Journal of Rehabilitation Research and Development Vol. 39 No. 6, November/December 2002 Pages 635–650
Shoulder kinematics and kinetics during two speeds of wheelchair propulsion Alicia M. Koontz, PhD, ATP; Rory A. Cooper, PhD; Michael L. Boninger, MD; Aaron L. Souza, MS; Brian T. Fay, PhD Department of Rehabilitation Science and Technology and Department of Bioengineering, University of Pittsburgh, PA; Department of Physical Medicine and Rehabilitation, University of Pittsburgh Medical Center Health System, Pittsburgh, PA; Human Engineering Research Laboratories, Department of Veterans Affairs Pittsburgh Healthcare System, Pittsburgh, PA Abstract— The primary objective of this study was to examine the kinematics and kinetics of the shoulder during wheelchair propulsion at a slow and moderate speed. Twenty-seven individuals with paraplegia propelled their wheelchairs at speeds of 0.9 m/s and 1.8 m/s while a motion analysis system captured movements of their upper limbs and SMARTWheels simultaneously recorded their pushrim kinetics. Intraclass R correlation and Cronbach’s coefficient alpha statistics revealed that all shoulder parameters were stable and consistent between strokes and speeds. The shoulder exhibited a greater range of motion, and forces and moments at the shoulder were 1.2 to 2.0 times greater (p < 0.05) during the 1.8 m/s speed trial. Peak posterior forces occurred near the end of the propulsion phase, and at the same time, the shoulder was maximally flexed and minimally abducted (p > 0.1). Shoulder positioning and the associated peak shoulder loads during propulsion may be important indicators for identifying manual wheelchair users at risk for developing shoulder pain and injury.
Key words: biomechanics, kinematics, kinetics, shoulder injury, shoulder pain, spinal cord injury, wheelchair propulsion.
INTRODUCTION Wheelchair propulsion has been implicated as a causative factor in developing shoulder pain and injury. An estimated 30 to 75 percent of manual wheelchair users will develop shoulder pain during their lifetime [1–6]. In one study, 72 percent of individuals with a spinal cord injury (SCI) had radiological evidence of degenerative shoulder changes and, in most cases, were asymptomatic [6]. The factors that predispose an individual to developing chronic shoulder pain and pathology are not well understood. These factors may be related to shoulder positioning and repetitive loading occurring at the joint during propulsion. Several researchers have recorded shoulder movement patterns during propulsion for various groups of wheelchair users. Many of the earlier studies presented a two-dimensional (2D) analysis of shoulder kinematics [7–9], whereas more recently, three-dimensional (3D) analyses have been performed with axial rotation as a third articulation of the humerus [10–13]. These studies have documented that during the propulsion phase of the cycle, the shoulder exhibits internal rotation, abduction, and flexion and extension. In addition, movement patterns vary depending on wheelchair type, level of injury, and speed [11–14].
This material was based on work supported by the U.S. Department of Veterans Affairs (Project B689-RA and B2674-CA), Eastern Paralyzed Veterans of America, National Institutes of Health (NIH K08 HD01122-01), and the National Institute on Disability and Rehabilitation Research Training grant (H133P970013-98). Portions of this work were presented at the Rehabilitation Engineering and Assistive Technology Society of North America (RESNA) 2000 Conference, June 28 to July 2, 2000, in Orlando, Florida, and at the American Society of Biomechanics Annual Meeting, July 19–22, 2000, in Chicago, Illinois. Address all correspondence and requests for reprints to Alicia M. Koontz, PhD; Human Engineering Research Laboratories, Department of Veterans Affairs Pittsburgh Healthcare System, 7180 Highland Drive, 151R-1, Pittsburgh, PA 15206; 412-365-4833; fax: 412365-4850; email:
[email protected]. 635
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Only a few investigators have obtained estimates of the net shoulder-joint forces and moments during wheelchair propulsion [8,15–17]. Shoulder kinetics have been determined for propulsion at various speeds, over simulated inclines, and for varying external power outputs [16,17]. Most of the studies on shoulder kinetics have been limited to 2D analyses or static data collection procedures. Limited data regarding 3D shoulder kinetics and kinematics during wheelchair propulsion are available. Peak shoulder forces and moments that occur at or near the end ranges of the shoulder motion may help explain the high prevalence of shoulder pain and pathology among wheelchair users. Therefore, the primary objective of this study was to provide a 3D description of both shoulder kinetics and range of motion during wheelchair propulsion. We hypothesized that— 1. Output parameters of a 3D dynamic shoulder model would produce stable and reliable parameters that describe shoulder biomechanics during propulsion. 2. Shoulder angles, forces, and moments would be statistically different between speeds of propulsion. 3. Maximum and minimum shoulder angles would occur near the same time during the propulsion phase as the peak shoulder forces and moments. The shoulder girdle is the primary source of power in most activities performed by wheelchair users but, consequently, is more prone to overuse injuries and chronic pain. Determining the overall shoulder stresses and positioning during propulsion may help identify harmful aspects to a manual wheelchair user’s stroke, which may be linked to the development of pain and injury. In addition, information about the applied loads to the shoulders during propulsion may be used for the optimization of wheelchair setup and performance and the user may be trained in more effective wheelchair techniques to avoid injuries in the future. METHODS This study took place at the Department of Veterans Affairs (VA) Medical Center in Pittsburgh, Pennsylvania, and was approved by the local VA Research and Development Committee, VA Human Studies Subcommittee, as well as the University of Pittsburgh’s Institutional Review Board. Study Participants The inclusion criteria were (1) complete or incomplete SCI below T1, (2) the use of a manual wheelchair
as a primary mode of mobility, (3) the use of a manual wheelchair for at least 1 year, and (4) ages between 18 and 65 years. In addition, brief medical histories were collected from each subject before performing the propulsion trials. This information was used to ensure that subjects did not have a heart condition that could possibly be exacerbated by propelling a wheelchair. Twentyseven individuals were recruited from within the VA Pittsburgh Healthcare System and from local rehabilitation hospitals and clinics. All subjects provided written informed consent before their participation in this study. The sample consisted of 10 women and 17 men. Their mean ages and years postinjury were 36 ± 10 (range = 20 to 65 years) and 11 ± 5 (range = 1 to 23 years), respectively. The average mass of the subjects was 74.6 kg ± 17.3 kg (range = 43.1 kg to 105.2 kg). Kinetic Measurement System Propulsion kinetics were obtained with a SMARTWheel, which is a 3D force and torque-sensing push rim [18]. Details concerning the system components, percent linearity, and precision of this device have been previously documented [18,19]. Kinetic data from two SMARTWheels were collected at 240 Hz and filtered with an 8th order Butterworth low-pass filter, zero lag, and a 20 Hz cutoff frequency. The kinetic data were linearly interpolated for synchronization with the kinematic data collection rate of 60 Hz. Kinematic Measurement System An OPTOTRAK 3020 3D motion analysis system (Northern Digital, Inc., Waterloo, Ontario, Canada) was used to collect position data of infrared-emitting diode markers placed on the wheelchair user’s body. The markers were attached with double-sided sticky tabs to the bony landmarks on the arm (Figure 1): most lateral part of the acromion process, lateral epicondyle, olecranon, third and fifth metacarpalphalangeal joints (not shown), and radial and ulnar styloids. This system was configured to collect the real-time movement patterns of the entire arm, head, and torso during propulsion. Two cameras were positioned, facing each other, on either side of the subjects to capture movement patterns bilaterally. A synchronization pulse from the OPTOTRAK was used to trigger the start of kinetic and kinematic data collection. Kinematic data were collected at a 60 Hz sampling frequency and filtered with a 4th order Butterworth low-pass filter, zero lag, and a 7 Hz cutoff frequency.
637 KOONTZ et al. Shoulder kinematics and kinetics during propulsion
Figure 1. Marker placement and global coordinate system axis representation.
Data Collection Subjects used their personal wheelchairs in this study, and their wheelchairs were not altered, with the exception of replacing their original rear wheels with the SMARTWheel . All subjects used a standard smooth push rim of s 0.88 cm in tube diameter and 52.7 cm overall rim diameter. Wheelchairs were centered between the two motion analysis cameras and secured to a dynamometer, with a resistance comparable to that of a smooth level tile surface [20]. Subjects were provided with an acclimation period for at least 5 minutes before data collection to become accustomed to propelling on the dynamometer. Afterward, participants were instructed to propel at two constant speeds of 0.9 m/s (2 mi/h) and 1.8 m/s (4 mi/h). They monitored their own speed using a video computer monitor that displayed the forward velocity and the differential velocity of the roller system. Subjects propelled at the target speeds for at least 1 minute before 20 s of kinetic and kinematic data were collected. Data Analysis Biomechanical Shoulder Model We determined shoulder joint angles and forces using a local coordinate system (LCS) approach described by Cooper and colleagues [21]. The analysis was limited to movements and kinetics of the glenohumeral joint because of the difficulties in measuring the positions of
the scapula, clavicle, and thoracic spine movement during propulsion. Shoulder motion was described relative to the trunk. In this analysis, the trunk was restricted to moving in the sagittal plane only. A recent finding of Cooper et al., who investigated various shoulder and trunk representations, concluded that only allowing the trunk to flex produced similar results as if the trunk were allowed to freely rotate [21]. Limb segments were assumed to be rigid with uniform density. Body weight, segment lengths, and circumferences were obtained from each wheelchair user in the study. Segment mass, segment center of mass, and inertias of the hand, forearm, and upper arm were computed according to the methods described by Hanavan and Clauser et al. and entered into the model [22,23]. The output variables of the biomechanical model were the time-varying 3D net joint muscle forces acting at the glenohumeral joint (along a anterior/posterior axis (x), superior/inferior axis (y), and medial/lateral axis (z), shoulder angles expressed in anatomical terms, and the net moments acting to flex/extend (z), abduct/adduct (x), and internally/externally rotate the arm (y′)). The “righthand rule” was used to define positive forces and moments. Sagittal flexion, horizontal flexion, abduction, and internal rotation angles were positive. All angles were determined in reference to a neutral anatomic position, that is, with the arm straight down to the side, palm facing in toward the body. All equations and postprocessing analyses were implemented in MATLAB (Mathworks, Inc., Natick, Massachusetts). Data Reduction Shoulder joint angles, forces, and moments were analyzed for five propulsion strokes across each speed trial. The stroke has generally been divided into two phases: a drive or propulsion phase and a recovery phase [24,25]. The onset of propulsion was defined as the point at which the propulsive moment at the push rim, as measured by the SMARTWheels, deviated from the baseline by 5 percent. The end of propulsion and the beginning of recovery were defined as the point at which the propulsive moment at the push rim returned to baseline and remained within 5 percent. From the motion, force, and moment curves, peak values during the propulsion phase were obtained. Each stroke was analyzed separately. The time instance at which the maximum and minimum angles, forces, and moments occurred was expressed as a percentage of the propulsion phase, since time spent in this phase varied across individuals and across strokes.
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Statistical Analysis We used a two-step approach to determine if the shoulder model produced consistent and reliable parameters across multiple strokes and speeds. First, the interstroke reliability was evaluated by intraclass R correlation coefficients (ICC). The ICCs were computed separately for each model parameter, each speed, and each side. A parameter was considered reliable if intraclass R was > 0.60 at both speeds [26]. Second, for parameters that met the preceding criterion, we computed Cronbach’s alpha coefficient to determine the increased reliability of creating aggregated scores across the repeated strokes. The summed score was considered to have good reliability if the alpha coefficient was > 0.80 at both speeds and sides [27]. For parameters that met both of these conditions, the peak forces, moments, and maximum and minimum angles were averaged across strokes for each side. We then performed Pearson product-moment correlations to assess the strength of the right and left side associations among the shoulder variables. Since the sides were correlated (p < 0.05, minimum r > 0.6), peak shoulder values were averaged across sides for each speed trial. Paired t-tests were conducted to test for significant performance differences between the two speed trials. A p value less than 0.05 was considered statistically significant. Paired t-tests were also used to test for similarities between the relative timing of the peak shoulder force and moment variables and the maximum and minimum angles (p > 0.10). Experiment-wise error rate control procedures were not used. All statistics were performed in the SPSS statistical package (SPSS, Inc., Chicago, Illinois). RESULTS Descriptive All model output parameters met the criteria for reliability and stability; therefore, all the parameters underwent further analysis. Table 1 lists propulsion
characteristics for the group. Participants had no difficulty maintaining the 0.9 m/s target speed, but some had trouble reaching and/or maintaining the fast speed of 1.8 m/s. The average number of strokes per second was greater, while total cycle time and the amount of time spent in contact with the push rim were shorter during the 1.8 m/s speed condition. The angle of hand contact relative to the horizontal (+x-axis) and the overall contact angle on the push rim during the propulsion phase are also shown in Table 1. Subjects propelled with smaller hand push rim contact angles at the fast speed (100.3° and 110.3°, respectively). Shoulder Motion During Propulsion Each figure in the paper (except Figures 2 and 3) represents a different subject. Figures containing data for five strokes demonstrate the stroke-to-stroke variability within a subject. Table 2 contains the group mean maximum and minimum shoulder angles and relative timing during the propulsion phase for both speed conditions. Also, shoulder angles for five stroke cycles at the slowest speed are presented for a single subject in Figure 4. As illustrated in Figure 4, the shoulder started in a position of extension and progressed to flexion in both the sagittal and horizontal plane views. The arms remained abducted throughout the entire cycle and reached an average peak 29 percent into the propulsion phase (Table 2). For most subjects, a second peak was observed later during the recovery phase and just before hand contact. Only three of the subjects externally rotated their shoulders during the 1.8 m/s speed trial, and therefore, only internal rotation angles were reported for the group. A general pattern of decreasing internal rotation during the propulsion phase and increasing rotation during the recovery phase was observed. Shoulders were minimally abducted and internally rotated near the end of the propulsion phase. For both speeds tested, the overall ranges of shoulder motion were larger for sagittal flexion/extension (62.3° and 68.6°) and horizontal flexion/ extension (84.6° and 93.6°).
Table 1. Group propulsion characteristics for both speed conditions: 0.9 m/s and 1.8 m/s. Group means and standard deviations are in parentheses. Push time is time during which hand was in contact with push rim. Start and contact angles are reported relative to horizontal axis.
Speed (m/s)
Actual Speed (m/s)
Cadence (stroke/s)
Total Cycle (s)
Push Time (s)
Start Angle (°)
Contact Angle (°)
0.9
0.97 (0.12)
0.97 (0.20)
1.06 (0.19)
0.52 (0.09)
119.55 (11.4)
100.3 (16.6)
1.8
1.61 (0.23)
1.32 (0.22)
0.77 (0.12)
0.35 (0.05)
119.03 (10.5)
110.3 (15.8)
639 KOONTZ et al. Shoulder kinematics and kinetics during propulsion
Figure 2. Shoulder forces versus time during propulsion at 0.9 m/s: Five consecutive strokes for a single subject are represented. Beginning of stroke is at time = 0 s. Anterior, medial, and superior force components were positive. Blue area represents transition between propulsion phase and recovery phase.
Shoulder Kinetics During Propulsion Table 3 and Table 4 display the group’s mean peak 3D net shoulder forces and moments and the relative timing of the peaks for both speed conditions.
Forces At both speeds, the greatest forces were present during the propulsion phase with the highest magnitudes in the inferior direction (90.0 N, 108.2 N), followed by the anterior (59.9 N, 86.6 N), and then the medial directions (34.0 N, 50.4 N). As the stroke progressed, the magnitudes of the anterior, inferior, and medial force components continued to increase until they reached their peaks for each speed at 44.5 percent and 38.5, 63.6 and 55.5 percent, and 57.9 and 56.4 percent of the propulsion phase, respectively. After this point, the anterior force rapidly decreased and changed directions at the end of the propulsion phase and the beginning of recovery. Inferior forces generally decreased and leveled off at a constant value equal to the weight of the limb. Forces in the anterior/posterior and medial/lateral directions diverged to values close to zero. Only four subjects displayed forces directed in the superior direction during the 1.8 m/s speed trial. The magnitude of this force was small (
