jogger’s knee. The prosthetic knee remained in a hyperextended position throughout stance to provide stable support to the subject. The swing control mechanism in the four-bar linkage knee was active in decelerating the lower limb during the latter part of swing; however, only 7.3 J were absorbed which is less than that absorbed by the knee flexors of a normal jogger, 24 J (Winter, 1983). This difference is due to the fact that a normal lower leg and foot have a somewhat greater mass and mass

moment of inertia than the prosthetic replacement; 4.4 kg and 0.75 kg m’, and 2.8 kg and 0.26 kg m2 for the non-amputated leg and TFJ prosthesis respectively. Hence. the hamstring muscles must absorb considerably more energy to decelerate the lower limb prior to heel contact. Hip. The power values at the hip are quite low, having peak positive power levels of up to 70 W during stance. For normal joggers, Winter (1983) observed the power patterns at the hip to be low, with the only consistent burst to be H2 (Fig. 10) which was due to concentric contraction of the hip flexors early in swing. The WP89B trial for the amputated side showed this burst to be absent. Figure9 indicates that hip extensor activity was present up to mid-stance, providing support and stability to the jogger. A similar period of hip muscle activity has been revealed in the walking patterns of above-knee amputees. The hip trajectory of the amputated limb differs from the normal pattern throughout most of the first half of stance, as shown in Fig. 5. The amputee restricts the advancement of the prosthesis by limiting hip flexion, hence minimizing the drop in the hip joint seen at heel contact. By limiting the hip trajectory, the amputee avoids an exaggerated tilting of the pelvis which contributes to the asymmetry detected during the stance phase. Non-amputated

side.

Ankle. The muscle power patterns of the normal ankle for trial WP89C are shown in Fig. 10, and closely resemble those of a normal jogger except that the peak power levels are much larger. During early stance, a peak negative power of 760 W was reached along with 53.3 J (Al) being absorbed. Figure 6 shows that by mid-stance the ankle has reversed its direction, and begun to rapidly plantarflex while generating an ankle plantarilexor moment. This resulted from the plantarflexors reaching a peak power of 1320 W and generating 90 J (A2), providing the majority of push-off energy to the jogger.

Knee. A similar analysis was carried out for the intact knee, as shown in Fig. 10, where seven separate power events labelled Kl-K5 occurred. Winter reported that a normal jogger has five distinct power events. The additional two, observed here, occurred during the latter phase of stance. A significant amount of power was absorbed by the knee extensors (K2A), followed by a period of lesser power generation (K2B). Aside from these two occurrences, the knee pattern remained quite similar to that of a normal jogger. The initial knee flexion of 20” at heel strike enabled the knee musculature to help absorb the shock load delivered at heel contact; here 45 J (Kl) were absorbed through eccentric contraction of the knee extensors. A large knee extensor moment was developed, as shown in Fig. 8, and the MPK,gcn,absincreased to a negative peak of 908 W. Shortening of the quadriceps prior to mid-stance resulted in 33 J (K2) being gener-

556

D. J. DIANGELOet al.

ated, and a positive peak power of 735 W being attained. This generation burst added potential energy to the body which was quickly converted to forward potential energy at toe-off. The ratio of Kl and K2 demonstrates that the amputee’s knee muscles absorbed only 1.4 times as much energy as they generate compared to 3.6 times as much for a normal jogger. The large increase in energy generation by the intact kne!: muscles was necessary to compensate for the passive artificial leg. As the knee continued to extend, a flexor moment was developed about the knee to control the forward and upward advancement of the body, resulting in the absorption of 12.7 J (K2A), and attaining a peak negative power of 368 W. The hamstring muscle group remained active by contracting concentrically after toe-off, as the knee reversed direction and began flexing, generating up to 4 J (K2B). Knee flexion continued through early swing, reaching a maximum of 120”, but the knee extensors lengthened as they decelerated the posterior swinging shank and foot. This action exceeded the peak knee flexion of 80” observed with non-amputee joggers, and resulted in 15.6 J (K3) being absorbed by the quadriceps. The hamstrings contracted eccentrically to decelerate the forward swinging shank and foot during the reach phase of swing, and absorbed a substantial amount of energy, 45 J (K4). Just prior to heel contact, the same hamstrings reversed the direction of the leg and foot, and although only 2 J (K5) were generated, they were sufficient to reduce the forward velocity of the foot to near zero and allow the knee to obtain a flexed position by heel contact, which was desired for weight bearing purposes.

gers for both stance and swing phases, with two exceptions. Firstly, the knee muscles generate as much energy as they absorb; this accounts for approximately one-third of total generated energy during weight bearing. Secondly, the intact leg produced almost 89% of the total energy generated by both legs throughout the entire jogging stride. This is mostly due to the prosthesis being completely inactive and the inability of the spring unit to contribute to the forward propulsion of the jogger. The power flow pattern of the spring unit is discussed for the complete jogging stride, commencing with heel contact. Stance phase. The knee unit provides a means of transmitting power from the body and thigh to the foot throughout the stance phase. All of the joint force power entering the knee unit through its proximal end, passes through the bottom of the knee unit into the spring unit. During the first quarter of stance, the coil spring of the spring unit collapsed rapidly due to the high ground reaction loads. Compression of the coil spring coincided with the transfer of energy from the thigh to the shank, JFP,,(SU, SU), having a peak positive power of 1300 W, and the development of a large negative muscle power at the prosthetic ankle, MP,(A, SU) reaching a negative peak of 600 W. Both of these patterns are shown in Fig. 11, and are similar

POWER GEN/ABS -WP89C vsNORMALS

Hip. A plot of the hip muscle power is shown in Fig. 10; three distinct events of power exchange were recorded, and have been labelled as Hl, H2, and H3. The first power burst occurred during stance where the hip extensors shortened, generated 22.8 J (Hl), and reached a peak positive power of 185 W. This energy generation by the hip extensors assisted in the forward propulsion of the trunk by pushing it from the rear. Throughout the first half of swing, the hip flexors shortened as they contracted concentrically to achieve a pull-off of the total limb, thereby generating 24.1 J (H2). This accounted for approximately 10% of the overall power required to accelerate the leg forward, as well as to provide the necessary clearance between the swinging foot and the ground. The third and final power phase occurred during late swing, when the hip extensors generated 18.5 J (H3) in an attempt to return the hip to a less flexed position prior to heel contact. The H3 burst, like K5, did not assist in the forward propulsion of the body, but was a necessary overhead to achieve a safe landing. All three power events are much more prominant than those observed for non-amputee joggers, as illustrated in Fig. 10.

Power flow analysis

The power flow patterns of the non-amputated limb were similar to those displayed by non-amputee jog-

Fig. 11. Power flow patterns of the shin-tube spring unit. Total power, TP, is the summation of the joint force power, JFP, and muscle power, MP, entering or leaving the proximal, SPR HT, and distal, ANKLE, ends of the segment.

Terry Fox jogging prosthesis to those observed by Winter (1983) for normal joggers during mid-stance of the jogging stride. Since both ends of the spring unit are rigidly attached to the adjacent segments, the corresponding net moments at either end of the spring unit are equal in magnitude and opposite in polarity. This is shown by the mirror imaging of the ANKLE MP and SPR UT MP patterns in Fig. 11. As a result the total power of the segment represents the relative difference in the JFP patterns at either end. Any difference is directly related to the operation of the large coil spring. Figure 11 indicates the stored energy in the coil spring was returned by toe-off, and consisted of joint force power transferred from the spring unit to the knee unit. A significant difference was recorded in the ankle muscle power of the artificial foot compared with a normal ankle. Winter reported that the average peak of muscle power generation at the ankle of a nonamputee jogger was 800 W. Here the muscle power generated at the artificial ankle was 260 W, and is largely due to the joint force power created during extension of the coil spring in the spring unit. This indicates that the spring unit was achieving only 25% of the desired limit. Prior to toe-off of the amputated leg (TOA), a high transfer of joint force power with a peak limit of 600 W, originating from spring extension, is passed on to the thigh and socket, raising its total power. The joint force power delivered to the thigh through the knee began to drop rapidly, while at the same time the joint force power to the thigh from the trunk increased quickly, causing the total thigh power to reach a positive peak. This occurs at approximately the time of ipsilateral toe-off and aids in accelerating the lower limbs during early swing. Swing phase. Throughout late stance and early swing of the amputated leg, the joint force power transmitted to the trunk from the swinging nonamputated leg (TOA to HCNA) passes via the pelvis and enters into the stump and socket (TOA to HCNA). This also assists the jogger in accelerating the residual hip joint and displays the body’s ability to conserve energy. Hip flexor activity has been considered to be the major contributor in accelerating the thigh and lower limbs. However, it has been shown that the principal supporter is joint force power from the trunk, and not hip flexor muscle power. The joint force power of the intact leg originates during its stance, where the ankle plantarflexors remained active up to the end of stance. Extended ankle plantarflexion at late stance was also seen in the walking patterns of above-knee amputees by Cappozzo et al. (1976), and was considered to be a compensating mechanism necessary to overcome the fact that the prosthetic foot cannot dorsiflex, and that the socket may move slightly downward along the stump during swing. The lower limbs were decelerated for preparation of heel contact during the second half of swing. By rotating the thigh and socket backwards while the knee unit rotates forward, the amputee develops iner-

557

tial forces on the knee unit which assist the prosthetic knee joint in reaching complete extension sooner. This was verified, as the angular velocity of the knee unit decreased drastically to almost zero during the same period of time, denoting complete extension of the prosthetic knee joint. This further contributes to the asymmetry observed in the jogging stride, and emphasizes the lack of adequate damping by the swing phase control mechanism in the knee unit. Its function was to provide controlled deceleration of the lower segments, and eliminate any jerky motion that occurs when an artificial knee joint comes to an abrupt stop. Unfortunately, a conventional knee unit cannot compensate for the increased knee angular velocity associated with jogging, as it is designed to operate over a range of walking speeds only. Hence, the knee unit must be redesigned or modified to satisfy the design criteria of a jogging prosthesis. In addition, by having the knee completely extended prior to heel contact, the amputee is unable to experience the energy absorption and desired hip motion analogous with having the knee flexed 15-20” at the onset of weight bearing. Transfer of power out of each of the swinging segments occurred during late swing and was confirmed by the high negative rate of transfer out of their proximal ends along with an overall negative total power. Similar patterns have been observed during late swing by Chapman and Caldwell (1982). Just prior to heel contact of the artificial leg, a slight increase in hip muscle power took place. This activity limited the forward motion of the swinging foot and positioned the artificial leg in an almost vertical orientation. The greater the reduction in hip flexion prior to heel contact, the less the hip extensor moment required for stability during the initial part of stance.

CONCLUSIONS

The jogging performance of an above-knee amputee wearing the Terry Fox jogging prosthesis was assessed, and compared to non-amputee jogging patterns. This biomechanical study revealed that the TFJ prosthesis was unable to satisfy all of the requisite design criteria for a jogging prosthesis. The amputee was able to adopt a style of gait completely different from what he was previously accustomed to, consisting of alternating periods of single support and non-support for both his amputated and non-amputated legs. However, the nonamputated leg compensated for the passive artificial leg by increasing the amount of energy it generates to 1.8 times the normal, i.e. the normal leg accounted for 90% of the total energy generated by both legs. In addition, the amputee still required to ‘pole-vault’ over the prosthesis into the toe-off position, which is associated with the two-to-one stepping pattern and produces an abnormal motion of the pelvis.

558

D. J. DIANCELO et al.

Since knee flexion was not attainable at heel strike or throughout the stance phase, considerably higher ground reaction forces with peak loads of 4.3 times body weight, almost double that for a normal jogger, were recorded. This was due to the artificial knee unit being completely extended at heel contact, and remaining in a locked position throughout the stance. These high loads are transmitted directly to the stump-socket interface and residual hip joint, and can reduce the prosthesis’ life and cause accelerated wear of the residual hip joint. To accommodate the increased loads, it is recommended that appropriate materials and design strengths be employed to ensure safe operation of the prosthesis without any mechanical failures. The spring unit collapsed completely, and absorbed only 60% of the peak vertical loads. The subsequent telescoping action of the shin-tube spring unit lowered the centre of gravity of the amputee, and reduced the amount of vaulting and pelvic rotation somewhat. However, the spring device did not provide the forward thrust of toe-off that is needed in jogging because the spring released its energy prior to mid-stance, and resulted in only a lifting action being provided to the jogger. It is recommended that a heavier spring be designed, allowing it to be active over the entire range of loads. The action of the air damper in the telescopic shank is not required during spring extension, as it can be self-controlled by the weight of the runner, and the ability of the spring to act against this weight. By overdesigning the coil spring, the air damper would no longer be required during the compression phase of the spring. It is therefore recommended that the air damper be completely removed from the design. Some asymmetry was detected with the jogging motion when wearing the TFJ prosthesis, as the time required for the swing phase of the prosthesis was greater than that for the opposite, normal leg, because the prosthesis had to be hyperextended at the end of the swing phase so that it would be in a locked-knee configuration ready for heel strike and weight bearing. The swing phase control mechanism incorporated in the conventional knee unit was unable to provide suflicient control of the prosthesis during swing activities. Unfortunately, it was designed to operate over a range of walking speeds, and cannot cater to the increased knee angular velocities associated with jogging. These findings clearly demonstrate the inability of the Terry Fox jogging prosthesis, in its presenf design form, to satisfy all of the required design criteria for a jogging prosthesis and can be attributed to the lack of appropriate design analysis.

REFERENCES

Burdett, R. G. (1982) Forces predicted at the ankle during running. Med. Sci. Sports Exercise 14, 308-316. Cappozzo, A., Figura, F., Leo, T. and Marchetti, M. (1976) Biomechanical evaluation of above-knee prostheses. Biomechanics V-A (Edited by Komi, P.V.), pp. 366-372. University Park Press, Baltimore, MD. Cavanagh, P. R. and LaFortune, M. A. (1980) Ground reaction forces in distance running. J. Biomechanics 13, 397406. Chapman, A. and Caldwell, E. (1982) Factors determining changes in lower limb energy during swing in treadmill running. J. Biomechanics 15, 69-77. Chapman, A. and Caldwell, E. (1983) Kinematic limitations of maximal sprinting speed. J. Biomechanics 16, 79-83. Enoka, R. M., Miller, D. I. and Burgess, E. M. (1982) Belowknee amputee running gait. Am. J. phys. Med. 61, 6684.

Hughes, J. and Jacobs, N. (1979)Normal human locomotion. Prosth. orthop lnt. 3, 4-12. Ito, A., Komi, P. V., SjBdin, B., Bosco, C. and Karlsson, J. (1983) Mechanical efficiency of positive work in running at different speeds. Med. Sci. Sports Exercise 15, 299-308. Kegal, B., Webster, J. C. and Burgess, E. M. (1980) Recreation activities of lower extremity amputees: a survey. Archs phys. Med. Rehabil. 61, 258-264. Martin, G. H. (1982) Kinematics and Dynamics ofMachines, 2nd edn. McGraw-Hill, New York. Quanbury, A. O., Winter, D. A. and Reimer, G. D. (1975) Instantaneous power and power flow in body segments during walking. J. Hum. Mumt Stud. 1, 59-67. Robertson, D. G. E. and Winter, D. A. (1980) Mechanical energy generation, absorption and transfer amongst segments during walking. J. Biomechanics 13, 845-854. Sinning, W. E.&d Forlyth, H. L. (1970) Lower-limb actions while running at different velocities. Med. Sci. Sports 2, 28-34. Vaughan, C. L. (1985) Biomechanics in Running Gait, CRC: Critical Reoiews in Biomedical Engineering. (Edited by Bourne, J. R.), pp. 148. CRC Press, Boca Raton, FL. Winter, D. A., Quanbury, A. 0. and Reimer, G. D. (1976) Instantaneous energy and power flow in normal human gait. Biomechanics V-A (Edited by Komi, P. V.), pp. 334-340. University Park Press, Baltimore, MD. Winter, D. A., Quanbury, A. 0. and Reimer, G. D. (1976) Analysis of instantaneous energy of normal human gait. J. Biomechanics

9, 253-257.

Winter, D. A. and Robertson, D. G. E. (1978) Joint torque and energy patterns in normal gait. Bio[. Cybernetics 29, 137-142. Winter, D. A. (1979) Biomechanics ofHuman Movement. John Wiley, New York. Winter, D. A. (1980) Overall principle of lower limb support during stance phase of gait. J. Biomechanics 13, 923-927. Winter, D. A. (1983) Moments of force and mechanical power in jogging. J. Biomechanics 16, 91-97. Winter, D. A. (1983) Energy generation and absorption at the ankle and knee during fast, natural, and slow cadences. C/in. Orthop. Rel. Res. 175, 147-154. Winter, D. A. and Sienko, S. E. (1988) Biomechanics of below-knee amputee gait. J. Biomechanics 21, 361-367.