admoss_msc_0803_small_file _size

Velocometer: a telemetry-based device to measure intra-push changes in racing

wheelchair velocity

Andrew D. Moss

A thesis submitted in partial fulfilment of the requirements of the Manchester

Metropolitan University for the degree of Master of Science by Research

Department of Exercise and Sport Science

Crewe+Alsager Faculty

Manchester Metropolitan University

August 2003

I certify that all material in this thesis that is not my own work has been identified

and that no material is included for which a degree has previously been conferred

upon me

ii

Abstract

Measurement of the intra-push changes that occur in racing wheelchair velocity is important because it assists in explaining how wheelchair athletes accelerate their wheelchairs. This information has direct application to training and coaching in wheelchair athletics. The purpose of this thesis is to present the design, functional characteristics and utility of a telemetry-based velocometer with the ability to measure intra-push changes in racing wheelchair velocity. Studies one to five describe the functional characteristics of the velocometer. Validity and system linearity: a linear relationship was found when velocity calculated from the velocometer was plotted against three test velocities. The average root mean square deviation (ARMSD) was used to compare velocity calculated from the velocometer with velocity calculated by manual digitising. The ARMSD calculated for each test speed from three trials was 0.06 ± 0.002, 0.27 ± 0.05 and 0.48 ± 0.16 m.s-1 at 1, 5 and 9 m.s-1 respectively. Dynamic response: the ARMSD calculated from the five acceleration and five deceleration trials was 0.29 ± 0.086 and 0.51 ± 0.115 m.s-1 respectively. Reliability: the ARMSD was used to compare the mean trial velocity calculated from velocometer and the speed of the wheelchair rear wheels spun using a DC servomotor. The mean and standard deviation of the differences were 0.079 ± 0.008 m.s-1, for the eight disc-wheel trials and -0.014 ± 0.019 m.s-1, for the eight spoke-wheel trials. Resistance: velocometer resistance calculated as a factor of the mechanical resistance of the wheelchair rear wheel spinning in air was 0.50 and 0.91 N, for the disc and spoke wheel trials respectively. Velocometer resistance calculated as a factor of the total mechanical resistance of the wheelchair/wheelchair-user system was 1.37 and 1.82 N, for the disc and spoke wheel trial respectively. The purpose of the sixth study was to use the velocometer in the analysis of the first six pushes of a sprint start in over-ground racing wheelchair propulsion. One experienced international male wheelchair athlete (age = 28 years; body mass = 60.6 kg; racing classification = T4) performed ten maximal over-ground sprint start trials, over approximately 10 m, in his own racing wheelchair fitted with a Velocometer. Each trial was filmed at 200 Hz using a “Pan and Tilt” system. Eight trials were manually digitised at 100 Hz. The raw co-ordinate data were smoothed using a quintic spline routine. The duration of each push cycle decreased from 0.82 ± 0.02 to 0.45 ± 0.01 s. Within each push the mean duration of the propulsive phase decreased from 0.62 ± 0.02 to 0.21 ± 0.01 s. The mean duration of the recovery phase increased from 0.20 ± 0.01 to 0.24 ± 0.02 s. The athlete contacted the rim progressively closer to top dead centre with each push. Similarly, the athlete released the rim progressively closer to bottom dead centre with each push. The data indicate that peak velocity occurs after release. This is due to the motion of the trunk. The main findings of this study support the observation that racing wheelchair propulsion is a complex form of locomotion and cannot be described accurately by using just the established definitions of a propulsive and a recovery phase. The velocometer provides an effective research tool for the measurement of intra-push changes in velocity, which can be used to further the body of knowledge with regard to racing wheelchair propulsion.

iii

Acknowledgements

My sincere thanks go to my supervisors, Dr Neil Fowler and Dr Vicky Tolfrey. They

have been helpful, supportive, encouraging throughout the duration of this M.Sc.

Neil, you have an amazing ability to explain clearly, and with some obvious

excitement, the most complicated biomechanical concepts. Vicky, your guidance

during my involvement with the British Wheelchair Racing Association (BWRA)

sport science support project, gave me an invaluable grounding in applied work on

which the foundations of this M.Sc thesis are based. I would also like to sincerely

thank Tom McKee for his vast knowledge and expertise in the field of electronics,

hard work and enthusiasm.

I would like to gratefully acknowledge Draft wheelchairs for allowing me the use of

a state of the art racing wheelchair and Edward Grazier for trusting me with his

carbon fibre wheels. As a cyclist I know how valuable these things are.

I am indebted to the individuals who gladly gave up their time for my studies. To

Tanni Grey-Thompson and Chris Hallam, my thanks are for educating me in all

things wheelchair racing. I wish all of you the best in your future racing.

I consider myself fortunate to have good friends. To Mark Johnson, Jason Martin and

Ellen Dawson. I offer my sincere thanks for their friendship, support and advice over

the last seven years.

Above all, I would really like to thank my mum for more things than I can possibly

list here, but mainly her love and kindness.

iv

Publications

The following parts of this thesis have been published or are under review for

publication.

Publication

Moss, A. D., Fowler, N. E., Tolfrey, V. L. (2003). A telemetry-based velocometer to

measure wheelchair velocity. Journal of Biomechanics, 36 (2), 253 – 257.

Under Review

Moss, A. D., Fowler, N. E., Tolfrey, V. L. An explanation of the intra-push velocity

profile of over-ground racing wheelchair propulsion during the first six pushes of the

sprint start.

v

List of contents

Contents Page

Title Page i

Abstract ii

Acknowledgements iii

Publications iv

List of contents v

List of tables ix

List of figures x

Glossary of abbreviations xii

Glossary of terms xiv

1. Chapter 1 16

1.1. Introduction 16

1.1.1. Wheelchair sports and the Paralympic Games 16

1.1.2. British Paralympic success 16

1.1.3. Wheelchair sprinting: Technical background 17

1.1.4. A deterministic model for wheelchair sprinting 18

1.1.5. Summary 21

1.1.6. Aim 21

1.1.7. Objectives 22

1.1.8. Hypothesis 22

1.2. Literature review 23

vi

1.2.1. Inclusion criteria 23

1.2.2. Wheelchair related research 24

1.2.3. Wheelchair racing: development of a sport 25

1.2.4. Ergonomics 26

1.2.4.1. Wheelchair-user interface: seat 28

1.2.4.2. Wheelchair-user interface: push-rim 29

1.2.4.3. Manual wheelchair propulsion daily use vs. sport 31

1.2.5. Assessment of athletic wheelchair performance 31

1.2.5.1. Simulated wheelchair propulsion under realistic 32

Conditions

1.2.5.1.1. Wheelchair ergometers (WERGs) 34

1.2.5.1.2. Motor driven treadmills (MDTs) 36

1.2.5.1.3. Over-ground manual wheelchair propulsion 37

1.2.5.1.4. Protocols 56

1.2.5.1.5. Physiological assessment of the wheelchair 61

Athlete

1.2.5.1.6. Biomechanical assessment of the wheelchair 61

1.2.6. Summary 75

2. Chapter 2 77

2.1. A telemetry-based velocometer to measure wheelchair velocity 77

2.1.1. Design of the device 78

2.1.2. Sampling 81

2.1.3. Mounting 81

2.1.4. Calibration 81

vii

2.2. Study 1: validity and system linearity 83

2.2.1. Introduction 83

2.2.2. Method 84

2.2.3. Results 86

2.2.4. Discussion 88

2.3. Study 2: dynamic response 91


2.3.2. Method 92

2.3.3. Results 94


2.4. Study 3: reliability 96


2.4.2. Method 97

2.4.3. Results 98


2.5. Studies 4 and 5: resistance 102


2.5.2. Method 103

2.5.3. Results 106


3. Chapter 3 110

3.1. Study 6: an explanation of the intra-push velocity profile of 110

over-ground racing wheelchair propulsion during the first six

pushes of the sprint start

viii


3.1.2. Method 111

3.1.2.1. Calibration 114

3.1.2.2. Pilot study 116

3.1.2.3. Data collection 119

3.1.2.4. Data analysis 121

3.1.2.5. Digitising error 123

3.1.3. Results 123

3.1.3.1. Coefficient of variation 131

3.1.3.2. Relative momentum analysis 131


3.1.5. Conclusion 139

4. Chapter 4 141

4.1. General Discussion 141

4.1.1. Limitations 143

4.2. Conclusion 145

4.3. Future Recommendations 145

References 147

Appendices 174

ix

List of tables

Table

Title Page

Table 1 Wheelchair coding for tables 2, 3 and 4 39

Table 2 Studies using a wheelchair ergometer to simulate

manual wheelchair propulsion

40

Table 3 Studies using a motor driven treadmill to

simulate manual wheelchair propulsion

51

Table 4 Studies employing over-ground manual

wheelchair propulsion

54

Table 5 Velocometer resistance calculated from rundown

trials

107

Table 6 Actual and calculated pan and tilt calibration

values

114

Table 7 Mean propulsive cycle data for the first six

pushes of the sprint start calculated from eight

trials

124

Table 8 Mean velocity data for the first six pushes of the

sprint start calculated from eight trials

126

Table 9 Mean acceleration data for the first six pushes of

the sprint start calculated from eight trials

127

x

List of figures

Figure

Title Page

Figure 1 A deterministic model for wheelchair sprinting 20

Figure 2 Optical encoder and transmitter assembly 79

Figure 3 Telemetry system block diagram 80

Figure 4 Calibration equation 82

Figure 5 Experimental set-up for studies 1, 2, 3 and 4

showing treadmill wheelchair mounting system

(TWMS)

85

Figure 6 Velocometer validity and system linearity 87

Figure 7 Wheelchair and velocometer wheel dimensions 90

Figure 8 Velocometer and manually digitised, 2D video

film data collected during (a) one acceleration

trial and (b) one deceleration trial

93

Figure 9 Agreement between the constant velocity of a

wheel spinning in air and mean velocity

calculated from the velocometer data, within a

five percent error band, from (a) Ten disc wheel

trials (b) Ten spoke wheel trials

99

Figure 10 Study 5 experimental set-up showing camera and

calibration pole placement in relation to the line

of progression

105

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Figure 11 Study 6 experimental set-up showing the pan and

tilt camera and calibration pole placement in

relation to the line of progression

113

Figure 12 Calibration procedure. Point denoted by cross is

digitised as follows: 1) Top point at bottom of

view, 2) Top point at top of view, 3) Bottom

point at bottom of view, 4) Bottom point at top of

view

115

Figure 13 Upper extremity calibration frame 118

Figure 14 Wheelchair/wheelchair-user system model used

in the manual digitising of the 3D video film

121

Figure 15 Intra-push wheelchair velocity and trunk,

shoulder and elbow angular displacement during

the first six pushes of the sprint start

129

Figure 16 Intra-push wheelchair velocity and trunk,

shoulder and elbow angular velocity during the

first six pushes of the sprint start

130

Figure 17 The relationship between relative, transfer and

total momentum of the head and trunk during the

first six pushes of the sprint start

132

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Glossary of abbreviations

Abbreviation

Clarification

ISMGF International Stoke Mandeville Games Federation

NWAA National Wheelchair Athletic Association

BPAA British Paraplegics Athletics Association

IOC International Olympic Committee

MDT Motor Driven Treadmill

WERG Wheelchair Ergometer

HAT Head, Arms and Trunk

SCI Spinal Cord Injury

CP Cerebral Palsy

SB Spina Bifida

AB Able Bodied

BSEN British Standard European Standards

ARMSD Average root mean square deviation

TDC Top Dead Centre

BDC Bottom Dead Centre

WAnT Wingate Anaerobic Test

P5 Highest mean power output from any five second period during

(WAnT)

P30 Mean power output measured during 30 second (WAnT)

IOF Index of Fatigue

Fiso Isometric Strength

xiii

HR Heart Rate

VE Ventilation rate

MTT Montreal progressive Tack Test

Vc Critical velocity test

Vch Maximal velocity with lactate steady state test

RPE Rating of Perceived Exertion

HLa Blood lactate

2OV! Oxygen Uptake

2OV! Peak Peak Oxygen Uptake

POaer Maximal Aerobic Power Output

ME Mechanical Efficiency

xiv

Glossary of terms

Term

Clarification

Quadriplegia. Condition resulting from SCI at the level of the cervical

vertebrae

Paraplegia Condition resulting from SCI at the level of the thoracic

vertebrae or below

Wheelchair /wheelchair

user system

Wheelchair and wheelchair user as one integrated unit

Wheelchair/wheelchair-

user interface

The point of integration between the wheelchair and the

wheelchair user e.g. Seat cage, push-rim and gloves

Manual wheelchair

propulsion

The act of locomotion in a push-rim wheelchair

Propulsive cycle The movements that bring about locomotion from hand

contact to subsequent hand contact at the start of the next

propulsive cycle

“propulsive” or “push”

phase

The period between the instant of hand contact to the

instant of release while the hand is in contact with the

push-rim

“non-propulsive” or

“recovery” phase

The period between the instant of release to the instant of

contact while the hand is not in contact with the push-rim

Total momentum The combined contribution of all body segments to

momentum of the system

Relative momentum The contribution of a particular body segment to the total

xv

momentum of the system

Transfer momentum The momentum that is transferred to a particular body

segment from the proximal segment

16

1. Chapter 1

1.1. Introduction

1.1.1. Wheelchair Sports and the Paralympic Games

Wheelchair sports were originally developed shortly after World War II by Sir

Ludwig Guttman and colleagues as a rehabilitation tool, a means to provide exercise

and recreation for young persons injured during the war. By 1952 the games had

developed into the first international wheelchair sporting competition for the

disabled. In the same year the International Stoke Mandeville Games Federation

(ISMGF) was formed to develop and govern wheelchair sports. The ISMGF later

established ties with the International Olympic Committee (IOC) and in 1960 the

first international games for the disabled held in conjunction with the Olympic

Games took place in Rome. During the 1964 Tokyo games the name “Paralympics”

was coined. Subsequently, the Paralympic Games have been held every four years.

1.1.2. British Paralympic success

Of all the 18 Paralympic sports wheelchair racing is arguably the most high profile

and, like mainstream athletics, sprint events take centre stage. Wheelchair sprinting

(events from 100 to 800 m) is also where Britain achieves most of its success in

international competition. British wheelchair athletes returned from the 1996

Paralympic Games in Atlanta, USA with nine medals. Two gold medals and new

17

World records (Tanni Grey, 800m, time: 1.55.12 mins and David Holding, 100 m,

time: 14.45 s), three silver medals (Tanni Grey, 100 m, 200 m and 400 m) and four

bronze medals (Nicola Jarvis, 100 m and 200 m, Paul Williams, 100 m and David

Holding 200 m). The success of British wheelchair athletes was shown to the world

thanks to the extensive media coverage of the 2000 Olympic and Paralympic Games

in Sydney, Australia. In the Paralympic Games British athletes finished second in the

medal table, only surpassed by the host nation. Great Britain’s athletes officially

became Britain’s most successful Paralympic Team ever. British wheelchair athletes

returned with seven medals. Five gold medals (Tanni Grey – Thompson 100 m, 200

m, 400 m and 800 m and Deborah Brennan 200 m) and Two bronze medals

(Deborah Brennan 200 m and David Holding 100 m). In addition Deborah Brennan

set a new World record over 200 m with a time of 33.87 s.

1.1.3. Wheelchair sprinting: Technical background

The goal of the wheelchair sprinter is the same as that of the sprint runner, which is

to cover the race distance in the shortest possible time. For the runner the race is

made up of a number of strides. Each stride can be broken down further into two

basic components, stride length and stride frequency. The same is true for the

wheelchair athlete, the race consists of a number of propulsive cycles consisting of a

push phase and a recovery phase. The push phase begins at the point of hand contact

with the push-rim. During the push phase the propulsive impulse that brings about

forward motion is imparted from the body to the push-rim. The recovery phase

begins at the point at which the hand releases the push-rim. The movements that

18

return the body to the point immediately before hand contact combine to make up the

recovery phase. The push phase can be broken down into pushing length (the

distance covered by the wheelchair with each push on the push-rim) and pushing

frequency (the number of pushes per unit of time). Walsh (1986) states wheelchair

velocity can only be increased through manipulation of one or both of these factors.

1.1.4. A deterministic model for wheelchair sprinting

The deterministic model for wheelchair sprinting (figure 1) identifies the key

components that determine the success of a wheelchair sprint athlete. As stated

previously the goal of the wheelchair sprinter is to cover the race distance in the

shortest possible time, therefore, the goal of the wheelchair sprinter is the

development of speed.

With the use of sophisticated laboratory based equipment sport scientists are able to

measure many of the components shown in figure 1 during simulated racing

wheelchair propulsion (RWP). Information relating to performance enhancement can

then be collated and disseminated to coaches and athletes. Unfortunately RWP

simulated in a laboratory environment is artificial compared to RWP in a competitive

environment (Vanlandewijck et al. 2001). RWP data collected in this artificial

environment provides a false description of RWP in a competitive environment and

therefore may not be directly applicable to enhance the performance of wheelchair

athletes. Scientists working to enhance the performance of wheelchair athletes must

19

develop methods of collecting data during over-ground RWP in competition in order

to gain an accurate picture of how wheelchair athletes propel their wheelchairs.

20

Figure 1 A deterministic model for wheelchair sprinting

Wheel Velocity

Point of Contact Point of Release

Contact Radius

Contact Time

Muscle CrossSectional Area

Activation Muscle Length

Total MuscleForce

Point of ForceApplication

Seating Position Joint Angles Segmental Lengths Pushrim Size

Segmental Motion

Direction

Direct PropulsionForce

Relative Momentumof Segments

Indirect PropulsionForce

Propulsive Impulse

Speed

Resistive Impulse

Friction Rolling Resistance

Mechanical Resistance

Wheelchair Athlete

Frontal Surface Area Coefficient of Drag Segmental Density Velocity

Drag Non-contact Time

21

1.1.5. Summary

The information above clearly identifies British wheelchair sprinting as being at the

forefront of international disability sport. However, at present the ability of the sport

scientist and coaches to further enhance the performances of these athletes is

hampered by methodological constraints. To ensure the continued success of British

wheelchair sprint athletes, equipment must be developed for the collection of data

during over-ground wheelchair sprinting.

A velocometer that could measure racing wheelchair velocity, would provide a

useful research tool in the study of propulsion technique. The device would allow the

velocity profile of the wheelchair to be constructed. The velocity profile would

provide information on the intra-push characteristics of propulsive cycle.

1.1.6. Aim

1. To design, produce and to test the utility of a velocometer to be used in the

assessment of intra-push changes in wheelchair velocity during over-ground

propulsion.

22

1.1.7. Objectives

1. To assess the functional requirements of the velocometer in relation to best

practice for the collection of data from wheelchair athletes.

2. To manufacture the velocometer in accordance with the functional requirements

assessed in objective 1.

3. To test the velocometer in accordance with the functional requirements assessed

in objective 1 by using the device to record the velocity profile of a racing

wheelchair during a sprint trial.

1.1.8. Hypothesis

The velocometer provides an accurate and reliable method for quantifying intra-push

changes in racing wheelchair velocity during over-ground propulsion.

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1.2. Literature review

This literature review is intended to provide the reader with a summary of the

findings of selected wheelchair related research. The literature under review covers

the period from the mid 1970’s, when manual wheelchair propulsion first became the

subject of scientific investigation, through to the present. In Sydney 2000 the world

witnessed the most integrated and successful Paralympic Games to date. Wheelchair

sport is now considered to be at the forefront of disability sport.

1.2.1. Inclusion criteria

The research reviewed in this section has been subjected to inclusion criteria. The

criteria are intended to ensure only studies that do not suffer from the major

limitations inherent in wheelchair related research are included. Preference has been

given to studies in which data has been collected from athletes, using their own

racing wheelchairs, during realistic simulated or actual over-ground manual

wheelchair propulsion. Where appropriate, only studies which have utilised over-

ground manual wheelchair propulsion or who have realistically simulated manual

wheelchair propulsion using a motor driven treadmill are included. Studies using

able-bodied subjects with little or no wheelchair experience have not been considered

for inclusion. Studies in which daily use, basketball or “active” wheelchairs,

interchanged between subjects, are also not included. Research findings related to

lever operated or hand crank wheelchairs has been excluded on the basis that manual

24

wheelchair propulsion is the most widely used method of locomotion for wheelchair

users.

1.2.2. Wheelchair related research

Previously the global aim of many researchers conducting wheelchair related

research has been to contribute to an improvement in the quality of life of lower limb

disabled persons who rely on wheelchairs for everyday mobility. However, many

researchers have used the growth and maturity of wheelchair sport as justification for

scientific investigation (Steadward and Walsh 1986). Cooper (1990c) states that in

recent years the progression of world records had slowed significantly, suggesting

that a point had been reached in terms of equipment and training at which small

differences become more significant. If continued improvements in wheelchair

racing are to be made, greater knowledge of the interaction between an individual

and their wheelchair will be required. To the sport scientist looking to enhance

performance the wheelchair/wheelchair-user system poses a similar problem to that

of any athlete whose interaction with a specific piece of equipment brings about a

sporting performance. Cooper (1996) states manual wheelchair research can be

divided into: design and testing; ergonomics and clinical assessment; physiology and

nutrition; and biomechanics.

For a comprehensive collection of wheelchair related research papers the reader is

directed to two published works edited by Woude et al. (1993) and Woude et al.

(1999). These compilations of wheelchair related research papers, based on the

25

proceedings of international workshops, show the variety and direction of wheelchair

related research in 1991 and 1999.

1.2.3. Wheelchair racing: development of a sport

In possibly the first study specifically targeting wheelchair racing, Higgs (1983)

characterised racing wheelchair construction in terms of success at the 1980 Olympic

games for the disabled. He found that the wheelchairs of more successful athletes

were characterised by lower seats, an increased seat angle to the horizontal, narrower

frame and smaller push-rims. In relative comparison the chairs used by the successful

sprinters had higher and more forward placed seats and a shorter chair length. No

significant differences in rear wheel camber were found.

Hedrich et al. (1990) provides an excellent description of the developments in

wheelchair racing between 1970 and 1990. Prior to the mid 1970s, wheelchair racing

existed as an accelerated version of conventional wheelchair propulsion mechanics.

The same wheelchairs used in everyday pursuits were used for sport (LaMere and

Labanowich 1984a). Recent advancements in wheelchair technology and training

have improved performance. However, the propulsion mechanics of wheelchair

racing have been dramatically altered (Higgs 1986; LaMere and Labanowich 1984a,

1984b, Sanderson and Sommer 1985, Steadward and Walsh 1986). Contemporary

wheelchair frames and wheels are built of aircraft quality alloys that are lighter and

stronger than steel or aluminium. Sealed precision bearings are now used in order to

26

reduce mechanical friction and in order to reduce rolling resistance, bicycle racing

wheels with narrow profiles and high pressure racing tyres are used.

To some degree the aerodynamic properties of the racer and the wheelchair have also

been addressed. Similar to cycling many wheelchair racers wear skin tight,

lightweight clothing to minimise aerodynamic drag. Athletes have chosen to reduce

the number of rear wheel spokes, adopt radial rather than crossing spoke patterns and

use flat rather than round spokes. These wheel modifications enhance the

aerodynamic properties of the racing wheelchair. Many athletes have adopted a

seating position with flexed upper trunk. Originally adopted because it assured upper

torso stability while concurrently allowing more severely disabled racers to push as

efficiently as their less disabled counterparts, athletes now believe that adopting this

position improves their propulsive efficiency and reduces drag.

1.2.4. Ergonomics

Woude et al. (1989a) described ergonomics as the “optimisation of human work”.

The ergonomic approach to the study of manual wheelchair propulsion seeks to

optimise the wheelchair-user interface, the fit between the wheelchair user and the

wheelchair itself. Cooper (1990c) states the seat cage and the push-rims are two of

the most critical interfaces between the individual and his/her racing wheelchair. The

seat cage provides support and stabilisation and determines body position with

respect to the push-rims. The efficiency of the force transference is dependent upon

the limb geometry with respect to the push-rim. The characteristics of the seat can be

27

broken down into position (in relation to the rear wheel axel and therefore the push-

rims, and height from the ground) and construction (upholstery). Seating can be

further broken down in terms of the angle of the base from the horizontal and height

of the backrest. Push-rims vary in the overall diameter, the diameter of the tubing

used in there construction, the distance they are mounted from the surface of the rear

wheels and the material covering the outer surface. These considerations have

obvious implications for the design of performance wheelchairs. In the design of

performance wheelchairs not only is the optimisation of the wheelchair-user

interface, maximising the ability of the athlete, a prime consideration but also the

performance characteristics of the wheelchair. Rolling resistance, internal friction

and aerodynamic drag must all be considered.

For most wheelchair athletes seating is highly individual. In most modern racing

wheelchairs the seat may be only a few pieces of strategically placed upholstery

strapped to the frame of the wheelchair. Similarly, the sizes of the push-rims are also

highly individual. Wheelchair athletes use push-rims that are of a smaller overall

diameter than those typically seen on “daily use” or “active” wheelchairs. The reason

is speed. Wheelchair athletes need to be able to accelerate their wheelchairs quickly

to top speed and then continue to propel them at a high percentage of that top speed

for the duration of the event. The size of the push-rim can be likened to the gearing

on a bicycle. The smaller the gear, the faster the bicycle will travel at any given

cadence.

28

1.2.4.1. Wheelchair-user interface: seat

The relationship between seat position and the biomechanics of manual wheelchair

propulsion has received great attention (Hughes et al., 1992, Mâsse et al., 1992,

Ruggles et al., 1994). Unfortunately a general lack of standardisation means that the

results of these studies are difficult to compare and generalise to other groups. It is

particularly difficult to infer useful information that can be applied to wheelchair

sprint athletes. Walsh et al. (1986) investigated the effect of seat position on maximal

linear velocity in wheelchair sprinting. The study utilised an adjustable wheelchair

fixed to a WERG to assess the effects of nine different seating positions believed to

cover the range of seating positions used by wheelchair athletes. The study found no

significant differences between the maximal linear velocities measured for each of

the nine seat positions. Meijs et al. (1989) investigated the effect of seat height on the

physiological response and propulsion technique in wheelchair propulsion. Meijs et

al. (1989) took into account the anthropometric dimensions of the nine male non-

wheelchair users in order to obtain better standardisation across trials. The study

found that seat height has a significant effect on physical load and propulsion

technique. The paper states that the reason some authors (Brattgård et al., 1970,

Brubaker et al., 1981, 1984) found no difference may have been due to the non-

standardisation of power output and seat height adjustment to individual’s

anthropometrical dimensions. Meijs et al. (1989) concluded the range in which the

wheelchair seat can be adjusted should cover an elbow angle of 100 to 120 °. The

author also states that the results may underline the importance of adjusting

wheelchair dimensions to the anthropometric characteristics of the user. These results

are similar to a previous study conducted by Woude et al. (1989a). Woude et al.

29

(1989a) indicated that, based on comparative physiological responses to propulsion,

the optimum angle of elbow flexion, is between 100 and 120 °. To date no studies

have successfully identified an optimal seating position for wheelchair sprint

athletes.

1.2.4.2. Wheelchair-user interface: push-rim

Gayle et al. (1990a) investigated the effect of two different sized push-rims (0.25 and

0.41 m overall diameter) on cardiorespiratory and perceptual responses to wheelchair

propulsion. Fifteen male paraplegics (3 track athletes, 12 recreational athletes)

performed three discontinuous laboratory based exercise tests and two 1600 m

performance based track trials. A racing wheelchair (Stainless Medical Products

Racer, San Diego, CA), modified for use with each subject, was used for the entire

series of laboratory and track based trials. The results reported no significant

differences in HR, 2OV! , VE, HLa or RPE using different sized push-rims at 4 km.h-1.

At 8 km.h-1 subjects demonstrated a 13 % lower steady state 2OV! (p<0.05) using the

0.25 m push-rims. HR was not significantly different. Under simulated race

conditions on an all weather track no significant differences were found for HR,

performance time, or RPE between trials. HLa was significantly lower using the 0.25

m push-rims. The authors concluded that although the data identified few significant

differences in the physiological responses between trials, there was a tendency for a

lower metabolic stress using the smaller push-rims.

30

Woude et al. (1988b) investigated the effects of five different diameter (0.3, 0.35,

0.38, 0.47 and 0.56 m) push-rims and varying speeds on a number of physiological

and kinematic variables. Eight wheelchair sportsmen (6 SCI [T2-LS], 1 Spina Bifida,

1 AB) used similar racing wheelchairs (weights ranged from 11 to 13 kg, rear wheel

camber 8.5°, tyre pressure standardised). The push-rims used all had a similar grip

profile and were constructed of 0.03 m tubing taped with soft plastic. Five

progressive exercise tests were randomly spaced on three subsequent days. Each test

consisted of five 3 min stages on a MDT. Tests were performed with a constant

treadmill inclination of 0.5 °. Belt velocity was increased by 0.83 m.s-1 every three

minutes. Speed ranged from 0.83 to 4.17 m.s-1.

The authors conclude, in terms of 2OV! , VE, HR, and gross ME; a smaller diameter

push-rim is more advantageous during high-speed wheelchair propulsion. Despite

inter-individual variation in movement technique and timing pattern, general patterns

of adaptation to rim diameter and wheelchair velocity were evident. Different push-

rim diameters were shown to lead to systematic shifts in the trajectories of the upper

arm, whereas no changes in timing parameters, push angle, and work per cycle were

seen. These findings may explain the increased cardiorespiratory stress observed at a

given velocity when using larger push-rims.

Based on a survey performed during the 1980 Olympics for the Disabled, Woude et

al. (1988b) stated that high level performance in wheelchair racing may be associated

with lower and more inclined seats, increased rear wheel camber, and smaller push-

rims. However, with the exception of Walsh et al. (1986), who reported the effect of

seat height on sprint performance, there is still a general lack of information

31

regarding optimum wheelchair-user interface characteristics for wheelchair sprint

athletes.

1.2.4.3. Manual wheelchair propulsion daily use vs. sport

Boninger et al. (1998) states that the nature of wheelchair propulsion means manual

wheelchair users are essentially walking with their arms. The upper extremity,

particularly the shoulder, is designed for freedom of movement and not repetitive

loading. Boninger et al. (1997) elaborated further. In order to propel a wheelchair a

force must be imparted to the push-rim. This force is analogous to the highly studied

ground reaction force of gait. The forces imparted to the push-rim are equally and

oppositely transmitted back to the upper limb of the wheelchair user. It is likely that

these joint reaction forces are responsible, in part, for a large majority of upper limb

injuries occurring in manual wheelchair users. Cooper (1990c) commented on the

increased demands of manual wheelchair propulsion in the sporting environment

compared to “daily use”. This seems logical when we consider the increased speed

and force requirements of accelerating a wheelchair and propelling a wheelchair at

high speed.

1.2.5. Assessment of athletic wheelchair performance

With the growth and maturity of wheelchair sport, practitioners began to train and

develop themselves in accordance with the general training principles of athletic

32

performance. The scientific community has taken an interest in wheelchair sports

persons. Equipment such as wheelchair ergometers (WERGS) were developed

(Glaser et al., 1978, Niesing et al., 1988, 1990, Vosse et al., 1990) and motor driven

treadmills (MDTs) were modified to accommodate wheelchairs (Horvat et al., 1984,

Claremont et al., 1985, Lakomy et al., 1987). Tables 2, 3 and 4 indicate the

prevalence of WERGs in relation to MDTs and over-ground manual wheelchair

propulsion in the manual wheelchair propulsion literature. Physiological testing

protocols were modified and tested with wheelchair users (Hartung et al., 1993,

Rasche et al., 1993, Goosey et al., 1995). The training practices of wheelchair

athletes were evaluated (Campbell et al., 1997) and investigated in relation to the

physiological characteristics of able-bodied athletes (Lakomy et al., 1987). The

laboratory based physiological testing of wheelchair athletes is now common.

However, Vanlandewijck et al. (2001) have called into question the realism of some

of the methods used to simulate manual wheelchair propulsion in the laboratory.

1.2.5.1. Simulated wheelchair propulsion under realistic conditions

The propulsive cycle has been the focus of many research studies. Like the running

stride the propulsive cycle has been broken down in terms of a contact, often termed

“propulsive”, and a non-contact, often termed “recovery”, phase. The contact phase

refers to the period between the instant the hand contacts the push-rim until the

instant the hand leaves the push-rim. The non-contact phase refers to the period

between the instant the hand leaves the push-rim until the instant before the hand

contacts the push-rim at the start of the next propulsive cycle.

33

These definitions have become standard terms in the manual wheelchair propulsion

literature. A wealth of research has been performed using these definitions.

Unfortunately this seems to have had the effect of simplifying manual wheelchair

propulsion research into an investigation purely of arm work, neglecting the

contribution of the head and trunk at a fundamental level. This is particularly

important in racing wheelchair propulsion in which the motion of the trunk and head

are exaggerated.

Vanlandewijck et al. (1994) provided an intra-push description of manual wheelchair

propulsion. The authors noted a twofold acceleration in the velocity curve of the

wheelchair-user system at 2.22 m.s-1. Propulsive forces acting on the push-rims were

responsible for acceleration during the propulsive phase. During the recovery phase,

a second, smaller acceleration was observed. This second acceleration was due to

experienced subjects accelerating their trunk and/or arms backward causing reaction

forces to act on the wheelchair. These actions delayed deceleration of the wheelchair.

The above findings demonstrate that wheelchair propulsion at velocities typically

observed in wheelchair racing do not consist of an “active” period (the propulsive

phase) and a “passive” period (the recovery phase) as argued by Veeger et al.

(1992b). The author’s state that wheelchair propulsion in experienced wheelchair

racers consists of three periods, each of which has specific energy demands. 1) An

acceleration period which occurs due to the forces applied to the push-rims; 2) A

second, smaller, acceleration period due to inertial forces acting on the wheelchair-

user system. This is caused by the backward trunk and/or arm- swing described

above; and 3) A deceleration period due to resistive forces acting on the wheelchair-

user system, caused by an increased forward segmental velocity in order to make

34

contact with the rims with increased hand speed. Vanlandewijck et al. (2001)

provides a similar description stating that manual wheelchair propulsion consists of:

1) An acceleration phase caused by forces applied to the push-rims, 2) A second

acceleration phase caused by the inertial forces acting on the wheelchair-user system,

caused by a backward arm and/or trunk swing and 3) A deceleration phase during the

second part of the recovery phase.

1.2.5.1.1. Wheelchair ergometers (WERGs)

Wheelchair ergometers are commonplace in manual wheelchair propulsion research.

According to Glaser et al. (1977), Arabi et al. (1997) and Bhambhani et al. (1991),

the use of wheelchair ergometry in the study of the physiology of manual wheelchair

propulsion with paraplegic and quadriplegic subjects is reliable and valid. Arabi et

al. (1997) examined the relationship between maximal oxygen uptake on a MDT and

WERG and concluded that the data obtained were similar and significantly correlated

when expressed in either l.min-1 or ml.kg.min-1 (1.25 ± 0.38 and 1.22 ± 0.28 l.min-1

or 19.5 ± 6.14 and 18.18 ± 4.27 ml.kg.min-1) for MDT and WERG respectively.

However, significant differences were found in maximal speed between the MDT

and WERG. This was probably due to the mechanical resistance of the rollers used in

the construction of the WERG. Bhambhani et al. (1994) performed a comparison

between simulated wheelchair racing on a WERG and track racing. The study

concluded that simulated wheelchair racing on a WERG is a valid measure of track

racing performance in male paraplegic and quadriplegic athletes. Generally speaking

35

the use of WERGs for physiological assessment is acceptable as the device is

bringing about a physiological response to a given workload.

Tables 2, 3 and 4 provide the reader with a comparison between studies that have

chosen to use WERGs, MDTs or over-ground manual wheelchair propulsion during

data collection. Table 1 provides the key to the wheelchair coding used in tables 2, 3

and 4 The main advantage of WERGs are that they can be used to simulate manual

wheelchair propulsion in a controlled laboratory environment. The laboratory

environment affords the researcher far greater opportunity for measurement,

unfortunately this is at the cost of realism. Wheelchair ergometers exist in two

common forms. 1) WERGs constructed as an approximation of a wheelchair with the

wheels and the seat mounted separately (Niesing et al., 1988, 1990, Vosse et al.,

1990). 2) WERGs incorporating either single (Goosey et al., 1998a) or twin

(Shimada et al., 1995) rollers on which the subject’s own wheelchair can be

mounted. In the table the former is indicated by an asterix after the study reference.

This type of WERG usually affords more sophisticated measurements due to the

independent mounting and therefore ease of instrumentation of the wheels and seat.

The latter addresses important issues relating to the wheelchair-user interface by

allowing the wheelchair user’s own wheelchair to be used during the testing. Goosey

et al. (1998b) indicates the importance of testing athletes in their own racing

wheelchairs stating that through training athletes become tuned to their own racing

wheelchairs.

From the point of view of realistically simulating manual wheelchair propulsion, in

relation to the use of WERGs, two main problems have to be overcome. These relate

36

specifically to the fact that during manual wheelchair propulsion the wheelchair/user

system is fixed in a stationary position.

1) The influence of the HAT motion on wheelchair motion during the recovery

phase when the hands are not in contact with the push-rim.

2) The effect of wind resistance and other environmental factors on the

metabolic cost of wheelchair propulsion and the variation with speed.

Writing specifically about the use of WERGs in anaerobic testing, Vanlandewijck et

al. (2001) highlights another important limitation. Backwards tilting is prevented on

most WERGs. For this reason the forces generated on the push-rims will be much

higher compared with the same task performed under field conditions.

1.2.5.1.2. Motor driven treadmills (MDTs)

MDTs are also common in manual wheelchair propulsion related research. However,

as tables 2, 3 and 4 indicate, MDTs are used less frequently compared to WERGs.

While many MDTs are now specifically manufactured for use in manual wheelchair

research with longer and wider treadmill belts and specific safety devices, much of

the early research was conducted on MDTs designed for runners, modified for use

with wheelchair users. A commonly held opinion is that by using MDTs many of the

disadvantages associated with the use of WERGs can be overcome. While this may

37

be true to a certain extent, MDTs have limitations when compared to over-ground

manual wheelchair propulsion.

MDTs allow accurately simulated manual wheelchair propulsion to be performed in

the laboratory environment. As stated previously, the laboratory environment affords

the researcher far greater opportunity for measurement than field based data

collection. However, A wheelchair fixed to a MDT is no different to a wheelchair

fixed to a WERG. The wheelchair must be fixed to the MDT in such a way that the

wheelchair is allowed to run freely along the whole length of the treadmill belt

(Horvat et al., 1984, Claremont et al., 1985, Lakomy et al., 1987). This allows the

wheelchair to accelerate and decelerate with the natural rhythm of propulsion.

Wheelchair ergometers and MDTs share one limitation in relation to the realistic

simulation of manual wheelchair propulsion. The effect of wind resistance and other

environmental factors on the metabolic cost of wheelchair propulsion, and their

variation with speed. This question has been addressed with respect to runners. Jones

and Doust (1996) state that a 1 % treadmill grade most accurately reflects the

energetic cost of outdoor running. However, to the best of the author’s knowledge

this has not been thoroughly researched with respect to simulated manual wheelchair

propulsion on the treadmill.

1.2.5.1.3. Over-ground manual wheelchair propulsion

The use of over-ground manual wheelchair propulsion provides the investigator with

the opportunity to study realistic propulsion. This is important in the study of manual

38

wheelchair propulsion kinematics and particularly important in the study of racing

wheelchair propulsion. The only limitation of using over-ground manual wheelchair

propulsion is the level of measurement that can be achieved. It is very difficult to

combine the realism of over-ground manual wheelchair propulsion and the carefully

controlled sophisticated measurement environment of the laboratory. At present the

level of measurement afforded by the laboratory environment cannot be replicated

when performing over-ground manual wheelchair propulsion trials. This is the reason

for the dearth of studies using over-ground manual wheelchair propulsion.

39

Table 1 Wheelchair coding for tables 2, 3 and 4

Wheelchair Code Wheelchair Code

Daily use 1 Traveller 11

Crank 2 Active or sport wheelchair 12

Synchronic lever 3 Basketball wheelchair 13

Fully adjustable 4 Racing wheelchair 14

Quickie GPV 5 Three wheeled racing wheelchair 15

Quickie 2HP 6 Four wheeled racing wheelchair 16

Quickie I 7 Own seat cushion used a

Quickie II 8 Personal wheelchair P

Premier II 9 Standard wheelchair S

Morrien Tornado 10 Wheelchair ergometer WERG

40

Table 2 Studies using a wheelchair ergometer to simulate manual wheelchair propulsion

Origin Study/Studies

Further studies using WERG

Brief description of WERG

Wheelchair Special features of WERG

Type of drive Type of resistance control Additional biomechanical data collection

Brattgård et al. (1970)*

None Platform with separately mounted adjustable seat and wheels

S Adjustable seat and wheels

Chain to Monark Flywheel friction brake None

Stoboy et al. (1971)

None Wheelchair driving platform equipped with rollers

P None Not stated Not stated None

Wicks et al. (1977, 1983)

None Design based on Brattgård et al. (1977). Combination wheelchair-cycle ergometer adapted to allow arm cranking

S None Chain to Monark Flywheel friction brake. Direct current generator attached to ergometer drive shaft to measure wheelchair strike frequency

None

Glaser (1977)* Combination wheelchair-cycle ergometer.

S None Chain to Monark Monark flywheel and belt with adjustable resistance via screw mechanism

None

Glaser et al. (1978, 1979)

" " " " " "

Glaser et al. (1980)

Modified to allow arm cranking

" " " " "

Brown et al. (1990)

" " " " " 2D analysis with a high speed camera

41

Table 2 Continued






Lundberg (1980) None Two cycle training rollers placed side by side with holding frame to steady front wheels

P None Direct to rollers

Not stated None

Ross and Brubaker (1984)*

None Motor compensated wheelchair dynamometer with independent bi-lateral inputs

Not stated Ability to sample dynamometer, push-rim torque, and velocity

Not stated Not stated EMG. Neuromuscular stimulator

Walsh et al. (1986)

None Custom made ergometer

S4 None Direct to rollers

Not stated 2D analysis

Burkett et al. (1987)*

None Hysterisis brake ergometer. Wheels and seat independently mounted on instrumented frame

WERG Horizontal seat adjustment

Wheels mounted on central drive shaft

Hysterisis brake None

Coutts and Stogryn (1987)

None Twin roller wheelchair ergometer

P12 Resistance and distance measurement

Direct to rollers

Torque wrench and electric motor

None

42

Table 2 Continued






Eriksson et al. (1988)

Custom designed frictionless roller ergometer with side mounted flywheels

P None Direct to rollers Flywheel None

Bhambhani et al. (1991)

" " " " " "

Lees and Arthur (1988)

None Twin roller wheelchair ergometer. Computer interfaced

P None Direct to rollers Weighted flywheel friction brake

None

Niesing et al. (1988 - Conference proceedings, 1990)*

Sophisticated computer controlled ergometer. Wheels and seat independently mounted

WERG Highly adjustable for investigation of wheelchair-user interface. Isokinetic and isometric force measurement

Wheels mounted independently

Motor controlled None

Woude et al. (1989b)

" " " " " "

Veeger et al. (1991b)

" " " " " EMG. 3D mirror analysis using a high speed camera

Veeger et al. (1991c)

" " " " " "

43

Table 2 Continued







Niesing et al. (1988 - Conference

proceedings, 1990)*

" Niesing et al. (1988 -

Conference proceedings,

1990)*

Niesing et al. (1988 - Conference proceedings,

1990)*

Niesing et al. (1988 -

Conference proceedings,

1990)*

Niesing et al. (1988 - Conference proceedings,

1990)*

EMG. 3D mirror analysis using a high speed camera

Veeger et al. (1992a, b, c)

" " " " " 2D analysis using high speed camera

Janssen et al. (1993)

" " " " " None

Woude et al. (1994)

" " " " " "

Dallmeijer et al. (1994, 1998)

" " " " " 2D analysis

Helm et al. (1996)

" " " " " EMG. 3D mirror analysis

Linden et al. (1996)


Dallmeijer et al. (1996), Woude et al. (1997, 1998)

" " " " " None

44

Table 2 Continued






Rozendaal et al. (2000)

See Niesing et al. (1988 - Conference proceedings, 1990)*

See Niesing et al. (1988 - Conference proceedings,

1990)*

See Niesing et al. (1988 - Conference proceedings, 1990)*

See Niesing et al. (1988 - Conference

proceedings, 1990)*

See Niesing et al. (1988 - Conference proceedings,

1990)*

3D analysis

Hughes et al. (1989, 1992)*

None Computer aided wheelchair data acquisition and physical simulator. Wheels and seat independently mounted on instrumented frame

S 0.35 m variation in seating position in each of the three orthogonal planes. Ability to record kinematics of trunk, shoulder, elbow and wrist. 8 channel EMG facility

Wheels mounted on central drive shaft

Not stated None

Samuelsson et al. (1989)*

None Frame mounted wheelchair connected to a Cybex II isokinetic dynamometer

S None Chain to Cybex 1/1 ratio

Cybex II None

Gehlsen et al. (1990)

None Pro Roller. Tach-generator interfaced to an Apple computer

P14 None Direct to rollers Not stated 2D analysis using a high speed camera

45

Table 2 Continued






Cooper (1990c) None Internal roller system equipped with a Maxon analogue tachometer and a Tektronix analogue data recorder

P14 None Direct to rollers Inertia adjustment Video records obtained

Gayle et al. (1990a, b)

Commercially available wheelchair roller with added electronic speedometer and wheel revolution counter

S14 None Direct to rollers Friction mechanism None

Rodgers et al. (1994)

" S12 " " " 3D motion analysis. EMG. Wheelchair instrumented with a force-measuring push-rim and potentiometers in the wheel hubs

46

Table 2 Continued






Rodgers et al. (1998)

See Gayle et al. (1990a, b)

S6 See Gayle et al. (1990a, b)

See Gayle et al. (1990a, b)

See Gayle et al. (1990a, b) 3D motion analysis. Wheelchair instrumented with AMTI multicomponent force/torque transducer

Vosse et al. (1990) Sophisticated computer controlled roller ergometer using a Proportional, Integral and Derivative (PID) controller

P Ability to simulate road/track conditions

Direct to rollers PID None

Robertson et al. (1996)

" S7 " " " SMARTwheel

Cooper et al. (1996)

" S5 " " " 3D motion analysis. SMARTwheel

Boninger et al. (1997), Shimada et al. (1998)

" S " " " "

Cooper et al. (1997)

" S7 " " " "

Boninger et al. (1998)

" S12 " " " "

47

Table 2 Continued






O’Connor et al. (1998), DiGiovine et al. (2000)

See Vosse et al. (1990)




See Vosse et al. (1990) 3D motion analysis

Mâsse et al. (1992)

None Commercially available wheelchair roller. Iron rings added to roller to increase inertia

S14 None Direct to rollers Not stated 3D mirror analysis. EMG

Cooper et al. (1992) (Conference proceedings)

CSUS Dynamometer (No description)

S7 Not stated Direct to rollers Not stated 2D analysis using two cameras. Modified three channel version of SMARTwheel

Asato et al. (1993)

" " " " " "

Meijs (1993) Motor driven single roller ergometer. Computer interfaced

P13 Continuous determination of torque

Direct to roller Electrically braked None

Hutzler et al. (1995)

" " " " " "

Woude et al. (1995)

" S15 " " " "

48

Table 2 Continued







Specially constructed, low friction steel roller system. Computer interfaced

P None Direct to rollers Not stated None


See Bhambhani et al. (1994)

P14 See Bhambhani et al. (1994)

See Bhambhani et al. (1994)

See Bhambhani et al. (1994) See Bhambhani et al. (1994)

Ruggles et al. (1994)

Two aluminium rollers connected to a Cybex II isokinetic dynamometer

S9, S8, S5 Angular position and torque measurement

Direct to rollers, rollers connected by chain to Cybex

Cybex II None

Davis et al. (1998)

" S11, S8 " " " 3D analysis

Wang et al. (1995)

Eagle roller with adjustable friction

P15 None Direct to rollers Adjustable friction control 3D mirror analysis using a high-speed camera. Electronic timing device to detect contact with the push-rim

Wang et al. (1996)

" P16 " " " Electronic timing device to detect contact with the push-rim

Chow et al. (2000, 2001)

" P14 " " " 3D analysis. EMG

Shimada et al. (1995)

Two-roller ergometer, electronically braked. Computer interfaced

P Torque measurement Direct to rollers Two independently wired single input electronic loads

None

49

Table 2 Continued






Koontz et al. (2001)

See Shimada et al. (1995)




See Shimada et al. (1995) Bilaterally mounted SMARTwheel

Mulroy et al. (1996)

Specially designed frame and split-roller drive assembly. Computer interfaced

S5a None Direct to rollers Inertia adjustment with removable flywheels proportional to the weight of the subject and the wheelchair

Wheelchair wheel instrumented with strain gauge force transducers. EMG

Newsam et al. (1996)

See Mulroy et al. (1996)

S5a See Mulroy et al. (1996)

See Mulroy et al. (1996)

See Mulroy et al. (1996) SMARTwheel

Rao et al. (1996), Kulig et al. (1998, 2001), Newsam et al. (1999)

" S5 " " " 3D analysis. SMARTwheel

Theisen et al. (1996)

Two interconnected rollers. Computer interfaced

S13 WERG has ability to simulate propulsion on inclines

Direct to rollers Electronic brake (Merobel) None

Arabi et al. (1997)

" S " " " Maximum voluntary force on push-rim measured using a strain gauge transducer

Goosey et al. (1998a)

Single roller ergometer. Computerised interfaced. Optical sensor used to count roller revolutions

P15 None Direct to roller Belt from roller drives fan 2D analysis

50

Table 2 Continued






Goosey et al. (1998c)

See Goosey et al. (1998a)




See Goosey et al. (1998a) 3D analysis

Goosey et al. (2000)


Goosey-Tolfrey et al. (2001)

" S15 " " " 2D analysis. On-line system tracking hand path. Manchester Metropolitan University force-measuring push-rim device

Malone et al. (1998)

None Commercially available roller system

S13 None Direct to rollers Not stated 3D analysis

51

Table 3 Studies using a motor driven treadmill to simulate manual wheelchair propulsion


Further studies using MDT

Brief description of MDT

Wheelchair Special features of MDT

Type of drive Type of resistance control

Additional biomechanical data collection

Engel and Hildebrandt (1973)

None Purpose built treadmill-ergometer

P1 None Direct to treadmill belt

None None

Gass and Camp (1979)

None MDT (No description) P None Direct to treadmill belt

None None

Gass and Camp (1984)

" " " " " None

Sanderson and Sommer (1985)


None 2D analysis

Woude et al. (1986)

Enraf Nonius, model 3446.


None Force transducer used to measure drag force

Woude et al. (1988a)

" P12, P13 " " " Force transducer used to measure drag force

Woude et al. (1988b)

" S14 " " " 2D analysis using a high-speed camera. Force transducer used to measure drag force

Meijs et al. (1989)

" S10 " " Pulley mechanism for normalisation of power output

EMG. Force transducer used to measure drag force

Veeger et al. (1989a)

" " " " " EMG. 3D mirror analysis using a high speed camera

Veeger et al. (1989b)

" S13 " " " 2D analysis using a high speed camera

52

Table 3 Continued




Wheelchair Special features of MDT


Woude et al. (1989c)

See Woude et al. (1986)

S10 See Woude et al. (1986)

See Woude et al. (1986)

Pulley mechanism for normalisation of power output

EMG. 3D mirror analysis using a high speed camera

Veeger et al. (1992a)

" " " " " Force transducer used to measure drag force


" Not stated " " " None

Rasche et al. (1993)

" P1 " " " Force transducer used to measure drag force


" S1 " " " "


" S2, S3, S1, S12 " " Not stated "

Woude et al. (1994)

" S10 " " Pulley mechanism for normalisation of power output

"

Lakomy et al. (1987)

Woodway model ELGZ adapted for wheelchairs

P None Direct to treadmill belt

None None

Campbell et al. (1997)

" P14 " " " "

Pitetti et al. (1987)


None None

53

Table 3 Continued




Wheelchair* Special features of MDT


Hartung et al. (1993)

None MDT (No description)

S None Direct to treadmill belt

None None

Vanlandewijck et al. (1994)


S5 None Direct to treadmill belt

Pulley mechanism including load cell for normalisation of power output

3D analysis using two video cameras. EMG

Spaepen et al. (1996) " S6 " " " " Goosey et al. (1995)

Woodway model ELGZ adapted for wheelchairs


None 2D analysis

Goosey et al. (1998b) " " " " " " Tropp et al. (1997)


P None Direct to treadmill belt

None Force transducer used to measure drag force

Arabi et al. (1997)

None Specially constructed MDT.

S None Direct to treadmill belt

None None

Arabi et al. (1999) " P14 " " " Maximal voluntary force measured by strain gauge transducer

54

Table 4 Studies employing over-ground manual wheelchair propulsion

Study Brief description of test Wheelchair Propulsion surface Additional biomechanical data collection

Higgs (1983) 400 m, 800 m and 1500 m P16 Outdoor running track 2D photographical analysis of wheelchairs from the front and rear

Higgs (1986) 200 m and 1500 m P14 Outdoor running track 2D Cine analysis Ridgeway et al. (1988) 800 m P14 Outdoor running track 2D analysis

Coutts and Schutz (1988) 100 m, 200 m, 400 m, 800 m, 1500 m, 5000 m and marathon

P14 Outdoor running track None

Lees and Arthur (1988) 100 m, 200 m and 400 m P14 Outdoor synthetic track None

Hedrich et al. (1990) Coast down trials P14 Smooth concrete apron around an indoor running track

Frontal cross-sectional body area.

Nadeau et al. (1990) 30 m sprint P14 Outdoor running track Motion detectors. Touch pad commenced data collection

Gayle et al. (1990) 1600 m track trials with two (10 inch and 16 inch) sized push-rims

P14 Outdoor running track None

Coutts (1991) Coast down trials S12 Hard-wood gymnasium floor

Wheelchair instrumented with a magnetic switch

55

Table 4 Continued

Study Brief description of test Wheelchair Propulsion surface Additional biomechanical data collection

Coutts (1992) Coast down trials S12 See Coutts (1991) See Coutts (1991) Coutts (1994) Coast down trials. Sprint

trials P13 " "

Bednarczyk and Sanderson (1994) Steady state propulsion S12 Long strips of smooth canvas placed on a wooden gymnasium floor

3D analysis. Hand switch to determine contact and release

Janssen et al. (1994) Activities of daily life P Not stated None Goosey et al. (1997) 800 m P14 Outdoor running track 2D analysis Vinet et al. (1998) Coast down trials P1 Tartan track field Deceleration profile calculated from video

recordings Arabi et al. (1999) Montreal progressive Track Test

(MTT) P Outdoor running track Maximal voluntary force measured by strain gauge

transducer

56

1.2.5.1.4. Protocols

Incremental testing protocols for wheelchair athletes are numerous. The protocols

used for testing on MDTs vary in terms of increments in speed (Lakomy et al., 1987,

Campbell et al., 1997, Arabi et al., 1999), increments in speed and % grade (Gass

and Camp 1979, 1984, Woude et al., 1986, 1988a, Hartung et al., 1993, Arabi et al.,

1997, Tropp et al., 1997, Goosey et al., 1995, 1998b) and increments in power output

(Rasche et al., 1993). Similar to the variety of protocols used for testing wheelchair

athletes on MDTs, WERG testing protocols share numerous variations. This is

largely due to the wide variety and complexity of their construction. Simple, single

or twin roller WERGs lend themselves to increasing speed protocols measured in

m.s-1 (Goosey et al., 1998c, Theisen et al., 1996), km.h-1 (Bhambhani et al., 1994,

1995), RPM (Coutts and Stogryn 1987) or wheel strike rate (Bhambhani et al.,

1991). Wheelchair ergometers constructed by connecting wheelchairs to Monark

cycle ergometers increase workload intensities at a constant cadence by increasing

the resistance on a flywheel. This is achieved by adding mass to a metal basket or by

tightening a screw mechanism (Brattgård et al., 1970, Wicks et al., 1977, 1983,

Glaser 1977, Glaser et al., 1978, 1979, 1980). More complex, computer interfaced,

designs use increasing increments of resistance through electronic braking to

measure power output in watts (Burkett et al., 1987, Niesing et al., 1988, 1990, Meijs

1993). Testing during over-ground racing wheelchair propulsion is limited due to the

level of measurement afforded by the environment. However, Arabi et al. (1999)

investigated the feasibility and practicality of performing a number of laboratory

based tests in the field. Despite the use of these varied protocols, few investigators

57

have investigated the possibility of an optimal protocol for testing wheelchair

athletes or looked at standardising the protocols used.

Woude et al. (1988a) investigated the effect of two workload strategies, 1)

Increments in velocity at a constant slope and 2) Increments in slope at a constant

velocity, using eight wheelchair marathon racers and basketball players in a standard

wheelchair. Woude et al. (1988a) justify the use of a MDT by citing the opinion of

Schenau (1980) that there is no actual mechanical difference between treadmill and

over-ground locomotion. Woude et al. (1988a) provide further justification stating

that Bassett et al. (1985) reported no variation in oxygen consumption between over-

ground and treadmill running at 0 and 5.7 % slope within the velocity range tested.

The authors reported no strategy effect in the cardio-respiratory parameters

mechanical efficiency (ME), ventilation rate (VE), oxygen uptake ( 2OV! ) and heart

rate (HR). However, the authors did report that the duration of the propulsive and

recovery phases appeared highly dependent on speed and slope respectively. Veeger

et al. (1989b) found that the duration of the propulsive cycle and recovery phase

were shorter for steeper slopes. Vanlandewijck et al. (2001) reported confirmation of

these findings with slope gradients between 1.5 to 6 %. Woude et al. (1988a) also

suggest that 3 minute stages appeared sufficiently long for experienced wheelchair

users to adapt to a given speed and slope combination.

Hartung et al. (1993) investigated the effect of three workload strategies: Increments

in velocity at a constant slope (S); increments in slope at a constant velocity (G); and

progressive increments in speed and slope (C), using seven wheelchair racing and

games players in a standard wheelchair. The authors reported that treadmill test

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protocols similar to (C) might be the optimal method. Variations in the kinematics of

manual wheelchair propulsion with each of the three protocols were not studied.

Hartung et al. (1993) voiced concerns about realistic testing protocols for wheelchair

athletes stating, a treadmill protocol using only increments in speed may be

unsatisfactory for athletes in order to elicit maximal responses for safety reasons. It is

this opinion that has fostered the use of inclined MDTs in the assessment of racing

wheelchair propulsion. The inclusion of gradients in racing wheelchair testing

increases the physiological response at any given speed (Goosey et al. 1995). This

reduces the need for tests at high velocities because data concerning maximal

performance ( 2OV! Peak, Power, speed or velocity at 2OV! peak, HR, respiratory

variables and blood lactate accumulation) can be at collected at lower velocities.

In response to safety concerns about physiological testing using MDTs at realistic

race speeds Goosey et al. (1995) studied the efficacy of using a 0.7 % treadmill

gradient in eliciting selected physiological responses at slower treadmill speeds using

11 wheelchair athletes in their own racing wheelchairs. Significant (p<0.01)

increases in HR, oxygen consumption and blood lactate were observed. The increase

in treadmill grade resulted in adaptations in the temporal data rather than the

displacement data. The cycle dynamics, cycle time and the number of pushes per

minute, were higher when the grade of the treadmill was increased (p<0.05 and

p<0.01, respectively). The increase in % grade was accompanied by a mean

reduction of 0.4 s in cycle time and an increase of 29 pushes per minute. The authors

concluded that a 0.7 % increase in gradient is sufficient to stimulate an increased

physiological demand without significantly affecting the movement pattern of

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wheelchair propulsion. The authors also state that this test protocol may be

recommended to examine the physiological and wheelchair propulsion techniques of

the athletes in their own racing wheelchair at realistic speeds

Both Goosey et al. (1995) and Woude et al. (1988a) found variation in the

kinematics of MWP when using gradients during treadmill testing. Therefore, it may

be questioned whether the kinematics of racing wheelchair propulsion on an inclined

treadmill are truly representative of over-ground racing wheelchair propulsion?

Woude et al. (1986) described a method of measuring the force Fd (resistance force

at a constant speed), made up of internal friction, rolling friction and a gravity

component but independent of velocity. The method of measuring the drag or

resistance force described involves performing a drag test during which the subject

remains passive in the wheelchair while it is moved by the treadmill at a constant

velocity. The force is measured using a force transducer fixed “in line”, on the drag

cable, between the wheelchair and the fixed point on the treadmill where the drag

cable attaches.

Woude et al. (1989a) pioneered a pulley system, off which various masses could be

hung, to standardise power output during manual wheelchair propulsion on an MDT.

Studies that have used this system can be seen in Table 3. The system works by

attaching a cable to the rear of the wheelchair and over a pulley suspended at the rear

of the treadmill. The other end of the cable is attached to a mass hanger, which is

suspended below the pulley. A standard power output is achieved first by measuring

Fd for all subjects, as described above, and then adding various masses in the mass

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hanger. By using these methods the physiological characteristics of racing

wheelchair propulsion can be measured in standardised trials without the use of

inclined treadmills.

Rasche et al. (1993) used the pulley mechanism described above, to increase the

intensity of trials at a constant velocity, during a study conducted to compare a

discontinuous (DP) and continuous-jump maximum oxygen uptake protocol (JMP) in

maximal wheelchair exercise on a treadmill. The DP protocol involved three minute

stages followed by two minutes relative rest. The JMP protocol involved increasing

power output via the pulley mechanism every minute. The paper concluded that both

the DP and JMP protocols were equally appropriate in determining 2OV! peak and

power output at 2OV! peak.

Arabi et al. (1999) investigated the feasibility of three tests, the Montreal progressive

Track Test (MTT), Critical velocity test (Vc) and maximal velocity with lactate

steady state (Vch), previously used in the assessment of runners. The MTT and Vc

were feasible in that the MTT could be performed, and Vc, determined, in the field.

The authors state that the measurements of Vch could not be used because of “many

absurd results”, (p. 489). A second study in a laboratory showed that the concept of

critical velocity and critical power could be used in wheelchair testing on a treadmill.

Similar to the test described by Woude et al. (1986), Vinet et al. (1998) described a

test for the measurement of drag or resistance force, which could be administered in

the field.

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1.2.5.1.5. Physiological assessment of the wheelchair athlete

The wheelchair/wheelchair-user system is required to perform optimally. However,

optimal performance is governed by the constraints of the athlete’s disability and the

mechanical constraints of the wheelchair. From a physiological standpoint, the

research was sought to determine whether the physiological characteristics that are

thought to govern athletic performance in able-bodied athletes apply to wheelchair

athletes across the disability range?

Physiologically orientated manual wheelchair propulsion research is divided into two

clear areas, aerobic and anaerobic performance. The anomaly that wheelchair athletes

tend to take part in most events from 100 m through to the marathon is very different

to the traditional distinction between sprint and endurance prevalent in able-bodied

athletics. In comparison to the 35 % decrease in the average velocity observed for

100 m and 5000 m World record performance for running, the decrease in the

average velocity for wheelchair racing is only 15 % (Coutts and Schutz 1988).

Hutzler (1998) explains this by stating that in the relatively small active muscle mass

of the arms, local fatigue precedes central factors as the limitation for peak

performance. Janssen et al. (1993) found that there was a strong positive relationship

between upper body isometric strength, sprint power and aerobic power in

individuals with SCI. The authors speculated that this relationship is due to the

shared dependency on active muscle mass together with peripheral muscular exercise

limiting factors. Janssen et al. (1993) postulate that measurement of one variable

might be sufficient to describe (within certain limits) the physical capacity of

individuals with spinal cord injuries. Although not fully longitudinally researched as

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yet, thoughts are that if significant relationships are found between measurements of

aerobic and anaerobic performance variables then one test could be developed to

measure the physical capacity of wheelchair athletes. This would reduce the

extensive requirement of time and laboratory instrumentation and also the

concomitant effort and cost. One limitation of this line of research is the use of

WERGs in collecting anaerobic performance data. Vanlandewijck et al. (2001) noted

that the fixed nature of WERG testing results in an increase in the force measured at

the push-rims compared with that measured during the same task performed under

field conditions. If one test is to become standard practice, a method of determining

push-rim forces that could be used to test wheelchair athletes under realistic

conditions needs to be developed first.

Conley and Krahenbuhl (1980) described running economy, the energy cost (oxygen

uptake) of working at a constant rate, as being essential to success in running. In

manual wheelchair propulsion, pushing economy is defined as the energy cost of

wheelchair propulsion at a constant speed (Lakomy and Williams 1996). Lakomy et

al. (1987) found pushing economy, defined as the oxygen cost of propulsion at 4 m.s-

1, returned a value of 0.39 when correlated with 5 km time trial time. The authors

concluded wheelchair propulsion economy did not appear to be major influence on

performance.

Jones et al. (1992) examined the relationship between pushing economy and

wheelchair propulsion technique at 2.69, 3.58, 4.69, 5.36 and 6.25 m.s-1 in male

wheelchair racers on a WERG. Ten athletes were selected from 15 and divided into

two groups, (five most and five least economical, grouped according to 2OV! ). Jones

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et al. (1992) reported that the economical group had: 1) Less head and trunk velocity

with more elbow and wrist velocity at the strike and release, 2) Released the wheel

with a straighter arm and higher wrist velocity, and 3) Stroked less frequently with

less time in contact with the rim. Jones et al. (1992) state that the economical group

had a more fluid, rhythmic motion, consistent across the speeds tested and concluded

that while an exact mechanism was not clear, a combination of these mechanical

factors may contribute to a decrease in 2OV! at a given speed.

Goosey et al. (1998b) examined the relationship between pushing economy and

selected kinematic variables at realistic racing speeds (6, 6.5 and 7 m.s-1) in eight

wheelchair racers on a MDT. Large variations in pushing economy were found

between individuals. Goosey et al. (1998b) state that at the speeds detailed above,

economy was associated with: the lighter athletes (r = 0.89, 0.86 and 0.83

respectively); a greater range of elbow movement (r = -0.85, -0.65 and –0.63

respectively) and a lower push rate (r = 0.73, 0.81 and 0.63 respectively). Goosey et

al. (1998b) concluded that the effects of lesion level and wheelchair design might be

more important in explaining differences in pushing economy than differences in

pushing technique. Goosey et al. (1998c) examined the relationship between

economy and selected kinematic variables. This study differed from Goosey et al.

(1998b) in that a 3D analysis was performed of propulsion technique on a roller

WERG at 4.70 and 6.58 m.s-1. The study found that higher ME and lower push rate

were associated with economy (p< 0.05) and concluded that the magnitude and

direction of forces may be important for determining economy of propulsion. Goosey

et al. (2000) investigated the effect of push frequency on propulsion economy at a set

speed of 6.58 m.s-1. The study was performed on a roller WERG using eight male

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wheelchair racers (T4-T8 and SB). Push frequencies of 60, 80, 120 and 140 % of the

individual’s freely chosen push frequency (FCF) were analysed. Goosey et al. (2000)

concluded that push frequency does have an effect on economy with the athlete’s

FCF being the most economical.

The studies detailed above share few uniform characteristics making strict

comparisons difficult. According to a number of researchers (Glaser et al., 1977;

Arabi et al., 1997; Bhambhani et al., 1991, 1994), the use of WERGs in the study of

the physiology of manual wheelchair propulsion is reliable and valid. However, the

validity of the wide variety of WERGs in the assessment of the kinematics of racing

wheelchair propulsion is not so clear (Tropp et al., 1997, Vanlandewijck et al.,

2001). The studies of Jones et al. (1992) and Goosey et al. (1998c, 2000) have

analysed racing wheelchair propulsion on WERGs and attempted to establish the

relationship between the mechanics of racing wheelchair propulsion and economy.

Although these studies provide a firm basis from which research into racing

wheelchair propulsion economy can be continued, the link between racing

wheelchair propulsion kinematics and economy must be studied under realistic

conditions.

The ability of muscles to produce maximal force over a short period of time is

typically referred to as anaerobic power. Originally developed for cycle and arm

crank ergometry, the updated Wingate Anaerobic Test (WAnT) described by Bar Or

et al. (1987) has been modified and adopted as the standard anaerobic power test for

wheelchair athletes (Hutzler 1998). For a more in depth review of the literature

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relating to the anaerobic fitness testing of wheelchair users, the reader is directed to

Hutzler (1998).

The WAnT protocol described by Bar Or et al. (1987) for wheelchair users facilitates

the measurement of peak power and mean power. Peak power (P5) is the highest

average power of any five-second period during the test. Mean power (P30) is the

average power produced during the test. P5 and P30 refer to the anaerobic maximal

and endurance capacity of the muscles used. In addition the index of fatigue (IOF)

can be calculated. The IOF provides a measure of the power drop off during the test.

Lees and Arthur (1988) conducted three experiments with seven British male athletes

(ISMGF classes 2-5). The first experiment investigated the stability of peak power,

mean power and maximum velocity measurements. Performing three tests over a

five-week period with resistive loads of 1.2 and 1.0 kg. No significant differences

were found between the measurements. The second experiment investigated changes

in peak and mean power output with varying resistive load. Both peak and mean

power showed a linear increase as resistive loads increased from 1.4 to 2.4 kg. In the

third experiment the relationship between peak power, mean power and sprint

performance time over 100, 200 and 400 m were examined. Significant negative

correlations (p<0.01) were found between peak power, mean power, and all

performance times. The authors concluded that the WERG produced reliable results,

that there was no clear optimum load for peak or mean power output and that peak

and mean power output was closely related to performance times.

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Woude et al. (1997, 1998) studied the anaerobic capacity of 48 elite wheelchair track

athletes (38 male, 10 female), classified into four different function classes. The

studies report class related P30 of 23, 68, 100 and 138 W for the male athletes and 38,

77 and 76 W for the female athletes (upper three classes). Sprint power for the mixed

sex cerebral palsy group was 35 W compared to the 121 W for the mixed sex

amputee group. A significant correlation (r = -0.79) was found between P30 and 200

m sprint performance times. No correlation was found between P5 or P30 and

marathon times. Specifically Woude et al. (1998) reported that sprint power relative

to body weight varied between 0.36 ± 0.03 and 1.85 ± 0.43 W.kg-1.BM-1 for the

different subject groups. The authors concluded that propulsion technique and

performance parameters are highly variable among wheelchair athletes.

Hutzler (1998) highlighted three main issues relating to the anaerobic fitness testing

literature. Firstly, the type of wheelchair ergometer used may have a considerable

effect on the results Secondly, a number of protocols based on the WAnT appear to

be in use to measure the same variables. Thirdly, There appears to be no agreed

resistance level, optimal or otherwise, for this type of testing.

Invariably anaerobic testing is conducted on a computer interfaced WERG. This

allows peak and mean power to be calculated using simple computer software.

Hutzler (1998) states that the type of ergometer used “reduces the applicability of

comparative interpretations”, (p. 105). This particular limitation relates to the

problems with comparing data from WERGs that provide a uniform wheelchair-user

interface (Niesing et al., 1988, 1990, Vosse et al., 1990) with those on which

individual wheelchairs can be mounted (Shimada et al., 1995, Goosey et al., 1998a).

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The type of WERG used in each of the studies reviewed above is detailed in table 2

As stated previously, the former usually affords more sophisticated measurements

due to the independent mounting and therefore ease of instrumentation of the wheels

and seat. The latter addresses important issues relating to the wheelchair-user

interface by allowing the wheelchair users own wheelchair to be used during the

testing.

Studies investigating the anaerobic performance of wheelchair users have utilised a

number of protocols modified from the WAnT protocol described by Bar Or et al.

(1987). A test duration of 30 s appears to be common to most studies. However,

Woude et al. (1994) performed tests of 20 seconds duration. One common variation

in the WAnT protocol is the use of and the intensity of the rolling start. The WAnT

protocol advocated by Bar-Or (1987) suggests a rolling start. Coutts and Stogryn

(1987) allowed subjects to perform a rolling start at 75 % of max speed. Lees and

Arthur (1988) used a set start speed of 60 % of the maximum flywheel velocity.

Janssen et al. (1993) used a rolling start at near maximal effort. Dallmeijer et al.

(1994) provided no quantification for the rolling start. Woude et al. (1997, 1998)

performed testing with no rolling start.

Lees and Arthur (1988) states that there appeared to be no clear optimum resistive

load. Studies conducted by Dallmeijer et al. (1994) and Janssen et al. (1994) selected

resistance loads of 0.25, 0.5 or 0.75 N.kg-1.BM-1 in order to restrict wheelchair

velocity to below 3 m.s-1 to avoid coordination problems at high velocities. Similarly

Woude et al. (1997, 1998) used resistance loads set at 2.5, 5, 7.5 or 10 % of the

combined subject and fictional wheelchair (20 kg) mass to restrict their subjects to a

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maximum velocity of 3 m.s-1. In Coutts and Stogryn (1987) tests were repeated using

a higher resistance (undisclosed) if the subject exceeded a maximum of 100 RPM.

Conversely, Hutzler (1995) reported the use of a maximal velocity protocol, which

used minimal resistance in order to achieve velocities representative of those during

actual track and basketball performance. To the best of the author’s knowledge the

optimal resistance for wheelchair users performing the WAnT is still not known.

Hutzler (1995, 1998) recommended the standardisation of braking load in roller

ergometry. It is the recommendation of the author that standardisation of the method

of reporting of resistive loads as a percentage of the subject or

wheelchair/wheelchair-user system mass should also be considered. Standardisation

of anaerobic testing of wheelchair users in terms of the protocols and resistive loads

used and the reporting of data would produce meaningful results and aid

comparisons between studies.

1.2.5.1.6. Biomechanical assessment of the wheelchair athlete

As stated previously, to the sport scientist the wheelchair/wheelchair-user system

poses a similar problem to that of any athlete whose interaction with a specific piece

of equipment brings about a sporting performance. From a biomechanical standpoint,

the interaction of the athlete and the wheelchair, brought together as a single

wheelchair/wheelchair-user system, poses an interesting performance question. How

can the wheelchair athlete bring about optimal performance given the relatively small

forces that can be produced by the muscle mass of the upper extremity?

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Invariably, when collecting kinematic data during racing wheelchair propulsion the

investigator is concerned with propulsion technique for descriptive analysis (Goosey

et al. 1997, Higgs 1983, Ridgeway et al. 1988) or in relation to an intervention such

as manipulation of the wheelchair/wheelchair-user interface (Walsh et al., 1986,

Gayle et al., 1990a, b, Woude et al., 1988b). The use of 2D analysis is limited in that

the particular subject of the analysis needs to be perpendicular to the optical axis of

the camera and be of sufficient size to facilitate accurate digitising of anatomical

landmarks or other points of interest when analysing the film. In this respect the most

reliable information that can be obtained from 2D film analysis relate to the timing

parameters of the propulsive cycle. Roeleveld et al. (1994) states that 2D analysis

was suitable for stroke, timing and displacements of segments in the sagittal plane.

For this reason studies employing 2D film analysis, with a single camera, have either

only been able to analyse one propulsive cycle during over-ground racing wheelchair

propulsion at specific points in an event (Goosey et al., 1997, Higgs 1983, Ridgeway

et al., 1988), or have had to compromise in order to analyse multiple propulsive

cycles using simulated racing wheelchair propulsion on stationary devices such as

WERGs (Cooper 1990, Gehlsen 1990, Goosey et al., 1998a, 2000) or MDTs

(Goosey et al., 1998b).

Three dimensional film analysis using the direct linear transformation (DLT) method

(Abdel-Aziz and Karara, 1971) is one of the most popular techniques for

reconstructing the location of objects in 3D space. 3D film analysis allows

investigation of the true complexity of racing wheelchair propulsion by making it

possible to map the locations of segments allowing accurate calculation of shoulder

and elbow angles during the propulsive cycle (Goosey et al., 1998c, O’Connor et al.,

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1998). However, an optimal 3D analysis, using gen-locked cameras, is restrictive.

Control points (points with known locations) must be distributed within the activity

space. In addition, the cameras need to be fixed. This usually precludes analysis of

over-ground propulsion. Veeger et al. (1991a) pioneered a method of performing a

3D analysis using a single camera and a mirror. This approach has since been used in

other studies (Helm et al., 1996, Mâsse et al., 1992, Wang et al., 1995). This method

precludes analysis of over-ground propulsion for the same reason.

Pan and tilt videography allows cameras to follow the motion of an object by rotating

about the horizontal and/or the vertical axes (pan and tilt respectively). This allows a

large subject image to be maintained at all times. These systems use specially

machined tripod heads each containing two optical encoders. The optical encoders

are aligned to sense the angular positions of the cameras. One encoder is aligned

vertically to measure pan positions, while the other horizontally, to measure tilt. This

method is called the integrated rotating camera (IRC) method. Systems using the IRC

method allow 3D film analysis to be performed over a large area by allowing

cameras to pan and tilt to follow the subject of the analysis within a pre-calibrated

space. These systems currently provide the best method of conducting film analysis

during over-ground manual wheelchair propulsion. However, to the best of the

author’s knowledge these systems have not yet been used in the study over-ground

racing wheelchair propulsion.

Typically the instant the hand contacts and releases the push-rim is identified from

the images recorded during the kinematic analysis. However, this can be difficult

even with the most sophisticated motion analysis systems. Bednarczyk and

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Sanderson (1994) and Wang et al. (1996) describe instruments designed to

accurately identify these stroke parameters. The devices used in both studies utilised

a switch mechanisms in the wheelchair user’s glove. Bednarczyk and Sanderson

(1994) wired the switch mechanism via a comparator to a LED placed in view of the

cameras. The resolution of the device was therefore determined by the 60 Hz sample

frequency of the cameras. Wang et al. (1996) independently wired switches from the

thumb, index and middle fingers to LED’s and sampled separately at 200 Hz using a

microcomputer. In this way Wang et al. (1996) were not only able to identify the

instants of contact and release, determining the durations of the propulsive and non-

propulsive phases, but were also able to identify the order in which the fingers

contacted the push-rim.

Nadeau et al. (1990) used a pressure sensitive pad and motion detectors in an

investigation of the mechanical power output of world-class wheelchair athletes.

Motion detectors, positioned at 4 and 5 m of every 5 m portion of a 30 m section of

running track, were activated when the wheelchair moved away from a pressure

sensitive pad over which it was positioned at the start line. As the wheelchair passed

in front of each motion detector a time was recorded. Split times were sent by

telemetry to a central receiver.

Coutts (1991) describes an instrument with the ability to detect and quantify

wheelchair motion. The device used a magnetic switch fixed to the wheelchair. The

switch was activated using two magnets, 180° apart, attached to the spokes of one

rear wheel. Coutts (1992, 1994) use the same instrument. Coutts (1992) uses one

wheelchair instrumented with the speed sensing system in order to describe the

72

dynamics of wheelchair basketball. In Coutts (1994) the device was transferred

between the wheelchairs of individual athletes in order to investigate the drag and

sprint performance of wheelchair basketball players. Although the device used by

Coutts provides a method of measuring wheelchair velocity during over-ground

athletic wheelchair propulsion, the resolution (two samples per revolution) is

insufficient to accurately determine intra-push wheelchair velocity. Vanlandewijck et

al. (2001) state that during the propulsive cycle the hand can be in contact with the

push-rim for anywhere between 71.0 and 121.7 ° of rear wheel rotation. Clearly

sampling every 180 ° of rear wheel rotation is insufficient to provide accurate data

regarding intra-push changes in wheelchair velocity. It is very important to be able to

accurately measure both the velocity of steady state wheelchair propulsion and the

changes in wheelchair velocity that occur due to the kinematics of propulsion.

Accurate measurement of these variables is fundamental to the assessment of

wheelchair athletes.

The ability to measure push-rim forces directly is important as it provides

information about how the force developed by the individual is directed. This

information can be used to describe and therefore improve stroke biomechanics

(Goosey-Tolfrey et al., 2001) (by maximising the force tangential to the push-rim),

to reduce injuries (by correcting for damaging stroke biomechanics), and to improve

wheelchair design. A number of investigators have attempted to tackle the problem

of how to collect kinetic data during racing wheelchair propulsion. The most popular

methods appear to be the use of instrumented, force-measuring, or SMARTwheels

(Cooper and Cheda 1989, Strauss et al., 1989, Watanabe et al., 1991, Asato et al.,

1993, Sickle et al., 1995, Stojak 1997, Wu et al., 1998) or complex WERGs with the

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ability to measure propulsion torque at the wheel hub (Niesing et al., 1988, 1990,

Ruggles et al., 1994). Other methods involve the use of static simulations of

wheelchair propulsion. Typically these devices are in the form of WERGs with

wheels that are blocked using force-measuring devices (Janssen et al., 1993, Arabi et

al., 1999, Brauer and Hertig 1981), instrumented, restrained platforms (Brubaker et

al., 1981) or force platforms (Tupling et al., 1986). Much of the credit for the

development of force-measuring wheels during the 1990’s should go to Cooper who

appears to have lead the drive for more in depth investigation of wheelchair

locomotion. Cooper authored and co-authored a number of papers in which the need

for instrumentation was emphasised (Cooper and Cheda, 1989, Cooper 1990a, b, c,

d, Vosse et al., 1990). Cooper et al. (1997), put forward a standardised method for

determining forces and moments.

Cooper and Cheda (1989) describe a wheel specifically designed for the

measurement of racing wheelchair propulsion forces/torques using beams

instrumented with strain gauges. The force/torque applied to the push-rim causes a

deflection of the beams, which is measured via strain gauges. The method outlined

has the ability to accommodate the individual athlete’s push-rims and racing

wheelchair. This early device is restrictive because it is wired directly into a

microcomputer. Variations of this device have been developed by Strauss et al.

(1989), Watanabe et al. (1991), Asato et al. (1993), Sickle et al. (1995), Stojak

(1997) and Wu et al. (1998).

Table 2 indicates the studies that have used these devices on a WERG during

simulated manual wheelchair propulsion. Very often the ideology of testing athletes

74

is infringed upon in these studies. A standard wheelchair equipped with the force-

measuring wheel is typically used in these studies. Force-measuring wheels are

heavier than normal wheels. It is also very difficult for investigators using and

developing these devices to allow for the wide variation of push-rim sizes and tube

diameters. Fundamental factors such as the distance of the push-rim from the surface

of the wheel and the variation in materials with which the push-rims are covered

cannot be completely replicated for each athlete tested. Experienced wheelchair

athletes know how important these factors are in bringing about optimal racing

wheelchair propulsion. Even those wheels that are able to closely replicate those of

the athletes (Cooper and Cheda 1989) suffer from one major limitation. To the best

of the author’s knowledge, a device still does not exist that can be used to measure

push-rim forces/torques during over-ground racing wheelchair propulsion. Therefore,

it can be concluded that we still do not have an accurate idea of the forces and

torques generated during over-ground racing wheelchair propulsion.

The use of electromyography (EMG) in manual wheelchair propulsion literature is

well documented Tables 2 and 3 detail the number of studies that have used EMG

during simulated manual wheelchair propulsion on WERGs and MDTs respectively.

Similar to the use of force-measuring wheels, measurement of the electrical activity

of the muscles used during manual wheelchair propulsion typically requires

connection to a microcomputer to allow the large volume of data that is generated to

be collected. For this reason studies using EMG simulate manual wheelchair

propulsion using WERGs or MDTs. However, unlike the collection of kinetic data,

the measurement instrument is applied to the subject and not the wheelchair. This

means that subjects are able to use their own wheelchairs. Unfortunately, out of all of

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the studies detailed in tables 2 and 3 that have used EMG, only Chow et al. (2000,

2001) has collected EMG data from athletes in their own wheelchairs.

Chow et al. (2000) investigates the effect of resistance load on the biomechanical

characteristics of racing wheelchair propulsion. Chow et al. (2001) provides a useful

comparison between the conventional and para-backhand pushing techniques. Both

studies performed 3D kinematic analyses and EMG analyses of eight muscles of the

right hand side of the body. Unfortunately, the data collected by Chow et al. (2000,

2001) were collected during simulated racing wheelchair propulsion on a WERG.

Although these studies provide the best description of the electrical activity and

activation pattern of the muscles during racing wheelchair propulsion, it can be

concluded that we still do not have data from EMG studies collected during over-

ground racing wheelchair propulsion.

1.2.6. Summary

This literature review has attempted to provide an overview of manual wheelchair

propulsion research with particular importance placed on research relating to racing

wheelchair propulsion under realistic conditions. The importance of the ergonomic

approach to the study of manual wheelchair propulsion by attempting to optimise the

wheelchair-user interface has been shown. The effect of the growth and maturity of

wheelchair sport in influencing the design of racing wheelchairs has also been

shown. The increased demands of wheelchair sport with respect to “daily use”

manual wheelchair propulsion provide a further argument for optimising the

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wheelchair-user interface with respect to the wheelchair athlete. The study of manual

wheelchair propulsion provides its own problems. Indeed as the quote from

Vanlandewijck et al. (2001) state, “Any uncontrolled deflection from reality will

influence movement/performance and consequently the relationships between the

parameters under study”, (p. 341). Many of the instruments used to either simulate or

gather data during manual wheelchair propulsion have been shown to deflect from

reality and therefore, the results of these studies should be viewed with caution. This

review has identified major methodological limitations in the literature relating to the

study of manual wheelchair propulsion, more specifically in relation to the study of

racing wheelchair propulsion. At present instrumentation for accurate analysis of

racing wheelchair propulsion under realistic conditions does not exist.

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2. Chapter 2

2.1. A telemetry-based velocometer to measure wheelchair velocity.

In order to address the concerns summarised in the previous chapter, a design brief

was formulated with a view to designing and manufacturing a velocometer. The

device would improve present testing methods by allowing data to be collected

during over-ground propulsion. Specifically the device could be used to address the

question of how wheelchair athletes accelerate their wheelchairs. The design brief

was as follows:

1. The device should provide a valid means of measuring racing wheelchair velocity

over the range of velocities typically experienced in wheelchair athletics.

2. The device should be reliable in order that repeated measurements of racing

wheelchair velocity can be made.

3. The device should be able to sample at frequencies that are sufficient to allow

intra-push changes in wheelchair velocity to be measured.

4. The device should cause the least possible disturbance to the natural function of

the wheelchair/wheelchair-user system.

78

5. The design should be such that it facilitates quick and simple transfer between

wheelchairs without alteration to the design of the wheelchair and without

damage to the wheelchair.

6. The device should be able to store or transmit real time data to a remote receiver

in a form acceptable for analysis.

According to the requirements detailed above a “velocometer” was designed and

constructed. A detailed description follows:

2.1.1. Design of the device

An optical encoder (HEDS-5700A00, RS Components Ltd., Northants, UK) driven

by a wheel with a plastic centre and sponge tyre, mounted directly onto the encoder

shaft. The wheel has been modified by fixing a 0.05 m x 0.005 m rubber O-ring

around the outer edge of the tyre about the midline to ensure a relatively small

contact area with the wheelchair wheel. The encoder is mounted so that the wheel is

in contact with the inside of one of the rear wheels of the racing wheelchair (figure

2). Each revolution of the encoder shaft results in the production of 500 pulses. The

pulses are then transmitted using a FM transmitter (418 MHz, RS Components Ltd.,

Northants, UK) with 0.18 m antenna and received using a FM receiver (418 MHz RS

Components Ltd., Northants, UK) with modified antenna for increased gain, giving a

range of approximately 200 m. The transmitter and receiver are housed inside

commercially available plastic boxes for protection. Once received the pulses are

79

converted to an analogue voltage proportional to the speed of the wheelchair rear

wheel using a frequency to voltage converter and passed through a 20 Hz low-pass

filter to remove any residual ripple voltage. The pulses are then converted from

analogue to digital data using a Picoscope (12 bit ADC 42, Pico Technology Ltd.,

Cambridgeshire, UK) analogue to digital converter interfaced via the serial port to a

laptop PC (figure 3).

Figure 2 Optical encoder and transmitter assembly

80

Figure 3 Telemetry System Block Diagram

418 MHz FM

Transmitter

Optical Encoder

418 MHz FM

Receiver

Frequency to Voltage Converter

Two Pole Low-Pass

Filter

ADC 42 Picoscope

ADC

Interface to Standard Notebook

PC

81

2.1.2. Sampling

A maximum of 24000 samples can be recorded at a frequency of up to 1000 Hz. This

frequency can be reduced to allow a longer sampling period.

2.1.3. Mounting

Two base plates were manufactured to allow the velocometer to be attached to a

wide variety of wheelchair designs. The plates were manufactured from Aluminium

alloy (BSEN 754-5 606) and can be attached to the wheelchair using Nylon clamping

blocks. To allow the velocometer to be used with spoke wheels, a Perspex disc

(approximate diameter 0.170 m) was manufactured. The disc can be mounted on the

inside surface of the wheelchair wheel. The mass of the device ranges between 0.345

kg and 0.845 kg depending on the mounting method.

2.1.4. Calibration

The velocometer calibration is performed using a specially designed interface. The

interface counts the number of pulses produced during one complete revolution of

the wheelchair rear wheel. Two lever micro-switches, sharing a common lever are

used to initiate and terminate the count. An actuating bar pressed onto the wall of the

tyre is used to bias the lever. When biased clockwise the system is in the “wait” state

with the right hand switch contact open. Rotating the wheel anticlockwise releases

82

the lever and closes the switch. This enables pulse counting. After one revolution the

actuating bar engages the lever and closes the left hand contact. This terminates

counting. This process is repeated three times. The mean value is then used in the

equation shown below, to calculate the calibration factor. Velocity is calculated by

multiplying the calibration factor by the velocometer output voltage.

Figure 4 Calibration equation

Nomenclature

C Calibration factor

d Diameter of the wheelchair wheel with which the velocometer is in

contact

CON Simplified constants

n Number of pulses counted during one complete revolution of the

wheelchair wheel with which the velocometer is in contact

nd

CONC =

83

2.2. Study 1: validity and system linearity

2.2.1. Introduction

In order to determine the validity of any measurement device (i.e. how accurately it

measures whatever it is designed to measure), the measurements made by the device

should be compared against an established “gold standard”. The validity of the

velocometer should therefore be determined by comparison against another method

of measuring velocity believed to be the best. Video analysis is widely agreed to be

the current “gold standard” for the assessment of motion in sport. Therefore, the aim

of this experiment was to test whether the velocometer could provide valid linear

measurements of racing wheelchair velocity, at velocities typically found in

wheelchair racing, compared with 2D video analysis.

Wheelchair athletes race in events from 100 m to the marathon. The 100 m gold

medal winner in the T54 racing classification at the 2000 Paralympic games recorded

a time of 14.46. This is an average velocity of approximately 6.92 m.s-1. Obviously

maximum velocity will be somewhat higher than this. Wheelchair athletes have been

observed at the Manchester Metropolitan University, Department of Exercise and

Sport Science, at treadmill velocities of 9 m.s-1. The author believes these speeds to

be a valid representation of current wheelchair performance.

Similar to able-bodied athletics, wheelchair racing events begin from a stationary

start. In wheelchair racing the shortest sprint event, the 100 m, involves wheelchair

athletes accelerating from a sprint start to top speed, which is then maintained for the

84

short duration of the rest of the race. As stated previously the author has observed

wheelchair athletes at propulsion velocities of up to 9 m.s-1. Therefore, the velocity

range 1 to 9 m.s-1 was chosen as a valid range over which the linearity of the

velocometer measurements should be assessed.

2.2.2. Method

The velocometer was fitted to a racing wheelchair (Mistral, Draft wheelchairs,

Godmanchester, UK) and then calibrated using the method explained previously. The

racing wheelchair was then fixed in a stationary position on a motor driven treadmill

(MDT) (Woodway, model ELG2) (figure 5). A 60 kg mass was placed in the

wheelchair to simulate the mass of a wheelchair athlete. The MDT was chosen in

order to provide repeatable steady state speed trials. Three speed trials were then

conducted at each of three speeds (1, 5 and 9 m.s-1). During each trial voltage data

from the velocometer and 2D video film data were recorded simultaneously at 200

Hz. A single high-speed camera (HSC-200 PS, Peak Performance Technologies Inc.,

Englewood, CO) mounted on a tripod and interfaced with a high-speed SVHS video

cassette recorder (AG-5700, Panasonic, Matsushita Electrical Industrial Co., Ltd.,

Japan) was used to record the motion of the rear wheel with which the velocometer

was in contact. Velocity was obtained from the video film by manually digitising

three points on the wheelchair using motion analysis software (Peak Motus, version

6.0). The raw co-ordinate data were smoothed using a quintic spline routine

(Woltring, 1986). The data were exported to a spreadsheet (Microsoft Excel 2000)

for further analysis. Average measured velocity was plotted against trial velocity.

85

Figure 5 Experimental set-up for studies 1, 2, 3 and 4 showing treadmill wheelchair

mounting system

2.5 m

Video camera &

tripod

86

2.2.3. Results

The relationships 0052.00076.1 −= χγ and 0063.00142.1 −= χγ were found for

the measurements of velocity from the velocometer and 2D video film data

respectively (figure 6). For both methods r2 = 1. This indicates that the measurements

made using the velocometer and video analysis are both linear. The average root

mean square deviation (ARMSD) was used to compare velocity calculated from the

velocometer with velocity calculated from manual digitising. The ARMSD

calculated for each test speed from three trials was 0.06 ± 0.002, 0.27 ± 0.05 and 0.48

± 0.16 m.s-1 at 1, 5 and 9 m.s-1 respectively. These data indicate a mean variation of

6, 5 and 5 % between the velocometer and 2D video film data at 1, 5 and 9 m.s-1

respectively.

87

Figure 6 Velocometer validity and system linearity.

y = 1.0142x - 0.0063R2 = 1.

y = 1.0076x - 0.0052R2 = 1.

0.00

1.00

2.00

3.00

4.00

5.00

6.00

7.00

8.00

9.00

10.00

0.00 1.00 2.00 3.00 4.00 5.00 6.00 7.00 8.00 9.00 10.00

Trial Velocity (m.s-1).

Cal

cula

ted

Vel

ocity

(m.s-1

).

Velocometer

Peak

Linear (Peak)

Linear (Velocometer)

88

2.2.4. Discussion

The aim of this study was to test whether the velocometer could provide valid linear

measurements of racing wheelchair velocity at velocities typically found in

wheelchair racing. The treadmill (figure 5) provided a method of spinning the

wheelchair wheels at the required velocity. The treadmill wheelchair mounting

system (TWMS) was used to ensure that the wheelchair would remain stable during

the experiment. These results indicate that measurements of racing wheelchair

velocity made using the velocometer are valid and linear. Similar to a bicycle

dynamo, the velocometer runs in contact with the wheelchair wheel. When the

wheelchair travels at a given velocity, the wheelchair wheel completes a certain

number of revolutions per unit of time. The velocometer wheel also completes a

certain number of revolutions per unit of time.

Figure 7 illustrates the dimensions of the wheelchair and velocometer wheels and the

path drawn by the velocometer wheel. The number of revolutions of the velocometer

wheel can be calculated by dividing the circumference of the velocometer wheel by

the circumference of the circle drawn by the path of the velocometer wheel. Thus, if

a wheelchair travels at 1 m.s-1 for 60 seconds, the rear wheels perform 28.30

revolutions. In this time the velocometer wheel performs 56.57 revolutions. Any

abnormality between the wheelchair wheel and the velocometer wheel will affect the

data approximately 56 times. A wheelchair travelling at 10 m.s-1 for 60 second, the

velocometer wheel performs 565.67 revolutions. Consequently the data is affected

approximately 565 times.

89

The optical encoder shaft, turned by the velocometer wheel, produces 500 electrical

pulses per revolution. Similar to the interface between the wheelchair and

velocometer wheels, this system is also linear. This linearity throughout the system

explains the systematic variation of approximately 5 % between the velocometer and

2D video film data across the three test velocities.

The cause of the 5 % variation between the velocometer and the 2D video film data

can be explained by a number of factors. The velocometer was calibrated before the

60 kg mass, simulating the mass of the wheelchair athlete, was added to the

wheelchair. Deformation of the wheelchair frame and wheels as a direct result of

adding the simulated athlete mass produces a small change in the relationship

between the velocometer wheel and the wheelchair wheel, measured previously

during calibration. The 200 Hz data sampling rate and use of the average root mean

square deviation (ARMSD) may result in exaggerated estimations of the

measurement errors in studies 1 and 2. Thus, the worst case scenario is represented.

The reasons for this are twofold. Firstly, The 200 Hz sample rate is higher than

strictly necessary for the steady state and relatively small accelerations used in

studies 1 and 2, and therefore generates a larger number of matched data points.

Secondly, the ARMSD compares the difference between these data points, which

may include errors in both the velocometer and video film data.

90

Figure 7 Wheelchair and velocometer wheel dimensions.

ro = 0.027 m

Velocometer wheel

rw = 0.338 m

rs = 0.054 m

Hub centre

91

2.3. Study 2: dynamic response

2.3.1. Introduction

Wheelchair racing is a dynamic activity. As stated previously the 100 m involves

rapid changes in velocity as athletes accelerate from the sprint start to top speed. To

the author’s knowledge the inter-push rate of change in velocity during wheelchair

athletic events has not been studied. It is also widely accepted that propulsion

kinematics bring about intra-push changes in the wheelchair/wheelchair-user system.

However, the magnitudes of these changes in velocity are also unknown. Any device

designed to measure racing wheelchair velocity should be able to measure these

changes. Therefore, the aim of this experiment was to assess the ability of the

velocometer to accurately measure changes in racing wheelchair velocity.

As in study 1, in order to quantify the accuracy of the measurements made using the

velocometer a direct comparison was required between those measurements and

measurements made by the current “gold standard”, video analysis. In order to

ensure repeatability the treadmill detailed in study 1 was used to accelerate the

wheelchair rear wheels in a controlled manner. Due to the absence of any reference

values regarding the magnitude of changes in velocity during over-ground racing

wheelchair propulsion the treadmill was operated at maximum acceleration.

92

2.3.2. Method

Five acceleration and deceleration trials were performed using the wheelchair/MDT

set-up explained previously (figure 5). During each trial the treadmill belt was

accelerated from standstill to 9.5 m.s-1 and then decelerated from 9.5 m.s-1 to

standstill. Velocometer voltage data and 2D video film data were recorded

simultaneously at 200 Hz. The video analysis equipment is detailed in study 1. The

raw co-ordinate data were smoothed using a Butterworth digital filter (4 Hz cut-off).

The data were exported to a spreadsheet (Microsoft Excel 2000) for further analysis.

The first five seconds of each acceleration and deceleration trial were analysed. Five

seconds was chosen as this time period was sufficient to allow the wheelchair wheel

to accelerate and decelerate between the two end velocities stated above.

2.3.3. Results

The ARMSD was used to compare the differences between velocity calculated from

the velocometer data with velocity calculated from manually digitising the 2D video

film. The ARMSD for the five acceleration and five deceleration trials was 0.29 ±

0.086 and 0.51 ± 0.115 m.s-1 respectively. These data indicate greater accuracy in

measuring acceleration than deceleration. However, these data also indicate an

acceptable level of accuracy in the measurement of wheelchair velocity under

acceleration and deceleration.

93

Figure 8 Velocometer and manually digitised, 2D video film data collected during (a)

one acceleration trial and (b) one deceleration trial.

0.00

2.00

4.00

6.00

8.00

10.00

0 200 400 600 800 1000

Data Point No.

Vel

ocity

(ms-1

).Velocometer Data Points

Manually Digitised Data Points

0.00

2.00

4.00

6.00

8.00

10.00

0 200 400 600 800 1000

Data Point No.

Vel

ocity

(ms-1

).

Velocometer Data Points

Manually Digitised Data Points

a

b

94

2.3.4. Discussion

The aim of this study was to assess the ability of the velocometer to accurately

measure changes in racing wheelchair velocity, such as those observed in racing

wheelchair propulsion during acceleration from a sprint start. The treadmill (figure 5)

provided a method of accelerating and decelerating the wheelchair. The TWMS was

used to ensure that the wheelchair would remain stable during the experiment. The

results indicate that the velocometer is able to accurately measure rapid changes in

racing wheelchair velocity. The velocometer uses an optical encoder with a virtually

frictionless encoder assembly. This minimises any resistance that may cause the

velocometer wheel to slip on the wheelchair wheel under acceleration. The low mass

of the velocometer wheel, made of small quantities of plastic, sponge and rubber,

ensures that the momentum of the wheel is kept low and therefore minimises any

slippage under deceleration. The effect is an accurate instantaneous response to any

motion of the wheelchair. The rubber O-ring fixed around the outer edge of the

velocometer wheel, about the midline, ensures adequate friction between either the

carbon wheels or the Perspex disc used with the spoke wheels, despite the relatively

small contact area.

The accuracy of the measurements made appear to vary with respect to acceleration

and deceleration. A possible explanation for this may be that the surface marker used

to denote a single point on the wheel for digitising purposes became unclear at high

speeds. This is an inherent limitation of video analysis when recording movements at

high speed. Any variation between the five acceleration and five declaration trials in

respect to the rate of change in velocity would bring about this variation.

95

As explained previously the sample rate and statistical procedure used to analyse the

date may provide exaggerated estimations of the measurement error.

96

2.4. Study 3: reliability

2.4.1. Introduction

A device can be said to be reliable if it is able to repeatedly perform a function to a

satisfactory or given standard of accuracy. A meter rule is said to be a reliable

measurement system because it will always provide the same answer when

measuring the length of the same object. In the case of the velocometer the

measurement quantity is velocity. The ability of the velocometer to repeatedly

measure racing wheelchair velocity needs to be assessed. Therefore, the aim of this

experiment was to assess the ability of the velocometer to perform repeated accurate

measurements of racing wheelchair velocity.

The velocometer will be required to fulfil the role of a research tool. In this role the

velocometer will be required to measure racing wheelchair velocity during repeated

trials on one or more racing wheelchairs. Specifically, the velocometer needs to

provide reliable measurements when transferred between racing wheelchairs. In

order to simulate the transfer between wheelchairs the velocometer was repeatedly

removed and re-attached between reliability trials.

Similar to studies one and two the accuracy of the measurements made using the

velocometer needs to be quantified by direct comparison with those made by the

current “gold standard”. In this case the “gold standard” was provided by the velocity

of the wheelchair rear wheel spun by a 24 V, 100 W DC, DC servo motor (unloaded

97

speed 2400 RPM). The motor was able to spin the wheelchair wheel at a calculated

constant velocity of 4.59 m.s-1.

The use of carbon fibre disc wheels is common among elite wheelchair athletes. The

wheelchair used in the previous two studies was fitted with these disc wheels.

However, these wheels are expensive and not available to all athletes. Although less

aerodynamic, spoke wheels are more accessible to entry-level athletes. Figure 2

shows the velocometer fitted to a racing wheelchair with spoke wheels using the

Perspex disc. The use of the Perspex disc increases the mass of the wheel on which it

is fitted and changes the interface between the velocometer wheel and running

surface. For this reason data were collected using both disc and spoke wheels to

assess the variation in reliability between the two conditions.

2.4.2. Method

The wheelchair/MDT set-up explained previously (figure 5) was used. The

wheelchair was elevated approximately 4 cm above the treadmill belt to allow the

left hand rear wheel, the wheel the velocometer was in contact with, to spin freely.

The motor detailed previously was used to spin the wheelchair wheel at a calculated

constant velocity of 4.59 m.s-1.Ten trials were performed using the same disc wheel

used in the previous two experiments. A further ten trials were performed using a

spoke wheel fitted with the Perspex disc. During each trial, velocometer data were

collected at 200 Hz for twenty seconds. Between each trial the velocometer was

98

removed, re-attached and calibrated. The data were exported to a spreadsheet

(Microsoft Excel 2000) for further analysis.

2.4.3. Results

A measurement of the agreement between the mean trial velocity obtained using the

velocometer and the constant wheel velocity calculated from the motor speed is

shown graphically in figure 9, a and b., The figure indicates the reliability of the

velocometer measurements within 95 % confidence limits (Bland and Altman, 1986).

Expressed as a percentage of the mean trial velocity, the mean ± SD of the

differences were 0.00 ± 0.17 %, for the disc wheel trials and 0.00 ± 0.41 %, for the

spoke wheel trials.

99

Figure 9 Agreement between the constant velocity of a wheel spinning in air and

mean velocity calculated from the velocometer data, within a five percent error band,

from (a) Ten disc wheel trials (b) Ten spoke wheel trials.

4.35

4.45

4.55

4.65

4.75

4.85

1 2 3 4 5 6 7 8Trial.

Vel

ocity

(m.s-1

).

4.35

4.45

4.55

4.65

4.75

4.85

1 2 3 4 5 6 7 8Trial.

Vel

ocity

(m.s-1

).

a

b

100

2.4.4. Discussion

The aim of this study was to assess the ability of the velocometer to perform repeated

accurate measurements of racing wheelchair velocity. The results of this study

suggest that accurate repeated measurements of wheelchair velocity can be made

using the velocometer. The results also suggest that the reliability of the

measurements is greater than the 5 % error widely acknowledged as being

acceptable.

The TWMS (figure 5) was used to ensure that the wheelchair would remain stable

and elevated to allow the rear wheel to spin freely during the experiment. The motor

detailed previously was used to ensure the wheelchair wheel would spin at a

calculated constant velocity of 4.59 m.s-1. Greater variation in the spread of the data

points can be seen clearly in figure 9 b, (spoke wheels) with respect to a, (disc

wheels). A possible explanation may be the use of the Perspex disc and variation in

its repeated fixing to the spoke wheel. The Perspex disc was employed to provide a

smooth running surface for the velocometer wheel. Any variation in the placement

and fixing of the disc to the spoke wheel would result in small variations in the data

collected. Due to the arrangement of the spokes, radiating from a wide central hub to

a narrow rim, even slight misalignment of the Perspex disc would have had the effect

of altering the diameter of the velocometer wheel by compressing and releasing it

with each revolution of the wheelchair wheel. This would cause the velocometer

wheel to spin slightly faster when compressed and slightly slower when released,

resulting in small over and under estimations of velocity when the wheelchair wheel

101

is travelling at a uniform velocity. This did not occur when using carbon fibre disc

wheels due to the smooth uniform inner surface.

102

2.5. Studies 4 and 5: resistance

2.5.1. Introduction

Vanlandewijck et al. (2001) states that human movement should be studied under

realistic conditions. In making this point Vanlandewijck is arguing against the

simulation of manual wheelchair propulsion in research studies. This point of view

relates specifically to the influence of artificial factors on the movement/performance

and consequently the relationship between the parameters under study. The effect of

wheelchair resistance on racing wheelchair propulsion has not been studied.

However, it is logical to assume that propelling a wheelchair with increased

resistance will require a larger propulsive force and have an increased physiological

demand. Therefore, it is very important that any device seeking to measure natural

manual wheelchair propulsion, which must be attached to a wheelchair/wheelchair-

user system, imposes minimal resistance. The aim of this study was to quantify the

resistance the velocometer imposes on the wheelchair/wheelchair-user system.

To quantify the resistance imposed on the wheelchair by the velocometer, two

experiments were performed. Study four was performed to quantify velocometer

resistance in relation to the natural deceleration of a wheelchair wheel spinning in

air. Study five was performed to quantify velocometer resistance in relation to the

natural deceleration of a wheelchair/wheelchair-user system. In these two

experiments velocometer data were not collected. 2D video film was recorded during

both experiments in order to analyse the rundown trials. For the reasons stated in

103

study three data were collected using both disc and spoke wheels to assess the

variation in resistance between the two conditions.

2.5.2. Method

Study four: The experimental set-up is shown in figure 5 As explained in study 3, the

wheelchair was elevated approximately 4 cm above the treadmill belt to allow the

left hand rear wheel to spin freely. The DC servomotor detailed in study 3 was used

to spin the wheelchair wheel at a constant velocity before being removed allowing

the wheel to decelerate naturally. Ten rundown trials were performed under each of

four conditions. The four conditions were as follows: spoke wheel/velocometer;

spoke wheel no velocometer; disc wheel/velocometer and disc wheel no

velocometer. 2D film of the trials was recorded at 200 Hz using the video analysis

equipment detailed in study 1. Wheel velocity was obtained by manually digitising

one point on the wheel for each run down trial.

Study five: The experimental set-up is shown in figure 10. One male athlete (age =

24 years; mass = 30.1 kg; wheelchair mass = 9.6 kg;) participated in this study. The

subject was an experienced wheelchair racer (racing classification = T54). The

subject performed ten over-ground rundown trials under each of four conditions.

Each trial involved accelerating between two markers 5 m apart, and then adopting a

position with hands on the steering mechanism until the wheelchair had passed a

further marker, 15 m away. The four conditions were the same as in experiment one.

2D film of the trials was recorded using at 50 Hz using a single SVHS video camera

104

(AG-DP200E, Panasonic, Matsushita Electrical Industrial Co., Japan) fixed to a

stationary tripod. Wheelchair velocity was obtained by manually digitising the centre

of the front wheel hub for each run down trial.

The raw co-ordinate data from both experiments were smoothed using a quintic

spline routine. The data were then exported to a spreadsheet (Microsoft Excel 2000)

for further analysis. A 5th order Polynomial trendline was drawn though all of the

trials. For each trial the trendline equation was differentiated to give acceleration.

Force was then calculated from the resolved differentiated equations.

105

Figure 10 Study 5 experimental set-up showing camera and calibration pole placement in relation to the line of progression.

12 m

Line of Progression

Video camera & tripod

5 m 15 m

106

2.5.3. Results

The results of the rundown trials are shown in table 5 Velocometer resistance is

shown as the difference between the rundown forces calculated for disc and spoke

wheel trials conducted with and without the velocometer. Velocometer resistance

calculated as a percentage increase in the deceleration force of the wheelchair rear

wheel, calculated from data collected during rundown trials with the wheel spinning

in air, was 78 and 61 % for the disc and spoke wheels respectively. Velocometer

resistance calculated as a percentage increase in the deceleration force of the

wheelchair/wheelchair-user system, calculated from data collected during over-

ground rundown trials, was 26 and 28 % for the disc and spoke wheels respectively.

107

Table 5 Velocometer resistance calculated from rundown trials.

Wheel trials Wheelchair/wheelchair-user system

trials Mean Force Variance Mean Force Variance (m.s-2) (N) (N) (m.s-2) (N) (N)

Disc wheel/velocometer -0.21 ± 0.13 -0.64 -0.50 -0.13 ± 0.03 -5.22 -1.38

Disc wheel no velocometer

-0.05 ± 0.17 -0.14 -0.10 ± 0.01 -3.84

Spoke

wheel/velocometer -0.47 ± 0.13 -1.49 -0.91 -0.16 ± 0.06 -6.53 -1.82

Spoke wheel no

velocometer -0.19 ± 0.17 -0.58 -0.12 ± 0.03 -4.71

108

2.5.4. Discussion

The aim of this study was to quantify the resistance the velocometer imposes on the

wheelchair/wheelchair-user system. Two experiments were performed to allow the

resistance of the velocometer to be quantified in both a controlled and specific

setting. The results of these two studies suggest that the velocometer caused a

relatively large increase in the mechanical resistance of the wheelchair wheel

spinning freely in air but a relatively small increase in comparison to the total

mechanical resistance of the racing wheelchair used in the study.

Experiment one: The TWMS (figure 5) was used to ensure that the wheelchair would

remain stable and elevated to allow the rear wheel to spin freely during the

experiment. The DC servomotor used explained previously was used to spin the

wheel. The motor was used to ensure the wheel spun at a uniform and sufficiently

high velocity to produce a long rundown time. The relatively small mechanical

resistance of the wheel bearings and air drag ensured that the velocometer resistance

would be sufficiently large to be measured.

Experiment two: Over-ground rundown trials were performed by an experienced

wheelchair athlete in his own racing wheelchair on an indoor running track surface to

create an environment specific to wheelchair racing. A single SVHS video camera

filming at 50 Hz was felt to be sufficient to allow the deceleration of the wheelchair

to be calculated from the video film.

109

The effects of the velocometer on athletic performance have not been assessed. The

technological advancements in wheelchair design and manufacture have meant that

the frictional resistance of racing wheelchairs is very small. Therefore, any device

coming into contact with the revolving wheels of the wheelchair will increase the

force required to propel the wheelchair. However, the approximate increase of 1.3

and 1.8 N in resistance (disc and spoke wheels respectively) represents a minimal

load. Over 100 m this load represents an additional 130-180 N.m of work.

110

3. Chapter 3

3.1. An explanation of the intra-push velocity profile of over-ground racing

wheelchair propulsion during the first six pushes of the sprint start.

3.1.1. Introduction

Chapter two details five studies performed to test the validity and system linearity,

dynamic response, reliability and resistance of the velocometer. The studies showed

that the velocometer provides a method of measuring wheelchair velocity that is

valid, reliable and able to detect rapid changes with minimal interference to the

wheelchair/wheelchair-user system. The studies detailed in chapter two were

performed in a laboratory environment using static simulations of wheelchair

locomotion. Clearly the utility of the velocometer in assessing over-ground

propulsion needed to be demonstrated.

In wheelchair sprint events, like able-bodied sprint events, the sprint start is

considered crucial to success (Nadeau et al., 1990). However, very little information

exists about the sprint start in wheelchair racing. During the sprint start rapid

increases in velocity are observed initially as the athlete accelerates. In able-bodied

athletics, the kinematic characteristics of the sprint start are different to those

observed in steady state running (Delecluse 1997). It is logical to assume that this is

also true in wheelchair racing. Therefore, it is not possible to draw comparisons with

studies investigating the kinematics of steady state racing wheelchair propulsion.

However, it is also logical to assume that the kinematics of racing wheelchair

111

propulsion during the sprint start may evolve gradually into the kinematics of steady

state racing wheelchair propulsion. The movement patterns during the sprint start in

wheelchair racing need to be established.

Due to the descriptive nature of the study a specific research hypothesis was not

formulated. The aim of this study was to analyse the kinematics of over-ground

racing wheelchair propulsion during the first six pushes of a sprint start. Temporal

information regarding the phases of the propulsive cycle, and the intra-push

occurrence of peak velocity and acceleration are included. The importance of the

trunk in terms of its contribution to the forward momentum of the

wheelchair/wheelchair-user system was analysed.

3.1.2. Method

A Pan and Tilt system (Peak Performance Technologies Inc., Colorado, USA) was

used to record the kinematics of wheelchair propulsion. By using this system, it was

possible to record the motion of the wheelchair/wheelchair-user system while

keeping a sufficiently large image size to allow accurate manual digitising of

anatomical landmarks. Two high-speed (200 Hz) cameras (Peak HSC-200) were

fixed onto two Tribach equipped surveying tripods with tripod heads incorporating

pan and tilt encoders. Each camera was interfaced with a high-speed VCR

(Panasonic AG-5700) and high-speed monitor (Viewmagic MD – 935A). The master

camera was interfaced with the event and video control unit through which the

112

master and slave cameras were gen-locked. Both cameras were interfaced with a

directional encoding unit (DEU). The experimental set-up is shown in figure 11

113

Figure 11 Study 6 experimental set-up showing the pan and tilt camera and

calibration pole placement in relation to the line of progression.

10 m

15 m

1 2 3

Master Camera Slave Camera

Line of Progression

114

3.1.2.1. Calibration

Three calibration poles of 0.031 m diameter and 3.220 m in length were positioned

perpendicular to the line of progression. Exact perpendicular alignment with the line

of progression was achieved using a specifically designed dual spirit level. Manual

digitisation of the poles for calibration purposes (figure 12) was performed within the

motion analysis software (Peak Motus, version 6.1). A comparison between the

actual and calculated pole vales is shown in table 6 below.

Table 6 Actual and calculated pan and tilt calibration values.

Actual

Calculated

Interpoint Distance

(m)

Interpoint Distance

(m)

Pole 1

3.220

3.218

Pole 2

3.220

3.218

Pole 3

3.220

3.222

115

Figure 12 Calibration procedure. Point denoted by cross is digitised as follows: 1)

Top point at bottom of view, 2) Top point at top of view, 3) Bottom point at bottom

of view, 4) Bottom point at top of view.

1) 2)

3) 4)

Field of

view

+

++

++

+

Calibration

pole

116

3.1.2.2. Pilot study

The accuracy of the Pan and Tilt system has not been assessed with regard to

analysis of the racing wheelchair propulsion in independent studies. Therefore, there

was a need to validate its use. In order to assess the standard of the calibration and

the subsequent accuracy of the reconstructed coordinate data, an object of known

length was filmed. Figure 13 details the upper extremity calibration frame fabricated

for use in the pilot study from four pieces of brass rod and white spheres

(approximate diameter 0.03 m). The frame has three points of reference, providing

two known segmental distances and an angle. In this respect the frame provided an

approximation of the upper extremity. The frame was fastened to a black board

which provided a contrast to the white spheres The board was orientated so that the

marker spheres were in the approximate positions of the shoulder, elbow and wrist

joint centres of a wheelchair athlete immediately prior to a sprint start. A volunteer

then ran with the board along a previously calibrated 10 m section of indoor running

track while being filmed using the PAT system. While running the volunteer moved

the board in a circular motion simulating the kinematics of the upper extremity

during wheelchair propulsion. Both cameras simultaneously tracked the subject and

calibration frame. The whole movement was filmed at 200 Hz. every second image

(100 Hz) of the calibration frame was then captured on a PC and the three white

spheres digitised. The mean and standard deviation of the simulated segmental

distances and the angle between the two segments was calculated.

The mean segmental distance between markers 1 and 2, and 2 and 3 was 0.504 ±

0.005 and 0.500 ± 0.004 m respectively. The angle between the two segments was

117

88.6 ± 0.4 °. After fabrication the calibration frame was measured using a standard

engineer’s rule and long arm goniometer. A comparison between the measured

values and the values calculated from the digitised data indicates a possible error of

0.02 m and 1.5 °. These values are generally considered within acceptable limits for

kinematic analysis. Compared to an elbow range of motion of 120 ° the error is

approximately 1.25 %. These results show that the PAT system is a valid

measurement tool with an acceptable degree of accuracy for this study.

118

Figure 13 Upper extremity calibration frame.

0.5 m

0.5 m

119

3.1.2.3. Data collection

One male athlete (age = 28 years; body mass = 60.6 kg; wheelchair mass = 8.5 kg;

condition Spina Bifida Hydrocephalus) gave informed consent to participate in this

study. The subject was an experienced international wheelchair racer (racing

classification = T4 specialising in 100, 200 and 400 m sprint events). The subject

adopted a seated position in his own racing wheelchair (Draft, Godmanchester,

England) fitted with Carbon fibre disc wheels (Corima, Loriol, France). The wheels

were set into the frame at a camber angle of 10 ° and were fitted with 0.36 m

diameter, push-rims.

During a familiarisation period the subject performed ten practice trials to become

familiar with the experimental protocol and the environment to help the subject

perform naturally. Starting commands used in athletics were employed to initiate

each of the ten maximal sprint trials. Each trial was performed in a seated position,

from a stationary start along a 10 m portion of an indoor running track. Between

three and five minutes were allowed between trials to minimise the effect of fatigue.

During each trial data from the velocometer and video film were recorded

simultaneously at 200 Hz. A shutter speed of 1/1000 s was used. Nine surface

markers were used in the digitising process, the model (figure 14) included 1 = top of

head, 2 = 7th cervical vertebrae, 3 = shoulder joint centre, 4 = base of trunk, 5 =

elbow joint centre, 6 = wrist joint centre, 7 = 4th metacarpalphalangeal joint. (markers

6 and 7 positioned on the athletes glove), 8 = hub of rear wheel, 9 = wheel rim.

Angles were defined as follows: shoulder angle (fragmented angle between the

humerus segment and the hub marker translated along the x-axis), elbow angle

120

(intersection of the humerus and ulna segments), trunk angle (fragmented angle

between the trunk segment and the hub marker translated along the x-axis) and

contact and release angle (intersection of segment 8 - 9 and the hub marker translated

along the x-axis, projected onto the x-y plane). Marker motion was described using a

Cartesian coordinate system where the x direction described anterior-posterior

motion, the y direction described superior-inferior motion, and the z direction

described medial-lateral motion. The point of intersection of the x-axis (traced

through the hub centre), and y-axis (positioned on the leading edge of the rear wheel)

= 0°. 90° = TDC and -90° = BDC.

121

Figure 14 Wheelchair/wheelchair-user system model used in the manual digitising of

the 3D video film

1

2 3

6

4

5

7

8

9

x-anterior-posterior

z-medial-lateral motion y-superior-inferior

TDC

BDC

0°

122

3.1.2.4. Data analysis

Two trials were excluded on the grounds that the video files were insufficiently clear

to allow accurate digitising. Video film data from the remaining eight trials were

manually digitised at 100 Hz using motion analysis software (Peak Motus, version

6.1). The raw co-ordinate data were smoothed using a quintic spline routine

(Woltring, 1986). The resulting data were exported to a spreadsheet (Microsoft Excel

2000).

Velocometer data were exported to a spreadsheet (Microsoft Excel 2000). A

numerical sort reduced the data to a frequency of 100 Hz. This met the frequency

requirements suggested for wheelchair propulsion (DiGiovine et al., 2000). Intra-

push velocities were calculated from the voltage data using the previously

determined calibration value.

The momentum of the head and trunk was assessed according to Lees and Barton

(1996). Linear velocity data calculated from displacement of the HAT and

wheelchair/wheelchair-user system along the x-axis was used with centre of mass

information from Dempster (1955) to calculate the relative, transfer and total

horizontal momentum.

123

3.1.2.5. Digitising error

The reliability of the digitising process (the association of x, y and z co-ordinates to

specific markers captured on film) was estimated by determining the coefficient of

repeatability (Bland and Altman 1986) for the elbow angle from one trial that was re-

digitised. The coefficient of repeatability was determined as ± 1 º.

3.1.3. Results

Table 7 shows the absolute and percentage mean and standard deviation data for the

propulsion and recovery phases, and the duration of each propulsive cycle. The

duration of each propulsive cycle decreases from 0.82 ± 0.02 to 0.45 ± 0.01 s.

Similarly the mean duration of the propulsion phase also decreases from 0.62 ± 0.02

to 0.21 ± 0.01 s. In contrast, the mean duration of the recovery phase increased from

0.20 ± 0.01 to 0.24 ± 0.02 s. With each subsequent propulsive cycle the subject had

less time to apply a propulsive force to the wheelchair and took more time to recover.

Table 7 also shows the push-rim angle at contact and release, and the range over

which propulsive force was applied. 0 ° = point of intersection of the x-axis, traced

through the hub centre, and y-axis on the leading edge of the rear wheel. The data

show that force was applied over a greater range over each push. This was due to

contact occurring closer to TDC and release occurring closer to BDC with each push.

124

Table 7 Mean propulsive cycle data for the first six pushes of the sprint start calculated from eight trials.

Push Propulsion Phase Recovery Phase Push Durations Contact Angle Release Angle Range

(s) (s) (s) (0) (0) (0)

P1 0.62 ± 0.02 0.20 ± 0.01 0.82 ± 0.02 65 ± 5 -5 ± 14 70 ± 11

% Cycle time 76 ± 1 24 ± 1

P2 0.33 ± 0.01 0.19 ± 0.01 0.52 v 0.01 72 ± 12 -27 ± 9 99 ± 19

% Cycle time 63 ± 1 37 ± 1

P3 0.28 ± 0.01 0.21 ± 0.01 0.49 ± 0.01 75 ± 13 -42 ± 5 117 ± 15

% Cycle time 57 ± 1 43 ± 1

P4 0.24 ± 0.01 0.23 ± 0.01 0.47 ± 0.02 81 ± 8 -43 ± 15 124 ± 15

% Cycle time 52 ± 2 48 ± 2

P5 0.21 ± 0.01 0.24 ± 0.01 0.45 ± 0.01 80 ± 11 -46 ± 29 125 ± 27

% Cycle time 47 ± 2 53 ± 2

P6 0.21 ± 0.01 0.24 ± 0.02 0.45 ± 0.01 83 ± 12 -50 ± 21 133 ± 27

% Cycle time 47 ± 3 53 ± 3

125

Table 8 shows mean velocity data at key events during each push calculated from the

velocity profiles of the eight trials. Velocity at contact, release and peak velocity

increase with each push. In pushes two to six velocity at contact is lower than that at

the point of release of the previous push. This is due to deceleration of the

wheelchair during the period of the recovery phase. The reader should note that the

time of peak velocity indicates that peak velocity occurs after release.

Table 9 shows mean peak acceleration, time of peak acceleration and push-rim angle

at peak acceleration data for each push. The data indicate that peak acceleration

increased from 6.61 ± 2.69 to 19.14 ± 2.56 m.s-2 at the fifth push. Peak acceleration

for push six was 17.15 ± 5.52 m.s-2. In contrast to peak velocity, peak acceleration

occurred during the propulsive phase and progressively later around the rim at a

push-rim angle of 27.08 ± 15.16 to –24.59 ± 6.03 ° at the fifth push. Peak

acceleration occurred at a push-rim angle of -24.29 ± 6.25 ° for push 6.

126

Table 8 Mean velocity data for the first six pushes of the sprint start calculated from

eight trials.

Push

Velocity at contact

Velocity at release

Peak velocity

Time of peak velocity

(relative to contact)

(m.s-1) (m.s-1) (m.s-1) (s)

P1 0 1.50 ± 0.05 1.56 ± 0.06 0.65 ± 0.04

P2 1.18 ± 0.03 2.38 ± 0.05 2.48 ± 0.11 0.37 ± 0.01

P3 1.97 ± 0.02 2.99 ± 0.06 3.09 ± 0.05 0.32 ± 0.00

P4 2.60 ± 0.05 3.46 ± 0.03 3.61 ± 0.06 0.31 ± 0.02

P5 3.13 ± 0.06 3.88 ± 0.06 4.05 ± 0.05 0.28 ± 0.01

P6 3.53 ± 0.07 4.21 ± 0.05 4.38 ± 0.05 0.28 ± 0.01

127

Table 9 Mean acceleration data for the first six pushes of the sprint start calculated

from eight trials.

Push

Peak acceleration

Time of peak acceleration

(relative to contact) Push-rim angle at peak

acceleration (m.s-2) (s) (°)

P1 6.61 ± 2.69 0.47 ± 0.04 27.08 ± 15.16

P2 15.84 ± 1.79 0.20 ± 0.00 1.78 ± 9.27

P3 16.71 ± 3.12 0.17 ± 0.00 -13.14 ± 9.28

P4 18.15 ± 3.15 0.15 ± 0.01 -19.03 ± 7.23

P5 19.14 ± 2.56 0.14 ± 0.00 -24.59 ± 6.03

P6 17.15 ± 5.52 0.13 ± 0.00 -24.29 ± 6.25

128

Figure 15 shows a velocity profile and angular displacements of the elbow, shoulder

and trunk from one of the eight trials. Figure 16 shows the same velocity profile with

elbow, shoulder and trunk angular velocities. Angular data for the shoulder and trunk

are presented in relation to the x-axis. The reader should note the amplitude of the

elbow angular displacement and angular velocity curves. The curves indicate that the

elbow angle at peak acceleration increased from 92.13 ± 8.64 to 111.44 ± 5.75 ° at

the fifth push. The elbow angle at peak acceleration for push six was 110.27 ± 7.81 °.

Angular velocity of the elbow increased with each push from 183.72 ± 102.64 to

497.55 ± 119.21 ° s-1.

129

Figure 15 Intra-push wheelchair velocity and trunk, shoulder and elbow angular

displacement during the first six pushes of the sprint start

a

b

0.00

1.00

2.00

3.00

4.00

5.00

2.34 2.54 2.74 2.94 3.14Time (s).

Vel

ocit

y (m

.s-1

).

0

20

40

60

80

100

120

140

160

180

Ang

le (

0 ).

Velocometer Velocity Data Release ContactElbow Angle Shoulder to X Trunk to X

0.00

1.00

2.00

3.00

4.00

5.00

0.00 0.20 0.40 0.60 0.80 1.00 1.20Time (s).

Vel

ocit

y (m

.s-1).

0

20

40

60

80

100

120

140

160

180

Ang

le (

0 ).


130

Figure 16 Intra-push wheelchair velocity and trunk, shoulder and elbow angular

velocity during the first six pushes of the sprint start

a

b

0.00

1.00

2.00

3.00

4.00

5.00

2.34 2.54 2.74 2.94 3.14Time (s).

Vel

ocit

y (m

.s-1

).

-800

-600

-400

-200

0

200

400

600

800

1000

Ang

ular

Vel

ocit

y (0 .s-1

).


0.00

1.00

2.00

3.00

4.00

5.00

0.00 0.20 0.40 0.60 0.80 1.00 1.20Time (s).

Vel

ocit

y (m

.s-1).

-800

-600

-400

-200

0

200

400

600

800

1000

Ang

ular

Vel

ocit

y (0 .s-1

).

Velocometer Velocity Data Release Contact

Elbow Angle Shoulder to X Trunk to X

131

3.1.3.1. Coefficient of variation

In order to provide a measurement of the variability in velocity between each push

across all trials and between trials themselves, the coefficient of variation was

calculated. The coefficient of variation calculated across all trials from the mean and

standard deviation of the increase in velocity during each push was 3.26, 3.93, 6.14,

5.27, 5.44 and 3.70 % respectively. The coefficient of variation calculated across all

trials from the mean and standard deviation of the velocities at the point of release of

push six was 1.25 %. These values indicate only small variations exist between the

velocity reached at the point of release during push six and the change in velocity

brought about by each push. From a performance point of view this indicates the

ability of this athlete to reproduce the sprint start.

3.1.3.2. Relative momentum analysis

The relationship between relative, transfer and total momentum of the head and trunk

is shown in Figure 17. The velocity profile has been included to aid explanation of

this data. The net relative momentum of the head and trunk is shown as zero.

However, the relative momentum curve indicates positive and negative influences on

the wheelchair/wheelchair-user system. The negative areas indicate backwards head

and trunk motion relative to the wheelchair. This increases the velocity of the

wheelchair/wheelchair-user system. Positive areas indicate forward head and trunk

motion and a retarding effect. This is illustrated in push five by the occurrence of

peak velocity at the point after release when the curve shows the head and trunk

passing the zero line from negative to positive relative momentum.

132

Figure 17 The relationship between relative, transfer and total momentum of the head

and trunk during the first six pushes of the sprint start

-2.00

-1.00

0.00

1.00

2.00

3.00

4.00

5.00

2.34 2.54 2.74 2.94 3.14

Time (s).

Vel

ocit

y (m

.s-1

).

-100

0

100

200

300

400

Mom

entu

m (

N.s

).

Velocometer Velocity Data P5 Peak VelocityP5 Peak Acceleration ReleaseContact Transfer MomentumRelative Momentum Total Horizontal MomentumRelative Momentum Trendline

-2.00

-1.00

0.00

1.00

2.00

3.00

4.00

5.00

0.00 0.20 0.40 0.60 0.80 1.00 1.20

Time (s).

Vel

ocit

y (m

.s-1

).

-100

0

100

200

300

400

Mom

entu

m (

N.s

).

Velocometer Velocity Data ReleaseContact Transfer MomentumRelative Momentum Total Horizontal MomentumRelative Momentum Trendline

a

b

133

3.1.4. Discussion

The findings of the present study provide support for the view that racing wheelchair

propulsion under realistic conditions is made up of more that two phases

(Vanlandewijck et al., 2001). This has implications for manual wheelchair

propulsion related studies using traditional definitions of a “propulsive” or “push”

and “non-propulsive” or “recovery” phases. The kinematics of racing wheelchair

propulsion differ considerably from those observed during manual wheelchair

propulsion in a daily use, “active” or basketball wheelchair. For this reason the data

presented here will be discussed in relation to data from studies employing over-

ground propulsion using wheelchair racers in their own racing wheelchairs.

The results show that from the first push the mean duration of the propulsive cycle

decreased in a concave curvilinear fashion over the six pushes from 0.82 ± 0.02 to

0.45 ± 0.01 s. Similarly the mean duration of the propulsive phase also decreased in a

concave curvilinear fashion from 0.62 ± 0.02 to 0.21 ± 0.01 s. However, the mean

duration of the recovery phase increased with each push from 0.20 ± 0.01 to 0.24 ±

0.02 s. In order to increase wheelchair velocity a larger force must be applied to the

push-rim in a shorter time than during the previous push. The results show that this is

achieved by increasing the range over which the hand is in contact with the push-rim.

The subject contacts the rim closer to TDC and releases the rim closer to BDC with

each push.

The angular displacement and angular velocity data (figures 15 and 16) show that as

wheeling velocity increased angular displacement at the elbow also increased. The

134

arm appears to be slightly more extended with each subsequent push. The angular

velocity of the elbow increased with each push until push four. Angular velocity

remained consistent during pushes five and six. The angular displacement and

angular velocity of the trunk and shoulder remained relatively consistent across

pushes one, two, three, four, five and six.

The first push differs from subsequent pushes due to the stationary start. To initiate

movement of the wheelchair/wheelchair-user system the inertia of that system must

first be overcome. For the wheelchair athlete, this is the task of the relatively small

muscle mass of the arms and chest aided by the momentum of the trunk. The first

push is unique in terms of the motion of the trunk during the propulsive phase.

Between the initiation of the first push to approximately 0.5 s the trunk angle remains

relatively consistent. At approximately 0.5 s the trunk begins to flex. The elbow

appears to flex slightly at the initiation of the push and then extends rapidly reaching

peak extension at the point of release before returning to a flexed position. The point

at which the trunk begins to flex and the elbow extends coincide allowing the hands

to follow further round the push-rim and adding to the propulsive effect of the arms.

During the propulsive phase the shoulder flexes, from an initially hyper extended

position, slowly at first then rapidly just before the point of release before returning

to a hyper extended position.

The second push is the first push in which the subject has to contact the push-rim

while the wheelchair wheel is in motion. Figure 15 shows the trunk to be

approaching peak flexion at contact. The trunk extends throughout the propulsive

phase. This is opposite to that observed in push one and appears to be counter

135

productive as simultaneously the elbow extends. This can be seen in pushes two,

three, four, five and six, although the magnitude of the effect reduces with each

subsequent push due to increased elbow and decreased trunk extension. For this and

subsequent pushes the shoulder follows a similar pattern, flexing from the point of

contact to the point of release and then and hyper extending before contact. In pushes

three and four the range of motion of the trunk is reduced. This is largely due to the

amplitude of the peak at contact, which indicates that the trunk is less flexed at this

point while peak extension remains relatively consistent.

As stated previously, in pushes five and six, the trunk range of motion continues to

decrease. The shoulder follows the pattern explained above. The elbow range of

motion continues to increase compared to subsequent pushes. Pushes five and six

show signs of steady state propulsion. The durations of the propulsive cycle, the

propulsive phase, the recovery phase (table 7) and the time of peak velocity relative

to contact (table 8) all appear to plateau. Peak acceleration, time of peak acceleration

and the push-rim angle at peak acceleration are all decreased with respect to the

previous push (table 9).

Higgs (1986) reported a mean propulsive cycle time of 0.62 ± 0.14 s for sprint racers

with the mean duration of the propulsion phase being 35 ± 7 % of the mean cycle

time. However, Higgs (1986) does not take into account the time during which the

hand is in contact with the rim but not applying a propulsive force. According to

Higgs (1986) this period of time is included in the recovery phase. Consequently,

values reported as a percentage of cycle time are lower for the propulsion phase.

Goosey et al. (1997) found a mean cycle time of 0.53 ± 0.02 s. Propulsion time was

136

28 ± 5 % for senior male T3 and T4 wheelchair racers. Ridgway et al. (1988) found

that in classification IV/V the mean cycle time was 0.59 s. The percentage of time

spent in propulsion and recovery was similar for all classes, with approximately 33

% spent in propulsion and 67 % in recovery.

Higgs (1986) found that sprint athletes contacted the push-rim at a mean angle of 31°

before and released the push-rim at a mean angle of 100 ° after TDC. This gives a

range of approximately 131 °. In this present study contact and release angles are

measured from the point of intersection of the x (traced through the hub centre) and

y-axis, on the leading edge of the rear wheel a difference of 90 °. Although there is

some agreement between the results in terms of the angular displacement of the

wheel during the propulsion phase, the results differ markedly in the angle at contact

and release. All of the subjects in the Higgs study contacted the push-rim before

TDC. In the present study contact was not observed until after TDC. Similarly in the

present study release occurred much closer to bottom dead centre (BDC) than in the

Higgs study.

The differences between Higgs (1986), Ridgway et al. (1988) and the present study

can be attributed largely to differences in the wheelchair/wheelchair-user interface,

specifically seat height, the diameter of the push-rims, and the stage of the race

during which the data were collected. However, allowing for these differences, there

are signs that the movement patterns described in the pushes five and six of the

present study may migrate into the patterns observed by other investigators studying

steady state over-ground racing wheelchair propulsion.

137

It appears that the limiting factor to racing wheelchair propulsion velocity could be

the force velocity relationship of the muscles involved in propulsion. It has been

shown that in order to increase wheeling velocity the athlete must apply an

increasingly large propulsive force to the push-rim over a decreasing period time. It

has also been shown that this is achieved by applying force over an increasingly

large range of the push-rim. However, the range over which the force can be applied

to the push-rim is not infinite. Therefore, the speed at which force can be developed

within the muscle acts as a ceiling effect to increasing propulsion velocity. The

acceleration data (table 9) support this. With each push, peak acceleration occurs at

point when the hand is further round the push-rim than in the previous push. The

increasing force requirement requires an increasingly long period of time to generate.

The time of peak velocity relative to contact is shown in table 8. Comparing these

times with the mean durations of the propulsive phase it can be seen that within each

push peak velocity occurs after release. The point at which peak velocity occurs may

be attributed to the motion of the head, arms and trunk (HAT). Specifically, the

change from trunk flexion to extension and the phasing with the arms.

Vanlandewijck et al. (1994) highlighted the importance of movements of the arms

and trunk during manual wheelchair propulsion stating that ME studies should focus

not only on the propulsive phase but also on the movement pattern during recovery.

Specifically relating to racing wheelchair propulsion Vanlandewijck and Chappel

(1996) described trunk range of motion as a key parameter in identifying the

functional potential of wheelchair athletes.

138

Wheelchair propulsion is a quasi-conservative system unlike the counter movement

jump and running stride described by Lees and Barton (1996). Losses resulting from

the mechanical interaction of wheel bearings with their housings and tyres with the

floor surface are minimal in modern racing wheelchairs. Independent of propulsion,

trunk flexion results in rearward movement of the wheelchair and extension of the

trunk results in forward movement of the wheelchair. Depending on the velocity of

the movement and the changes in the location of the mass centre, the net change is

approximately zero (figure 17). During the push phase trunk flexion serves to place

the arms and hands in advantageous positions to apply force to the push-rims. The

force applied to the push-rims masks the small negative effect of trunk flexion.

During the recovery phase extension of the trunk forces the wheelchair forward. It is

at this point peak velocity is reached.

By using the velocometer in this study, it was possible to gather important

information concerning the intra-push changes in racing wheelchair velocity during

the first six pushes of the sprint start. The unique telemetry-based research tool with

the ability to sample at high frequencies allowed high-resolution data to be gathered

under realistic conditions. From this data, it was possible to easily construct the

velocity and acceleration profiles of the sprint starts. The instances of hand contact

and release were clearly visible from the velocity profile, allowing simple

identification of the phases of the propulsive cycle and facilitating analysis of the

associated timing parameters. The velocity and acceleration profiles also provided

important information regarding the temporal instances of intra-push peak velocity

and acceleration, which confirmed the importance of the contribution of the HAT

during the sprint start. Previously the information presented here would have been

139

obtained using film analysis. In most cases, the use of film analysis precludes the use

of realistic propulsion and typically racing wheelchair propulsion is simulated with

the use of a WERG, which produces an unrealistic propulsion environment. In this

respect, the velocometer represents an important methodological development in the

study of racing wheelchair propulsion.

3.1.5. Conclusion

The aim of this study was to analyse the first six pushes of a sprint start during over-

ground racing wheelchair propulsion. The results presented describe the temporal

characteristics of over-ground racing wheelchair propulsion during first six pushes of

the sprint start. Signs are present that the movement patterns described may migrate

into the patterns observed by other investigators studying steady state over-ground


The occurrences of peak velocity in what is commonly referred to as the “non-

propulsive” or “recovery” phase clearly demonstrates the importance of the HAT in

racing wheelchair propulsion. The results of this study support the view that in order

to gain meaningful data, racing wheelchair propulsion should be simulated under

realistic conditions as described by Vanlandewijck et al. (2001). The telemetry-based

velocometer used in this study is an important methodological development in the

study of racing wheelchair propulsion, making data under realistic conditions

possible.

140

Future research into the characteristics of racing wheelchair propulsion in an effort to

improve performance should look to manipulate the temporal phasing of segments.

This would aid in the investigation of the intra-push velocity and acceleration

profiles wheelchair racers. Further investigation of the relative contribution of

segments to racing wheelchair propulsion is warranted. Understanding and

optimising the contribution of the HAT during the non-propulsive phase may be a

key element in improving wheelchair racing performance.

141

4. Chapter 4

4.1. General Discussion

From the original development as a tool for rehabilitation, shortly after World War II

by Sir Ludwig Guttman and colleagues, wheelchair sports have grown into

internationally recognised athletic events. British wheelchair athletes have enjoyed

Paralympic success in wheelchair sprinting, the most high profile of these events,

and can be clearly identified as being at the forefront of international disability sport.

Studies focusing on racing wheelchair propulsion have suffered from a number of

limitations. The use of WERGs threatens the ecological validity of the results.

Typically, during simulated manual wheelchair propulsion on a WERG, the

wheelchair is fixed in place. Under these conditions any movement of the head, arms

or trunk (HAT) that would normally contribute to the forward momentum of the

wheelchair during over-ground wheelchair ambulation is not considered. Treadmill

testing is preferred because all of the energy from the subject contributes to the

motion of the wheelchair. However, the use of MDTs is unsuitable for measuring

sprint performance and neither WERG nor treadmill testing are able to reproduce the

effects of increasing air resistance with increasing velocity. To ensure the continued

success of British wheelchair sprint athletes, equipment must be developed for the

collection of data during over-ground wheelchair sprinting.

The purpose of this thesis was to present the design, manufacture and results of

detailed investigation to test the utility of a telemetry-based velocometer with the

ability to measure intra-push changes in racing wheelchair velocity. The results of

142

these investigations suggest that the velocometer can be used to collect highly

ecologically valid measurements of velocity during over-ground racing wheelchair

propulsion. Measurements of racing wheelchair velocity were shown to be valid and

linear across a range of speeds typically encountered in racing wheelchair

propulsion. The measurements were also shown to be accurate and reliable during

rapid changes in racing wheelchair velocity. The resistance produced by the

velocometer was shown to be minimal in relation to the normal mechanical

resistance of a racing wheelchair.

In an investigation of the sprint start in wheelchair racing performed using over-

ground trials, the velocometer facilitated the measurement of the velocity profile of

one elite wheelchair athlete during the first six pushes. The results indicate that the

occurrences of peak velocity during the “non-propulsive” or “recovery” phase clearly

demonstrates the importance of the HAT in racing wheelchair propulsion, supporting

the opinion that in order to gain meaningful data, racing wheelchair propulsion

should be simulated under realistic conditions.

The information presented in this document indicates that the telemetry-based

velocometer is an important methodological development in the study of racing

wheelchair propulsion. The ability of the velocometer to sample at high frequencies

provides a method for accurately measuring intra-push changes in velocity during

over-ground racing wheelchair propulsion. During over-ground racing wheelchair

propulsion, the movements of the athlete that influence the motion of the wheelchair

are transferred through the rear wheels. Data collected from this point is the product

of all of the factors that influence propulsion. The resulting velocity profile provides

143

accurate information regarding intra-push changes in velocity, which can be

differentiated to provide an acceleration profile, and the timing parameters of the

propulsive cycle. Previously, to achieve these results investigators have had to use

2D, or more accurately 3D film analysis.

4.1.1. Limitations

Suitable camera and calibration pole positions for filming a wheelchair/wheelchair-

user system during over-ground locomotion were required for a two-camera pan and

tilt protocol. It was found that the finalised camera positions were largely governed

by the internal constraints of the building in which data collection took place. This

meant that both cameras had to film one side of the body from different views. And

therefore, symmetry was assumed.

The maximum distance that could be satisfactory filmed using the equipment and

indoor space available was approximately 10 m. This allowed the first six complete

propulsive cycles to be filmed. The digitising of an object of known length over 10 m

showed that with this particular pan and tilt set-up that the system was extremely

accurate in reconstructing three-dimensional coordinate data. The integrated rotating

camera method provided an ideal approach when investigating over-ground racing

wheelchair propulsion.

The instants of contact and release were observed from the video film. This method

has been used widely in the manual wheelchair propulsion literature. However,

accurate identification of contact and release using this method is sometimes

144

difficult. Although the instant of contact may not necessarily be the point of positive

force application and vice versa for the point of release, clearly identifiable features

in the velocity profile were consistent with the incidents of contact and release

observed from the video.

The instant the rear wheels started to rotate was used to align the velocometer and 3D

video film data. In hindsight a method of marking both sets of data, such as a flash

from an LED positioned in the field of view of the cameras would have provided a

clearer, less time consuming and arguably more accurate method of aligning the two

data sets.

Segmental mass information was taken from the cadaver study conducted by

Dempster (1955). Information relating to the segmental mass percentage of total

body mass for wheelchair athletes with Spina Bifida Hydrocephalus was not

available. It should be stated that the results presented here are probably an

underestimation of the specific contribution of the HAT. This is due to the fact that

wheelchair athletes in general have over developed upper body and under developed

lower body musculature. This being the case, the upper body segmental mass of a

wheelchair athlete will be greater than that of an able-bodied individual, expressed as

a percentage of total body mass. The opposite will be true of lower body segmental

mass. Although this would not affect the general pattern of the data or alter the

interpretation presented, this can be considered a limitation of the study.

A further limitation of the study is that velocometer calibration has to be performed

with an unloaded wheelchair to allow the rear wheel to turn with the velocometer

145

attached. It is the author’s opinion that, when loaded, mechanical stress on the seat

cage and wheels will bring about minor variation in the dimension between the

centre of the velocometer wheel and the running surface on the wheelchair wheel.

This is similar to the concerns expressed in section 2.4.4.

4.2. Conclusion

The velocometer facilitates high-resolution measurements, which allow intra-push

changes in racing wheelchair velocity to be quantified. Its unique function in being

telemetry-based, light and generating minimal resistance permits measurements that

previously have been confined to the laboratory, to be performed in the field. For

wheelchair athletes wishing to analyse performance, the velocometer can provide a

valuable measurement tool that can be used in training and competition to monitor

performance. The velocometer can be considered a valid, reliable research tool for

the ecologically valid collection of data pertaining to the changes that occur during


4.4. Future Recommendations

The role of the velocometer is to measure changes in racing wheelchair velocity. The

velocometer should continue be used to assess, under realistic conditions, how

wheelchair athletes propel their wheelchairs. The importance and contribution of the

kinematics of the HAT should be the prime focus of this research.

146

Some of the conclusions of previous studies, formulated from the results of

investigations using simulated manual wheelchair propulsion on WERGs, could be

investigated under realistic conditions using the velocometer. The development of an

over-ground sprint test using the velocometer would provide useful performance

information for wheelchair athletes specialising in sprint events.

Continued development of the velocometer, incorporated into a telemetry-based

force-measuring wheel, would allow the kinetics of racing wheelchair propulsion to

be investigated under realistic conditions. To the best of the author’s knowledge a

telemetry-based force-measuring wheel capable of being used during over-ground

propulsion does not exist.

147

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Appendices

1 A telemetry-based velocometer to measure wheelchair velocity.

2 An explanation of the intra-push velocity profile of over-ground

racing wheelchair propulsion during the first six pushes of the sprint

start.

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