Locomotor Response of Persons With Upper Limb Loss to Treadmill Perturbations
Phase: NA
Conditions: Amputation
Sponsor: VA Office of Research and Development

Study Title: Locomotor Response of Persons with Upper Limb Loss to Treadmill 
Perturbations 
 
NCT: NCT04274218 
 
Document Date: April 2nd 2024 
 

Background and Study Design 
The role of the upper limbs in gait dynamics has received little attention historically, 
however surmounting evidence supports an important function of the arms in balance during 
locomotion and other dynamic tasks. Regaining balance after a perturbation is contingent on 
reactive mechanisms for re-stabilization. Rapid arm elevation is commonly observed in able-
bodied individuals following a perturbation, which acts to increase the moment of inertia (I), 
slowing trunk motion. In any movement context, it is plausible that the result of a perturbation, 
i.e., a fall or recovery, could rest on the ability to effectively manipulate arm motion. A high fall 
rate has been observed in individuals with upper limb absence (ULA). A large proportion of falls 
in persons with ULA (67%) has been reported to occur during walking activities. Effects of ULA-
induced imbalance can be observed in an increase in sound arm swing during walking. Use of an 
upper limb prosthesis, which effectively reduces the mass discrepancy, might be a solution to this 
issue of imbalance. Reactive responses are yet to be explored in individuals with ULA with and 
without use of a prosthesis. The objective of this cross-over intervention study was to investigate 
the reactive responses of people with ULA to perturbations during walking. 
 
Protocol 
Participants completed two baseline walking trials and 12 perturbation trials on an 
instrumented treadmill (Motek, Enschede, Netherlands). Baseline trials consisted of 30 seconds of 
steady-state walking, first at a self-selected speed then 1.0 m/s. The perturbation was designed to 
disrupt walking and instigate trip-like responses: the treadmill belt was programmed to accelerate 
(6.5 m/s2) to 3.3 m/s, then immediately decelerate (6.5 m/s2) to 1.0 m/s. During each trial a 
perturbation was triggered at initial contact of a randomly-selected step between 21 and 40 after 

the belt reached 1.0 m/s and delivered during single limb stance of the sound/dominant or 
impaired/non-dominant arm side (6 per side). Participants walked for twenty strides after 
restoration of reciprocal gait before the treadmill was stopped. Side/step randomization minimized 
anticipatory effects of knowledge of the perturbation timing on motor response. Prosthesis users 
completed all tasks with and without their customary prosthesis, order randomized.  
Mass, COM location and I of each prosthesis were estimated using an oscillation cage and 
reaction board. I was estimated about the medial-lateral axis of rotation at the prosthesis proximal 
end and translated to its COM using the parallel axis theorem.   
 
Data Analysis 
Kinematic data were collected with a 12-camera motion capture system (Motion Analysis, 
Rohnert Park, CA) at 120 Hz. Data were exported to Visual 3D (C-motion, Germantown, MD), 
filtered using a 4th-order low-pass (9 Hz) Butterworth filter, and a customized 12-segment model 
was applied. The assumption of uniform density [36] was applied to estimate segment mass as a 
percentage of body mass, COM position and I, except for the ULA impaired limb where the 
forearm was modified based on the prosthesis condition.   
Initial contact events were labelled using a velocity-based algorithm and corrected 
manually. L was estimated by calculating the instantaneous angular momentum of each segment 
about the COM in three planes, summing the values according to Eq. 1, where, for each segment 
i, 𝜔𝑖
⃗⃗⃗⃗  is the angular velocity, 𝑟𝑖⃗  is the distance from segment COM to whole-body COM, 𝑚𝑖 is the 
mass and 𝑣𝑖⃗⃗⃗  is relative velocity.  
𝐿⃗ = ∑
𝐼𝑖⃗ 𝜔𝑖
⃗⃗⃗⃗ + 𝑟𝑖⃗ × 𝑚𝑖𝑣𝑖⃗⃗⃗ 
𝑛
𝑖=1
 
Eq. 1 

L was normalized by treadmill speed and participant mass and height. Lrange was calculated 
as maximum minus minimum L over the period of interest. For baseline trials, Lrange was calculated 
over each stride and averaged over the middle 10 strides. The perturbation recovery phase was 
defined as perturbation onset time to the seventh initial contact, and Lrange was calculated from 
global maximum and minimum across that period. To aid interpretation of Lrange, bilateral shoulder 
flexion-extension and add-abduction ranges were calculated similarly for each period of interest. 
Shoulder flexion-extension was defined as the sagittal-plane angle between the upper arm long 
axis (line connecting shoulder and elbow markers) and thorax long axis (line connecting midpoint 
between shoulder markers and pelvis center) and add-abduction angle was that relative angle in 
the coronal-plane. Five perturbation trials on each side were averaged with the first (naïve) trial 
excluded because of exaggerated reactions to first-side perturbation. 
 
Statistical Analysis 
Descriptive statistics are produced for Lrange (recovery minus baseline), shoulder joint 
range-of-motion during recovery, maximum trunk flexion angle, and maximum trunk flexion 
velocity for each condition (controls: with arms free and with non-dominant arm bound; upper 
limb absence: without wearing the prosthesis and with wearing the prosthesis).