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Development and Evaluation of a Sensor System to Monitor the Stance-Phase Control Function of the Automatic Stance-Phase Lock (ASPL) Mechanism
by
Jessica N. Tomasi
A thesis submitted in conformity with the requirements for the degree of Master of Health Science, Clinical Engineering
Institute of Biomaterials and Biomedical Engineering University of Toronto
© Copyright by Jessica N. Tomasi 2016
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Development and Evaluation of a Sensor System to monitor the
Stance-Phase Control Function of the Automatic Stance-Phase
Lock (ASPL) Mechanism
Jessica N. Tomasi
Master of Health Science, Clinical Engineering
Institute of Biomaterials and Biomedical Engineering
University of Toronto
2016
Abstract
The Automatic Stance-Phase Lock is the novel stance-phase control mechanism employed by the
All-Terrain Knee. Gait analysis tools are often limited to controlled environments and cannot
directly monitor the ASPL. The objective of this project was to design and test a sensor
system to measure ASPL function and to begin to explore the effects of relevant alignment,
terrain, and mobility conditions on its performance.
The results of this study indicate that the developed system is sensitive to knee lock position
changes, knee extension and flexion, and gait events. Data collected by the system confirms the
fundamental relationships between applied moments and knee lock engagement which defines
ASPL stance-phase control. Measurable differences in ASPL function allude to its
responsiveness to variable gait conditions.
The developed system has the proven potential for use in larger biomechanical and clinical
studies to inform All-Terrain Knee design iterations and optimize patient-specific prosthetic
alignment and set-up.
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Acknowledgments
I would first like to acknowledge my thesis advisor, Dr. Jan Andrysek, of the Institute of
Biomaterials and Biomedical Engineering at the University of Toronto for his continual
leadership and support throughout this research and thesis-writing process. I would like to thank
Dr. Matthew Leineweber of the Bloorview Research Institute for his generous guidance and
input over the last two years and Brandon Burke of LegWorks for his invaluable insight. I would
like to acknowledge Amy Richardson, Brian Steinnagel, Neil Ready, and Sandra Ramdial of the
Holland Bloorview department of Prosthetics and Myoelectrics for so kindly sharing their
clinical knowledge and resources; Drs. Steve Ryan and Emil Schemitsch for their valuable
advice and direction; and Rhonda Marley for every prompt response to countless frantic emails.
To Daniel, Rachel, and Victoria, thank you for your contributions to this study and for believing
in its clinical implications.
I have been tremendously fortunate to be surrounded by so many classmates and colleagues who
have motivated and inspired me and whose friendship has undoubtedly enriched my Master’s
experience. Lauren and Emily, I can’t thank you enough. To Dr. Anne-Marie Guerguerian of the
Hospital for Sick Children, thank you for your unwavering positivity and encouragement.
I am sincerely grateful to my best friends, Alexa, Alycia, Amanda, Jordana, and Nicole, for their
infallible support and kind reassurance throughout these years. It means more to me than you
know.
And finally, to my incredible family, whose humbling and unshakable confidence in me saw me
through to the finish: thank you, thank you, thank you! Mom, Dad, and Becca, this is for you.
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Table of Contents
Acknowledgments.......................................................................................................................... iii
Table of Contents ........................................................................................................................... iv
List of Tables ................................................................................................................................. vi
List of Figures ................................................................................................................................ ix
Introduction .................................................................................................................................1
1.1 Thesis Roadmap ...................................................................................................................1
1.2 Background ..........................................................................................................................2
1.2.1 Lower Limb Loss .....................................................................................................2
1.2.2 Amputee Gait and Function .....................................................................................2
1.2.3 Prosthetic Knee Joints ..............................................................................................3
1.2.4 Automatic Stance-Phase Lock .................................................................................5
1.3 Research Problem ..............................................................................................................11
1.4 Research Objectives ...........................................................................................................12
Research Methods .....................................................................................................................13
2.1 Instrumentation ..................................................................................................................13
2.1.1 Automatic Stance-Phase Lock-Sensing System (ASPL-SS) .................................13
2.1.2 Portable Force and Torque Transducer ..................................................................20
2.2 Study Design ......................................................................................................................23
2.2.1 Participants .............................................................................................................23
2.2.2 Experimental Procedure .........................................................................................26
2.3 Data Analysis .....................................................................................................................30
2.3.1 Data Synchronization .............................................................................................31
2.3.2 Engineering Validation ..........................................................................................32
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2.3.3 Measuring ASPL Function ....................................................................................38
2.3.4 Comparing Conditions ...........................................................................................46
Results and Discussion ..............................................................................................................50
3.1 Sensitivity Analysis ...........................................................................................................50
3.2 Spatiotemporal Parameters ................................................................................................53
3.3 Engineering Validation ......................................................................................................54
3.3.1 Inductive Proximity Sensor: Bench Tests ..............................................................54
3.3.2 Force-Sensing Resistor ..........................................................................................55
3.3.3 Accelerometer ........................................................................................................56
3.4 ASPL Function...................................................................................................................58
3.4.1 Lock Displacement and Control Axis Moment .....................................................59
3.4.2 Knee Extension and Knee Axis Moment ...............................................................60
3.4.3 Knee Stability.........................................................................................................61
3.5 Limitations .........................................................................................................................68
3.6 Future Work .......................................................................................................................69
Conclusions ...............................................................................................................................71
References .................................................................................................................................73
Abbreviations and Glossary ...........................................................................................................79
Appendix A: Code .........................................................................................................................82
A.1 Arduino Data Logger .........................................................................................................82
A.2 MATLAB Synchronization and Analysis ..........................................................................85
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List of Tables
Table 1: Comparison of lock position sensor alternatives with respect to requirements. ............. 16
Table 2: The ASPL function and representative parameters measured by each component of the
ASPL-SS. ...................................................................................................................................... 17
Table 3: Test condition acronyms and conditions. ....................................................................... 27
Table 4: Order of conditions tested by able-bodied participants. ................................................. 29
Table 5: Sensor data sign convention. .......................................................................................... 30
Table 6: Objective 2 test and analysis overview. .......................................................................... 33
Table 7: Accelerometer calibration values. .................................................................................. 35
Table 8: Thresholds selected to define accelerometer and vertical force data features for temporal
comparison (Figure 17). Increases/decreases in acceleration and force were measured between
consecutive data points (collected at ~100Hz). ............................................................................ 37
Table 9: Objective 3 test and analysis overview. .......................................................................... 38
Table 10: Thresholds selected to define lock displacement and control axis moment data features
for temporal comparison (Figure 18). Increases/decreases in lock position and moment were
measured between consecutive data points (collected at ~100Hz). .............................................. 40
Table 11: Thresholds selected to define contact force and knee axis moment features for
temporal comparison (Figure 19). Increases/decreases in moment were measured between
consecutive data points (collected at ~100Hz). ............................................................................ 42
Table 12: Thresholds selected to define features for temporal comparison of knee stabilizing
events (Figure 20). ........................................................................................................................ 44
Table 13: Objective 4 analysis overview. ..................................................................................... 46
Table 14: Test conditions, their clinical relevance, and theoretically expected effects. ............... 47
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Table 15: Thresholds selected to define data features for comparison of lock engagement
duration (Figure 21). Increases/decreases in lock position and force were measured between
consecutive data points (collected at ~100Hz). ............................................................................ 49
Table 16: All able-bodied participant gait cycles in the NEUT condition were identified as either
drag or no-drag based on the presence or absence of the toe-drag features described above. The
results of paired t-tests used to compare drag and no-drag cycles for each parameter are shown
below. Statistically significant differences (p<0.05) are indicated by *. ...................................... 52
Table 17: A comparison of measured and literature averages for stride time (seconds) and stance
time (% gait cycle) in the NEUT condition. ................................................................................. 53
Table 18: Terminal impact events detected as a percentage of total gait cycles analyzed, listed by
condition. ...................................................................................................................................... 55
Table 19: Average horizontal and vertical acceleration values in stationary NEUT condition. ... 56
Table 20: Average temporal offset in gait event detection by the accelerometer and load
transducer (heel strike and toe-off) and accelerometer and force-sensing resistor (terminal
impact). Values represent gait cycles in the NEUT condition and are shown as %GC. Negative
values indicate that acceleration events preceded vertical or contact force events. ..................... 57
Table 21: Mean and standard deviation of temporal offset in gait event detection under different
translational alignment conditions for able-bodied participants. Values are shown as % gait
cycle. Negative values indicate that acceleration events preceded vertical and contact force
events. Statistically significant (p < 0.05) results of repeated measures ANOVA indicated by *.
....................................................................................................................................................... 57
Table 22: Average temporal offset between lock displacement and control axis moment events at
heel strike and mid-stance. Values are for able-bodied participant gait cycles in each condition
and are shown as %GC. Negative values indicate that lock displacement events preceded
moment events. Statistically significant (p < 0.05) results of repeated measures ANOVA
indicated by *. ............................................................................................................................... 59
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Table 23: Average temporal offset between contact force and knee axis moment events at heel
strike and mid-stance. Values are for able-bodied participant gait cycles in each condition and
are shown as %GC. Negative values indicate that contact force events preceded moment events.
Statistically significant (p < 0.05) results of repeated measures ANOVA indicated by *. .......... 60
Table 24: Average temporal offset between contact force and knee lock, and between knee axis
(KA) and control axis (CA) moment events at mid-stance and toe-off. Values are for able-bodied
participant gait cycles in each condition and are shown as %GC. Negative values indicate that
lock or control axis events preceded contact force or knee moment events. Statistically
significant (p < 0.05) results of repeated measures ANOVA indicated by *. .............................. 62
Table 25: The results of paired t-tests used to compare knee lock engagement duration between
each set of conditions. Mean and standard deviation values are shown as % stance-phase.
Statistically significant (p ≤ 0.017) results of paired t-tests indicated by *. ................................. 65
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List of Figures
Figure 1: The gait cycle. Stance-phase, beginning with heel-strike and ending with toe-off,
describes the period of ground contact and weight-bearing by a limb. Swing-phase is the period
of limb advancement and ends when the limb again makes contact with the ground at the
following heel strike. ...................................................................................................................... 4
Figure 2: The All-Terrain Knee: (A) assembled in a prosthesis; (B) with knee lock engaged to
prevent knee flexion in stance-phase; (C) with knee lock disengaged as it would be at toe off,
shown by red circle; and (D) shown flexed to about 120 degrees [31]. Throughout swing-phase,
the degree of flexion varies from approximately 0-70 degrees. ..................................................... 6
Figure 3: A depiction of the Ground Reaction Force Vector (GRFV) in (left) early and (right)
late stance-phase. (Left) Knee and control axes experiencing external flexion moments, tending
to flex the knee about the knee axis and drive the knee lock into the engaged position. (Right)
Knee axis and control axis experiencing external extension moments which drive the contact
interface closed and the knee lock back into the disengaged position. ........................................... 7
Figure 4: Internal springs and bumpers. ......................................................................................... 8
Figure 5: Sensor locations [36]. ...................................................................................................... 9
Figure 6: A sample of the data collected by Chen and Andrysek (2014) [36]. ............................ 10
Figure 7: Inductive proximity sensor and placement. ................................................................... 15
Figure 8: ASPL-SS sensor and data logger circuit schematic. ..................................................... 18
Figure 9: ASPL-SS sensor placement and mounting. (A) Inductive proximity sensor, (B) force-
sensing resistor, (C) accelerometer, (D) load transducer. ............................................................. 19
Figure 10: Front and side view of All-Terrain Knee instrumented with an ATI Mini58 F/T
Transducer using custom adapter plates. ...................................................................................... 21
Figure 11: Adjustable adapter plates used to modify anterior-posterior (A-P) translational
alignment between the All-Terrain Knee and prosthetic foot. Shown here in the (A) NEUT, (B)
POST, (C) POST1, (D) ANT1, and (E) ANT conditions. ............................................................ 21
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Figure 12: Schematic illustrating the offsets between the ATI Mini58 F/T Transducer Y-axis
(red), the ASPL control axis (pink) [Δx1 = 23.5mm, Δz1 = 64.3mm], and the ASPL knee axis
(maroon) [Δx2 = 16.5mm, Δz2 = 195.3mm] used to derive the moment applied at the control and
knee axes from the forces and torques acting at the ATI origin (coordinate axes shown) based on
the following equations: ................................................................................................................ 22
Figure 13: Prosthetic gait simulator assembly (left) and prosthetic gait simulator equipped with
load transducer (right). .................................................................................................................. 24
Figure 14: Walking trial data collection protocol. ........................................................................ 28
Figure 15: Vertical force and acceleration stomp feature synchronization. Stomp start features
were set at 0 seconds and subsequent timestamps were derived relative to them. ....................... 31
Figure 16: Pale blue shading indicates the series of vertical and horizontal acceleration points
between initial stomp and first gait cycle used to calculate average accelerometer values while
stationary on level ground in the NEUT condition. ...................................................................... 35
Figure 17: Acceleration, vertical force, and contact force data from one able-bodied participant
gait cycle in the NEUT condition. Features defined in Table 8 are indicated by the corresponding
number and colour. ....................................................................................................................... 37
Figure 18: Lock position and control axis moment data from one able-bodied participant gait
cycle in the NEUT condition. Negative lock position values indicate forward displacement/lock
engagement and positive values indicate backward displacement/lock disengagement. Negative
and positive moment values indicate flexion and extension, respectively. Features defined in
Table 10 are indicated by the corresponding number and colour. ................................................ 40
Figure 19: Contact force and knee axis moment data from one able-bodied participant gait cycle
in the NEUT condition. Negative and positive moment values indicate flexion and extension,
respectively. Features defined in Table 11 are indicated by the corresponding number and colour.
Inset shows a zoomed in view of the force and moment features at heel strike. .......................... 42
Figure 20: Lock position, contact force, knee and control axis moment data from one able-bodied
participant gait cycle in the NEUT condition. Negative lock position values indicate forward
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displacement/lock engagement and positive values indicate backward displacement/lock
disengagement. Negative and positive moment values indicate flexion and extension,
respectively. Features defined in Table 12 are indicated by the corresponding number and colour.
Inset shows a zoomed in view of the contact force event at heel strike. ...................................... 45
Figure 21: Lock position and vertical force data from one able-bodied participant gait cycle in
the NEUT condition. Negative lock position values indicate forward displacement/lock
engagement and positive values indicate backward displacement/lock disengagement. Features
defined in Table 15 are indicated by the corresponding number and colour. ............................... 49
Figure 22: Data from one above-knee amputee gait cycle (top), one able-bodied gait cycle with
no toe-drag (middle), and one able-bodied gait cycle with toe drag (bottom). Graphs on the left
depict lock position and control axis moment, graphs on the right depict contact force and knee
axis moment. Shaded grey regions highlight the period between toe-off and terminal impact
events. Negative peaks in lock position and moment which occur between toe-off and terminal
impact (indicated by red arrows) are not present in above-knee amputee gait cycles and suggest
that there was an external force applied to the limb during swing-phase. .................................... 51
Figure 23: An example of two consecutive above-knee amputee gait cycles under the same
condition (fast gait) with significantly different contact forces at terminal impact (red circles). 55
Figure 24: Average and standard deviation lock engagement duration for able-bodied
participants (n=8) under each condition (shown in black). Individual able-bodied participant
means shown in grey. Above-knee amputee means and standard deviations shown in dark grey
and broken lines for comparison. .................................................................................................. 65
Figure 25: Average lock engagement duration under different prosthetic alignment, All-Terrain
Knee setting, terrain, and walking speed conditions tested by the above-knee amputee
participant. .................................................................................................................................... 67
1
Introduction
1.1 Thesis Roadmap
The following document has been divided into four sections. Beginning here, section 1 provides
an overview of lower limb amputation and its clinical implications, a glimpse into the field of
prosthetic technology, and delves into the research and development of the All-Terrain Knee and
Automatic Stance-Phase Lock mechanism. With this background, section 1.3 highlights the
motivation for this study and section 1.4 outlines the specific research objectives and questions
which it seeks to address.
Section 2 details the methods by which the research problem was investigated. Sections 2.1 and
2.2 describe the development of the Automatic Stance-Phase Lock-Sensing System (ASPL-SS)
and the instrumentation and experimental protocols used to collect data for its evaluation.
Section 2.3 explains the data processing done to extract data features for comparison, and the
relevance of the analyses performed.
Section 3 presents and interprets the collected data and analysis outcomes, their applicability to
the research objectives, and implications for future work exploring and optimizing the function
of the Automatic Stance-Phase Lock.
To conclude, section 4 restates the research problem and summarizes the main findings of this
research and its clinical significance.
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1.2 Background
1.2.1 Lower Limb Loss
An estimated one in 150 North Americans are currently living with an amputation carried out to
treat conditions such as vascular disease, cancer, trauma, or congenital deformities; as average
life expectancy and the manifestations of conditions such as obesity and diabetes continue to
increase, the prevalence of limb loss is also expected to rise [1-3].
Studies suggest that approximately 65% of amputations involve the lower limb, of which about
60% are ‘major’, resulting in loss more proximal than the toes [1]. While amputation is often
seen as a final measure in vascular or orthopaedic treatment, it can relieve a patient of a painful
limb, and with the prescription of appropriate prosthetic components and proper treatment, can
allow for rehabilitation into a functional prosthetic ambulator [4].
1.2.2 Amputee Gait and Function
Recovery from transfemoral, or above-knee (AK), limb loss is a challenge demanding the
attention and expertise of a multifaceted clinical team of physicians, therapists, and prosthetists.
Even with the use of a prosthesis, lower limb amputees are often faced with persistent health
threats and physical challenges. Studies show that individuals with lower limb amputation have
decreased balance, energy efficiency, and walking speed, and demonstrate gait asymmetries,
reduced activity level, and difficulty ambulating over rough terrain, stairs, and hills.
Inappropriate load distribution at the stump-socket interface may result in vascular or neural
damage, skin breakdown, and pressure ulcers; the resultant discomfort often leads to over-
dependence on the contralateral limb, joint degeneration, and pain [5-23].
Aside from gait limitations, people with major limb amputation often experience environmental
barriers, participation restriction, and limited functioning levels as defined by the World Health
Organization’s International Classification of Functioning, Disability, and Health (ICF). The
most common environmental barriers encountered include climate, physical environment, and
income; participation restriction is often experienced in sports and physical recreation, leisure
and cultural activities, and job-seeking; and daily activities such as standing for long periods,
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walking long distances, and the emotional effects of disability are among the most challenging
[24].
1.2.3 Prosthetic Knee Joints
Componentry
The prescription of prosthetic components that address patient-specific functionality needs aims
to mitigate secondary injury and improve overall quality of life. The selection of an appropriate
knee component is especially critical to successful prosthetic function and sustained utilization
by an AK amputee [25]. In swing-phase, knee flexion is required to avoid unhealthy gait
deviations and, due to the loss of muscle control for the knee, AK amputees also depend on
prosthetic joints for stability in stance-phase (Figure 1).
A variety of knee joints exist, classified by their mode of articulation and means of controlling
this articulation. The two major classifications of articulation are monocentric and polycentric.
Monocentric mechanisms, such as a single axis knees, have fixed centers of rotation, relative to
the joint, which do not change during their operation [26]. The centers of rotation of polycentric
mechanisms, like four- and six-bar linkages, mimic the function of human knees and change
throughout the gait cycle, rendering them substantially more complex [26]. Generally,
polycentric joints add stability during early stance-phase and have better toe clearance during
swing-phase than monocentric alternatives [15]. External prostheses are further differentiated on
the basis of gait phase control. Controlling flexion in weight-bearing, also known as stance-
phase control, can be achieved by the relative alignment of prosthetic components, manual locks,
weight activated stance mechanisms, mechanical friction, polycentric mechanisms, or a
combination of these elements [27]. Toe clearance and the extent and rate of knee flexion are
dictated by swing-phase control, which may be applied by mechanical, hydraulic, or pneumatic
systems, and energy storing components such as springs [27]. Without swing-phase control, both
mono- and polycentric knees lack cadence response and are thus indicated for individuals who
ambulate at constant speeds [15].
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Stance-Phase Stability
Stability in stance-phase is the ability of the leg to resist flexion and remain supportive
throughout weight-bearing [15]. The Ground Reaction Force Vector (GRFV) is commonly used
to quantify the forces applied to the plantar surface of the foot during stance-phase. This single
vector is a summation of all external forces acting on the foot as a result of weight-bearing, and
acts through the foot’s center of pressure. The direction of the GRFV in the sagittal plane, its
magnitude, and its orientation with respect to stability-affecting axes of an articulating knee
joint, directly determine the external joint moments produced during stance-phase, and are thus
related to knee stability [28-30]. An AK amputee with weak residual musculature may have
trouble generating the joint moments necessary to maintain knee stability without changes to
alignment or the use of a brake or lock mechanism such as those mentioned in section 1.2.3.1.
The importance of stance-phase stability is highly rated by lower limb amputees and correlated
to their feeling of security, comfort, and fatigue during gait [27]. Like able-bodied ambulators,
lower limb amputees may be faced with uneven and inconsistent terrain, necessitating gait
modification and increasing the stability demands on their prostheses. Falls can occur when a
prosthetic knee joint is not fully extended at weight-acceptance or if it does not remain extended
throughout weight-bearing [29]. As discussed, prosthetic knee components provide stance-phase
Heel Strike Mid-stance Toe-off
Figure 1: The gait cycle. Stance-phase, beginning with heel-strike and ending with toe-off, describes the
period of ground contact and weight-bearing by a limb. Swing-phase is the period of limb advancement
and ends when the limb again makes contact with the ground at the following heel strike.
5
control, including stability, in a number of ways, usually in response to the loading of the joint
and the resultant magnitude and direction of the GRFV throughout the gait cycle [31].
Alignment
The spatial relationship of the socket relative to the other prosthetic components is referred to as
prosthetic alignment, and affects the stance-phase stability of a prosthetic leg. Besides the knee
joint, components of an above-knee prosthesis include the socket, thigh component (depending
on residuum length), shank component, and foot.
Unsuitable alignment can cause many of the functional and physiological challenges faced by
amputees and compromise their rehabilitation and health; thus, great care should be taken to
achieve appropriate alignment in clinical practice [9, 16, 17, 20, 32]. Understanding how
alignment affects the function of a particular prosthetic knee joint accelerates the alignment
process and improves patient comfort and safety. This information is generally provided by
component manufacturers. Knee axis offset from a vertical reference line and ankle joint flexion
often play a role in stability.
1.2.4 Automatic Stance-Phase Lock
Function
The All-Terrain Knee is a monocentric mechanical prosthetic knee joint suitable for transfemoral
amputees ranging from community ambulators to athletes and active children [33]. Three main
body components linked by two axes comprise the All-Terrain Knee: the thigh component
articulates about the knee body along the knee axis, and the control axis allows small rotations of
the knee body relative to the shin component (Figure 2). The All-Terrain Knee also houses the
novel Automatic Stance-Phase Lock (ASPL) mechanism, designed to provide the user with
stability in stance-phase without inhibiting swing-phase flexion and the natural progression of
gait.
The All-Terrain Knee is a multi-axis design: the ASPL employs a secondary “control” axis to
facilitate supplementary stability in stance-phase by manipulating the knee lock component. The
knee lock extends from the shin component interface toward the thigh component where it is
6
designed to engage and disengage to prevent or provide flexion as needed based on the moment
applied at the control axis. The most unstable phase of gait for an AK amputee is shortly after
heel strike. During loading response, the ground reaction force vector is directed upward and
backward, offset posteriorly from the knee joint and generating an external flexion moment
which creates an instant of potential knee instability [34]. In the ASPL mechanism, this early-
stance flexion moment at the control axis secures knee lock engagement, and ensures full
extension throughout weight-acceptance until an external extension moment naturally
hyperextends the knee and disengages the knee lock (Figure 3).
In terminal stance, the upper body moves forward faster than the tibia and the force vector again
moves behind the knee which, along with voluntary hip flexion, allows the joint to flex about the
knee axis and bend the limb in preparation for swing-phase [34].
Articulation during swing-phase is controlled by both the extension-assist spring assembly and
by adjustable variable friction where the thigh component and knee body are in contact.
Figure 2: The All-Terrain Knee: (A) assembled in a prosthesis; (B) with knee lock engaged to prevent
knee flexion in stance-phase; (C) with knee lock disengaged as it would be at toe off, shown by red circle;
and (D) shown flexed to about 120 degrees [31]. Throughout swing-phase, the degree of flexion varies
from approximately 0-70 degrees.
7
In terminal swing-phase, the extension-assist spring assembly automatically biases the knee
toward full extension while the knee lock spring drives the knee lock forward to again prepare
the limb for weight-bearing. Springs of different stiffness can be substituted and the knee lock
spring can be adjusted to accommodate the needs of different users.
Inside the knee joint, a series of springs and bumpers ensures that the knee is not noisy during
terminal impact and that there is adequate locking force to maintain stability (Figure 4). The
extension bumper is present to cushion terminal impact and reduce noise during knee extension.
The front lock bumper pad is held in place by the lock bumper post. The bumper pad cushions
the impact and reduces the noise of the knee lock as it pivots forward and comes in contact with
the knee body while moving into the locked position. Behind the knee lock, a second lock
bumper pad is held in place by the knee lock spring.
Figure 3: A depiction of the Ground Reaction Force Vector (GRFV) in (left) early and (right) late stance-
phase. (Left) Knee and control axes experiencing external flexion moments, tending to flex the knee
about the knee axis and drive the knee lock into the engaged position. (Right) Knee axis and control axis
experiencing external extension moments which drive the contact interface closed and the knee lock back
into the disengaged position.
8
Lock Bumper Post
Extension Bumper
Knee Lock Spring
Lock Bumper Pads Set Screw
Figure 4: Internal springs and bumpers.
9
Clinical Results
This simple design has demonstrated comparable performance to complex polycentric knees, and
studies show that the stance-phase control strategy is effective in reducing falls without
adversely affecting the biomechanics of gait. Walking speed with the All-Terrain Knee has been
more closely matched with high-end, than low-end knee component alternatives, offering a cost-
effective, reliable option for AK amputees [29, 31, 35]. Additional anecdotal evidence from
ongoing field trials also suggests that the already promising performance of the ASPL
mechanism could be improved further by optimizing its configuration and patient-specific setup
protocol. Confirming which points during the gait cycle cause the knee lock to engage and
disengage, and how this function varies with patient physiology and comorbidities under relevant
mobility conditions, and during different activities, would contribute to functionally enhancing
the design.
Some preliminary work has been done to study the status of the knee lock during amputee gait
with an All-Terrain Knee. Based on temporal correlation to motion capture software data, the
study by Chen and Andrysek (2014) was successful in using low cost force-sensing resistors to
identify the status of the knee lock during specific gait cycle events (Figure 5). While the
sensors used provided some insightful data, fixation issues and other challenges rendered them
impractical for use in more general and robust applications. Quantitative correlation of sensor
readings to moment magnitudes could not be carried out due to the low accuracy and
repeatability of the sensors [36].
Figure 5: Sensor locations [36].
Top sensor contact location
Front sensor contact
Back sensor contact
10
Figure 6 shows a sample of the data collected by Chen and Andrysek (2014) [36]. The first
graph shows the voltages, representing force, recorded from sensors placed at the ‘top’, ‘front’,
and ‘back’ of the All-Terrain Knee labelled in Figure 5. The thigh component of the knee exerts
pressure on the top sensor when the leg is in extension. The front and back sensors measure the
force with which the lock is pushing forward into the locked position and backward into the
unlocked position, respectively. The second graph illustrates the external moments acting on the
knee and on the lock itself. As discussed in section 1.2.4.1, the application of a flexion moment
at the control axis secures the knee in extension, and an extension moment disengages the lock to
prepare for knee flexion. The final graph shows the degree of knee flexion as measured by
motion capture cameras.
Extension Moment (+)
Flexion Moment (-)
Figure 6: A sample of the data collected by Chen and Andrysek (2014) [36].
11
1.3 Research Problem
Prosthetic kinetic and kinematic gait analysis is common and well-documented. Gold standard
gait analysis is often performed in instrumented laboratories with motion capture cameras and
floor-mounted load transducers for measuring the kinetics and kinematics of joints and limb
segments. These methods are costly and limited to highly controlled, indoor environments.
While technologies, such as high-precision load cells and inertial sensors, also exist and make
portable gait analysis possible, they often necessitate time-consuming sensor mounting and
calibration.
While the results of conventional gait analysis (1.2.4.2) offer compelling insight to the kinetic
and kinematic performance of the All-Terrain Knee and its functional advantages over
competing technologies, anecdotal evidence from ongoing clinical trials suggests that monitoring
the internal function of the All-Terrain Knee, and doing so outside the restrictions of a gait lab
would have significant research and clinical implications. Such data would inform mechanism
redesign and future All-Terrain Knee iterations, allow researchers to confirm whether or not the
ASPL functions as intended, demonstrate and improve ASPL function outside of a gait lab and
under different mobility conditions, and help the manufacturer provide the appropriate alignment
and set-up recommendations for different user populations. Clinically, such monitoring would
offer real-time feedback about ASPL function and knee stability, subjective data for the
systematic optimization of patient-specific alignment and set-up, and identification of long-term
component wear.
Early efforts to address this gap were limited by poor sensor fixation and accuracy, restricted to
use in a gait lab, and not suitable for conducting trials on a larger, more robust scale [36]. Thus,
the overall objective of this project was to design and test a sensor system for the purpose of
detecting and monitoring the ASPL stance-phase control function under relevant mobility
conditions.
12
1.4 Research Objectives
The explicit aims of this work and their related research questions were as follows:
1. Design a stand-alone, portable Automatic Stance-Phase Lock-sensing system (ASPL-
SS) for temporary application to an All-Terrain Knee, comprised of the necessary
sensors, their power supply, and means to record and convey sensor data to
researchers and clinicians.
2. Perform engineering validation tests on each sensor in the developed system to ensure
it performs to design goals and specifications. Does each sensor selected in Objective 1
provide the information it was intended to? Does the data logged from each sensor by the
developed system exhibit the known relationships between force, acceleration, and
displacement? Does the developed system perform comparably under variable conditions?
3. Describe the temporal relationships between external moment application at the
control and knee axes of the All-Terrain Knee and the resultant changes in knee lock
position and joint flexion at heel strike (HS), mid-stance (MS), toe-off (TO), and
terminal impact (TI). How do transitions between external extension and flexion
moments applied at the control axis and sudden increases or decreases in control axis
moment magnitude affect the movement of the knee lock as measured by the inductive
proximity sensor (IPS)? How do transitions between external extension and flexion
moments applied at the knee axis and sudden increases or decreases in knee axis moment
magnitude affect the flexion and extension of the All-Terrain Knee as measured by the
force-sensing resistor (FSR)? Does the ASPL mechanism provide stability without
impeding the natural progression of gait, i.e. does the knee lock engage and disengage in
time to secure knee extension at heel strike, ensure stability through mid-stance, and allow
knee flexion for toe-off?
4. Begin to explore the relationships between ASPL mechanism function and prosthetic
alignment, All-Terrain Knee settings, terrain, and walking speed. Can the developed
system detect differences in ASPL function between conditions? How do the tested
conditions affect ASPL function, and is it the expected effect?
The secondary aim of this study was to generate feasibility data that will inform both the design
of a clinic-ready model of the Automatic Stance-Phase Lock-Sensing System, and the planning
of larger clinical trials to evaluate ASPL performance with clinically relevant populations.
13
Research Methods
2.1 Instrumentation
2.1.1 Automatic Stance-Phase Lock-Sensing System (ASPL-SS)
The first objective of this work was to design a stand-alone, portable sensor system for
temporary application to an All-Terrain Knee, comprised of the necessary sensors, their power
supply, and means to record and convey sensor data regarding ASPL function to researchers and
clinicians. Section 2.1.1.1 outlines the design requirements for the ASPL-SS and its constituents.
ASPL-SS Design Requirements
The design requirements for the ASPL-SS were as follows:
Knee lock position detection: The knee lock is the crux of the Automatic Stance-Phase
Lock mechanism. It was designed to pivot around the control axis of the All-Terrain Knee,
engaging and disengaging based on the applied moment at the axis in order to prevent and
allow flexion of the joint as required by the user. In addition to moment about the control
axis, knee lock position is influenced by the lock spring. Depending on its tightness, the lock
spring can either bias the knee lock forward into the engaged position and increase the
stance-phase stability of the joint, or not make contact with the knee lock at all, leaving
voluntary control and externally applied extension moment to resist flexion. Measuring the
position of the knee lock throughout the gait cycle, relating it to control axis moment, and
comparing its motion under different lock spring conditions will aid in understanding the role
of each in ASPL function. Developing an instrument that can collect this data in larger
clinical studies will aid in informing future All-Terrain Knee design iterations. Clinically,
prosthetists will be able to use this information to monitor knee function under different
conditions while determining the appropriate patient-specific All-Terrain Knee setup.
Comparing the range of lock motion over time in a given setting will also provide insight to
the condition of internal components, such as bumper or lock spring deterioration.
Considerations for selecting a knee lock position sensor included the material of and access
to the knee lock, the portability of the system, precision, resolution, and cost.
14
Knee flexion/extension detection: If the All-Terrain Knee were a simple hinge, the flexion
and extension of the joint would be entirely dependent on the moment applied at its axis. The
All-Terrain Knee flexes and extends like a hinge about the knee axis, however, depending on
the moment applied at the control axis and lock spring tightness, the knee lock can oppose
the natural flexion of the knee. This should be the case at weight-acceptance when an applied
flexion moment at the knee axis would be expected to cause the knee to buckle without
additional support from the knee lock. Determining whether the knee is flexed or extended
throughout the gait cycle and comparing its status to lock position and applied moments at
the knee and control axes, will assist in identifying gait cycle events like terminal impact and
aid in understanding the role of the knee lock in maintaining knee extension under different
conditions. The measurement of this parameter should not impede the function of the lock.
Gait cycle event detection: In order to make gait analysis accessible and to contextualize
knee lock and flexion information with respect to the gait cycle outside a dedicated gait
laboratory, accurate detection of events such as the beginning and end of stance-phase is
required without using force plates. The selected sensor should be able to detect gait events
while affixed near the All-Terrain Knee.
Data logging: For research purposes, the data collected by the sensors should be sampled at a
consistent rate (≥100Hz), synchronized, stored, and accessible for post-processing. Data
collection trials should be separated for analysis.
Data transmission: In order to monitor the collected data in real time to ensure sensor
function, and eventually for clinical applications, data collected by the sensors should be
transmitted wirelessly for visualization by an investigator.
Portable: In order to perform analysis on any terrain or obstacle, the system must be
portable. The design must include the power supply for each of the sensors and data logger,
and a wireless means of data storage and transmission. The system should be relatively
lightweight and compact.
15
Easy-to-apply to and -remove from the All-Terrain Knee: Ultimately, the sensor system
should be a self-contained, easy-to-apply and -remove module which can be temporarily
attached to an All-Terrain Knee for intermittent monitoring on multiple users.
ASPL-SS Components
2.1.1.2.1 Knee Lock Position Detection
An inductive proximity sensor (1600Hz; 0-4mm; resolution: ≤1μm; output voltage: 0-10V; Ø
8mm; length: 45mm) (DW-AD-509-M8-390 Contrinex; Givisiez, Switzerland) was used to
measure the anterior-posterior movement of the knee lock continuously throughout the gait
cycle. Inductive proximity sensors emit an alternating electro-magnetic sensing field. When the
metal knee lock enters the sensing field, eddy currents are induced in the knee lock, which
reduce the signal amplitude and trigger a change of state at the sensor output.
An inductive proximity sensor was selected for this application for a number of reasons: the knee
lock is made of tool steel and is located inside the knee with limited access to, and space around,
it; the sensor system should be portable, wireless, and relatively light-weight; and the sensor
itself needs to detect small increments precisely and with high resolution, should not have many
moving parts (for ease of maintenance), and should be relatively low cost (Table 1).
In this application, the sensor is positioned orthogonally to the knee lock component through an
opening in the front of the All-Terrain Knee knee body (Figure 7).
Figure 7: Inductive proximity sensor and placement.
16
Table 1: Comparison of lock position sensor alternatives with respect to requirements.
Sensor type Met
al
det
ecti
on
Co
mp
act
Ea
sy-t
o-a
pp
ly t
o A
-T K
nee
Ba
tter
y-p
ow
ered
Lig
htw
eig
ht
Hig
h-r
esolu
tio
n a
nalo
g
inp
ut
Ea
sy-t
o-m
ain
tain
Rel
ati
ve
cost
Notes
Pressure sensor N/A Noisy $
Optical sensor N/A $$$ Bulky, alignment issues, too
sensitive
Lasers N/A $$$ Bulky, alignment issues, too
sensitive
Limit switch N/A N/A $ Most digital; analog
positioning limit switches $$$
Capacitive proximity
sensor $$
More suitable for plastic
Mechanical position
sensor N/A $$
Inductive proximity
sensor $$
2.1.1.2.2 Knee Flexion/Extension Detection
FlexiForce A201 sensors (response time: ≤5μs; 0-110N; thickness: 0.208mm; length: 51mm;
sensing Ø: 9.53mm) (Tekscan, Inc.; South Boston, Massachusetts) are force-sensing resistors
(FSR) which can be incorporated into a force-to-voltage circuit in order to measure the contact
force between two surfaces. One FlexiForce sensor was placed at the contact interface between
the thigh component and the knee body, similar to the “top sensor” of Chen and Andrysek (2014)
(1.2.4.2) (Figure 5). When the force sensor is unloaded, its resistance is very high and decreases
when a force is applied. This resistance can be measured and monitored. In this application, the
FlexiForce readings were used to detect the flexion and extension of the All-Terrain Knee by
monitoring binary changes (i.e. ≥1 or <1) in contact force. The thin profile of the sensor not only
allowed for the collection of data without disrupting knee locking and unlocking, but facilitated
the distinction between knee hyperextension, wherein the thigh component and knee body of the
All-Terrain Knee are in contact, and extension supported by the knee lock (i.e. when the
components are drawn apart by an applied flexion moment but the engaged knee lock prevents
17
the All-Terrain Knee from flexing). Since the limb is essentially extended in both natural and
lock-supported extension, monitoring knee angle would not necessarily indicate transitions
between the two states, although instruments such as goniometers or inertial measurement units
would provide additional kinematic details, especially in swing-phase. In terms of the final
application of the ASPL-SS -- being an easy-to-apply and -remove module -- a goniometer
would be more challenging to incorporate than a force-sensing resistor.
2.1.1.2.3 Gait Cycle Event Detection
A linear accelerometer (0.5-550Hz; ±3g; sensitivity: 360mV/g) (ADXL 335 Analog Devices,
Inc.; Norwood, Massachusetts) was used to record anterior/posterior and vertical components of
acceleration at the proximal end of the prosthetic shank. Body-mounted accelerometers have
become a viable alternative to foot contact switches, which are common for gait event detection
but prone to breakage and necessarily mounted on the foot, rather than the knee [37]. Use of a
knee-mounted accelerometer also allowed for event detection during the swing-phase of gait. A
summary of the physical parameters measured by each sensor is presented in Table 2.
Table 2: The ASPL function and representative parameters measured by each component of the ASPL-
SS.
Inductive Proximity Sensor Force-Sensing Resistor Accelerometer
Physical
parameters
measured
Anterior/posterior lock
position relative to front of
the All-Terrain Knee
Force applied at the
interface between thigh
component and knee body
of All-Terrain Knee
Acceleration of the All-
Terrain Knee in the
vertical and
anterior/posterior
directions
Relevant load
transducer
(§2.1.2)
measurements
External moment about the
control axis
External moment about
the knee axis
Vertical force
Relevant
aspect of
ASPL
function
Relationship between the
orientation and magnitude of
the applied control axis
moment and the position and
transitions of the knee lock
Initiation of extension and
flexion of the knee joint,
and differentiation
between extension
moment applied at the
knee axis and knee lock-
supported extension
Gait cycle event detection
in both stance- and swing-
phase
18
2.1.1.2.4 Data Logging and Transmission
Values measured by the inductive proximity sensor, force-sensing resistor, and accelerometer
were collected by an Arduino UNO data logger (Figure 8), transmitted wirelessly to a nearby
laptop for real-time viewing in the Arduino IDE serial monitor, and stored in a comma separated
values (.csv) file on an SD card for post-processing using MATLAB (MathWorks; Natick,
Massachusetts) (Appendix A). An Arduino Wireless/SD Shield mounted on the central
microprocessor housed an XBee Pro Series 1 Wireless Networking RF Module (Digi
International; Minnetonka, Minnesota) and a MicroSD port. The data logger collected sensor
readings at >100 Hz. A wearer-operated switch started and stopped writing to the SD card to
separate data collection trials.
2.1.1.2.5 Power Supply
Power for the Arduino UNO, XBee module, force-sensing resistor, accelerometer, switch, and
LED was provided by 6 AA batteries which provided 9V and about 2400mAh. The inductive
proximity sensor was powered by 2 9V batteries (18V, ~500mAh).
Figure 8: ASPL-SS sensor and data logger circuit schematic.
19
2.1.1.2.6 Mounting
The ASPL-SS was designed with the ultimate goal, to produce a self-contained, easy-to-apply
and -remove sensing module, in mind. The designed system (i.e. sensors, data logger, circuitry,
and power supply (Figure 8)) is portable and can be made to fit in a 15x10x7cm container. The
design of the final mounting enclosure is outside the scope of this project and is being developed
by another student in the PROPEL Lab.
For testing purposes, each of the sensors was individually mounted on the All-Terrain Knee
component used by all participants. The inductive proximity sensor was held in place by a
system of off-the-shelf brackets affixed by nuts and bolts. The threaded design of the sensor
itself facilitated precise positioning using nuts and plastic washers. The force-sensing resistor
was adhered to an L-bracket extension of the proximity sensor system such that it was flush with
the top face of the knee component and its force-sensing area was compressed by the extension
bumper during knee extension. The accelerometer was held in place on the anterior face of the
knee by adhesive Velcro pads with its ‘+X’ and ‘+Z’ axes oriented in the superior and anterior
directions, respectively (Figure 9). There was no formal protocol in place for confirming the
position and fixation of the sensors throughout the experiments, but physical adjustments were
made when issues were identified visually throughout data collection. The sensors were
connected to the batteries and circuitry via wires extending from the All-Terrain Knee to a fanny
pack worn by study participants.
Figure 9: ASPL-SS sensor placement and mounting. (A) Inductive proximity sensor, (B) force-sensing
resistor, (C) accelerometer, (D) load transducer.
A
B
C
D
20
2.1.2 Portable Force and Torque Transducer
In order to validate the function of the ASPL-SS components, as well as to examine the
relationships between knee lock position, kinematic gait events, the force and torque applied at
the knee (Objective 3), prosthetic legs used for data collection were instrumented with an ATI
Mini58 three-axis (six-degree-of-freedom) force and torque transducer (ATI Industrial
Automation, Inc.; Apex, North Carolina) (Figure 10). Portable, six-degree-of-freedom load
transducers have been cross-validated for use in lower-limb prosthetics research, and can be used
to identify specific gait cycle events, such as heel strike and toe-off during complex mobility
tasks in unconstrained environments [38-39].
Consultation was sought from a certified prosthetist in order to integrate the transducer with the
appropriate adapters to interface and variably align the prosthesis. A Computer Aided Design
(CAD) model, was used to ensure that the adapters fit within the appropriate limb length and to
design custom interfacing parts (Figure 11).
Since the load transducer has a different coordinate system origin than the All-Terrain Knee
axes, a transformation equation was used to calculate the equivalent forces and moments acting
on the control and knee axes (Appendix A.2) (Error! Reference source not found.). Data were
sampled at over 100 Hz using the transducer, transferred via a wearable battery-powered ATI
Wireless F/T Transmitter (ATI Industrial Automation, Inc.; Apex, North Carolina) to a nearby
laptop, and displayed by the ATI Wireless F/T Java software from which they were recorded to
.csv files for analysis.
21
Figure 11: Adjustable adapter plates used to modify anterior-posterior (A-P) translational alignment
between the All-Terrain Knee and prosthetic foot. Shown here in the (A) NEUT, (B) POST, (C) POST1,
(D) ANT1, and (E) ANT conditions.
( )
( )
( ) ( ) ( )
All-Terrain Knee
Custom Interface
Adapters
ATI Mini58
F/T Transducer
Control Axis
ATI Y-Axis
Knee Axis
Figure 10: Front and side view of All-Terrain Knee instrumented with an ATI Mini58 F/T Transducer using
custom adapter plates.
22
+z
+x
+y
Figure 12: Schematic illustrating the offsets between the ATI Mini58 F/T Transducer Y-axis (red), the
ASPL control axis (pink) [Δx1 = 23.5mm, Δz1 = 64.3mm], and the ASPL knee axis (maroon) [Δx2 =
16.5mm, Δz2 = 195.3mm] used to derive the moment applied at the control and knee axes from the forces
and torques acting at the ATI origin (coordinate axes shown) based on the following equations:
Control Axis Moment = TY+ FX*Δz1 + FZ*Δx1
Knee Axis Moment = TY+ FX*Δz2 + FZ*Δx2
Where TY represents torque about the y-axis (in the sagittal plane), and FX and FY represent force along
the x- and y-axes, respectively.
23
2.2 Study Design
2.2.1 Participants
Inclusion Criteria: Above-Knee Amputee Participants
In order to be considered for participation in this experiment, participants with AK amputations
had to (1) have a well-fitting and functional prosthesis that included the All-Terrain Knee; (2) be
experienced with using the All-Terrain Knee; (3) be capable of community level ambulation with
variable cadence without the use of ambulatory aids; (4) be able to negotiate stairs and slopes;
(5) have adequate shin length (minimum 10cm/4” between base of All-Terrain Knee and the top
of a prosthetic SACH foot) to be fitted with a prosthetic shank component including load
transducer and the applicable adapters; (6) weigh less than the maximum loading capacity of any
component of the instrumented prosthesis (maximum 100kg/220lb); (7) be above the age of 18;
and (8) be able to communicate in English.
Inclusion Criteria: Able-Bodied Participants
To be considered for participation in this experiment, able-bodied participants had to (1) be tall
enough (at least 140cm/4’7”) to be fitted with a prosthetic simulator including an adult sized All-
Terrain Knee, load transducer, and the applicable adapters; (2) weigh less than the maximum
loading capacity of any component of the instrumented prosthetic simulator (maximum
100kg/220lb); (3) be above the age of 18; (4) be able to communicate in English; and (5) be
strong, independent ambulators.
Sample Size
One AK amputee and eight able-bodied participants were recruited for participation in the study.
In relevant previous studies evaluating the use of load cells and prosthetic gait simulators, and
the effects of prosthetic alignment changes, common sample sizes ranged from one to 10
participants [38, 40, 41]. The applicable AK amputee population is limited; there was a pool of
four All-Terrain Knee users from which we could draw. In order to augment the sample size, a
convenience sample of eight able-bodied participants were recruited to test the ASPL-SS while
wearing a prosthetic gait simulator (Figure 13).
24
With one AK amputee and eight able-bodied participants each completing at least 20 data
collection trials under various conditions, this exploratory study aimed to generate sufficient data
to provide a performance assessment of the ASPL-SS and to begin to assess relationships
between applied force/torque, knee lock position, and kinematic gait events.
Recruitment
The participant with amputation was recruited from the known population of adults currently
using an All-Terrain Knee component on a regular basis (n=4).
A convenience sample of able-bodied participants was recruited primarily from the Bloorview
Research Institute. Potential participants were identified through self-referral in response to
advertisement on Holland Bloorview’s Participate in Research webpage.
Once potential participants were identified, they were contacted by email and the experiment was
described in more detail. Data collection sessions were scheduled at the convenience of
interested participants. Before commencing the sessions, additional questions were addressed
and written consent was obtained from each participant.
A
Figure 13: Prosthetic gait simulator assembly (left) and prosthetic gait simulator equipped with load
transducer (right).
25
Study Participants
One participant with unilateral transfemoral amputation (male; age: 24 yr; weight: 81.5kg
(wearing everyday prosthesis); height: 180cm; knee to ground: 33cm) and 8 able-bodied
participants (5 male and 3 female; age [mean ± standard deviation]: 27 ± 11yr; weight: 66 ± 10kg
(measured without prosthetic simulator); height: 171 ± 8cm; shin length: 44 ± 3cm) participated
in this study. The above-knee amputee participant had had an amputation of the left leg for 6.5
years at the time of the study, and had used the All-Terrain Knee for the last three. He wore a
skin fit suspension system and ischial containment socket and had a Locomotor Capability Index
(LCI) of 56. His amputation was caused by cancer. The able-bodied participants all wore the
prosthetic simulator on their right legs.
Ethical Considerations
Prospective participants were assured that they were under no obligation to participate, and that
their consent could be withdrawn without prejudice to pre-existing entitlements; this included the
option to withdraw any collected data up until study completion when identifying information
was destroyed. Prospective participants were informed of the potential risks associated with the
study protocol. For able-bodied participants, risks included discomfort caused by the simulator
socket-skin interface and muscle or joint stiffness due to bearing weight on the knee. It was
expected that simulated prosthetic gait would be unstable initially. Ambulatory aids (parallel
bars) were available throughout the training and data collection sessions. For amputees, the
prosthetic instrumentation itself posed no additional risks to those experienced in everyday wear.
Alternate alignment conditions tested were within the acceptable perturbation range as
determined by past research [41-42], and if the participants had expressed any concern or feeling
of discomfort about a particular testing condition, that portion of the trial would not have been
executed. Breaks were requested and taken as necessary by all participants. There was a
monetary incentive for participants.
At the start of the study, each participant was assigned a random identification (ID) code. The
identifying information linked to the code was stored separately from the collected gait data in a
locked cabinet. Only the principal investigator (Jan Andrysek) and the research coordinator
(Jessica Tomasi) had access to the data. Upon completion of the study, the link between
26
participant and ID code was destroyed. Data were analyzed and presented using only the
identification code. All personal information was kept confidential. If the results of the study are
published, participant names and identifying information will not be used. Following the
completion of the study, data will be saved in its anonymized state for seven years as required by
Holland Bloorview, after which it will be destroyed.
This experiment was approved by the Holland Bloorview Research Ethics Board and University
of Toronto Office of Research Ethics. Express written consent was obtained from all participants
prior to beginning data collection.
2.2.2 Experimental Procedure
Protocol: Above-Knee Amputee Participants
The prosthetic knee and shank of the AK amputee participant were replaced with the test All-
Terrain Knee equipped with the ASPL-SS prototype as well as a load transducer and custom
interface plates (Figure 10). The participant’s prosthesis was then realigned. In order to limit the
number of variables affecting the outcomes, all participants wore a solid ankle cushion heel
(SACH) foot for data collection trials. Lock spring and friction shim tightness were set on the
test knee and held constant for all participants and trials aside from the lock spring variation
conditions (i.e. conditions TTLS and LSLS, described in Table 3). All participants also used an
extension assist spring of equal stiffness.
The above-knee amputee participant was asked to perform a series of warm-up trials along a 7m
walkway in Holland Bloorview’s Gait Laboratory before completing data collection trials. For
the AK amputee participant, a trial was defined as one length of the walkway in one direction, a
complete set of stairs in one direction, or ascending or descending the length of a ramp once. In
addition to the alignment conditions tested by able-bodied participants (i.e. below-knee sagittal
translations), the AK amputee was asked to complete trials with changes to the angular
alignment of his prosthetic foot (i.e. trials with 4.6° of plantar- and dorsiflexion, one full turn of a
set screw in either direction), different lock spring settings (i.e. one half turn of the set screw in
either direction), and with the addition of the stance-phase flexion mechanism (Table 3).
27
Based on his practice and prior experience with prosthetic gait and use of an All-Terrain Knee,
the amputee participant was also asked to complete a series of data collection trials outside of the
lab, under different mobility conditions. Mobility conditions included fast walking, ascending
and descending stairs and a slope of 10°, and stationary cyclic knee loading (Table 3). These
alternate conditions were tested with the neutral alignment. These conditions were chosen based
on their proximity to our facility, as well as for their applicability to everyday ambulation.
Table 3: Test condition acronyms and conditions.
Code Condition Alignment/ All-
Terrain Knee Setting Terrain Speed AK AB
NEUT Neutral Neutral Flat, level Self-selected
ANT Anterior Foot
Translation
2cm anterior Flat, level Self-selected
POST Posterior Foot
Translation
2cm posterior Flat, level Self-selected
PFLX Plantarflexion 4.6° plantarflexion Flat, level Self-selected
DFLX Dorsiflexion 4.6° dorsiflexion Flat, level Self-selected
TTLS Tight lock
spring
½ CW turn of lock
spring set screw
Flat, level Self-selected
LSLS Loose lock
spring
½ CCW turn of lock
spring set screw
Flat, level Self-selected
FST Fast Neutral Flat, level Fast
SLW Slow Neutral Flat, level Slow
LZY Lazy gait Neutral Flat, level ‘Lazy’
RCK Rocks Neutral Small rocks Self-selected
RKSL Slow rocks Neutral Small rocks Slow
GRS Grass Neutral Grass Self-selected
INC Incline Neutral 80° Ramp, up Self-selected
DEC Decline Neutral 80° Ramp,
down
Self-selected
ASND Ascend Neutral Stairs, up Self-selected
DSND Descend Neutral Stairs, down Self-selected
SFSS Stance-flexion Stance-flexion Flat, level Self-selected
SFFS Stance-flexion
slow
Stance-flexion Flat, level Fast
SFSL Stance-flexion
fast
Stance-flexion Flat, level Slow
SFLZ Stance-flexion
lazy
Stance-flexion Flat, level ‘Lazy’
ANT1 Anterior Foot
Translation 1
1cm anterior Flat, level Self-selected
POST1 Posterior Foot
Translation 1
1cm posterior Flat, level Self-selected
28
The participant completed three data collection trials under each condition. Due to the number of
conditions tested, the amputee participant was asked to attend two sessions, each lasting about
two hours, at Holland Bloorview Kids Rehabilitation Hospital. A fanny pack containing the data
logger circuitry and power supply for the ASPL-SS was donned by the participant. Cables from
the sensors were anchored on the participant’s leg or belt and extended to the fanny pack. The
load transducer was connected to the wireless transmitter and clipped to the participant’s belt or
fanny pack strap. All wireless signal transmission was verified on a nearby laptop by an
investigator prior to and throughout data collection. For each new trial, the load transducer
software was started and stopped by the investigator at the laptop while the participant was asked
to turn on the data logger switch, stomp with the instrumented limb, pause, complete the walking
trial, and turn the switch off (Figure 14). Data were collected in the following order: STAT,
DEC, INC, DSND, ASND, NEUT, LZY, FST, SLW, ANT, POST, PFLX, DFLX, TTLS, LSLS,
RCK, RKSL, GRS, SFSS, SFFS, SFSL, SFLZ.
Protocol: Able-Bodied Participants
In order to participate in the study, able-bodied participants were fitted with a prosthetic
simulator to mimic the function of an above-knee prosthesis. Above-knee prosthetic simulators
with articulating knee joints have been shown to produce stride parameters, joint kinematics, and
net joint kinetics that are consistent with those from prosthetic users [40]. The socket of the
Figure 14: Walking trial data collection protocol.
29
simulator strapped on to the bent knee of the participant with Velcro and included the test All-
Terrain Knee fitted with the ASPL-SS prototype and the load transducer (Figure 13). The
simulator was aligned in the NEUT condition and adjusted for each participant by a member of
the research team trained to do so by a certified prosthetist.
Once the simulator had been properly configured, each subject completed a gait training session
in the NEUT condition with the support of parallel bars. Gait training was provided by a member
of the research team. Fitting, training, and data collection took place during a two hour session at
Holland Bloorview Kids Rehabilitation Hospital.
Data collection trials began once the prosthetic simulator was properly fitted and aligned, and the
participant felt comfortable walking back and forth along the walkway with minimal support
from the parallel bars. One trial was defined as walking one length of the parallel bars (4m) in
one direction, at the participant’s self-selected walking speed. Participants were asked not to rely
on the parallel bars for weight-bearing or support while completing data collection trials but were
encouraged to keep their hands on the bars while walking in the event of lost balance. In addition
to trials with the original, neutral alignment, trials were done with each alternate alignment
condition (Table 3). Alignment conditions included ±1cm, and ±2 cm translations in the sagittal
plane, facilitated by the custom adapter plates which were built into each prosthesis (Figure 11).
Translational adjustments are common in the process of dynamic prosthetic alignment and did
not require an acclimatization period. The testing order of all alignment conditions was randomly
selected by each participant prior to completing the trials to avoid systematic error. Data were
collected in the order shown in Table 4. Each participant completed four data collection trials
under each alignment condition. Able-bodied participant data collection followed the same
protocol as for the AK amputee participant (Figure 14).
Table 4: Order of conditions tested by able-bodied participants.
Participant NEUT POST1 POST ANT1 ANT AB1 1 5 2 4 3
AB2 1 2 3 5 4
AB3 1 5 3 4 2
AB4 4 2 5 1 3
AB5 5 4 1 2 3
AB6 1 2 4 5 3
AB7 1 2 3 4 5
AB8 1 2 4 5 3
30
2.3 Data Analysis
Features of the collected sensor data were identified and compared temporally in order to achieve
Objectives 2, 3, and 4. This section describes the synchronization of ASPL-SS data with those
recorded by the load transducer, the temporal relationships analyzed, and the thresholds chosen
to define data features for comparison. The sensor data sign convention used throughout this
report is summarized in Table 5.
Feature-defining threshold values for all analyses were selected using a trial and error approach
to identify the relevant events as specifically and sensitively as possible. Specific events for
temporal comparison were then identified and extracted manually from ASPL-SS and load
transducer data.
From each able-bodied participant walking trial recorded, only the second and third complete
gait cycles were analyzed to ensure steady-state gait. Similarly, the second, third, and fourth
complete gait cycles from each above-knee amputee waking trial were analyzed. Data from each
analyzed cycle were isolated and used as an individual set of points for comparison. Analysis
began with a total of 219 trials and 497 gait cycles.
All statistical analysis was performed using Microsoft Excel.
Table 5: Sensor data sign convention.
Positive (+) Negative (-)
Inductive Proximity Sensor Backward/Disengaged (away
from front of A-T Knee)
Forward/Engaged (toward
front of A-T Knee)
Force-Sensing Resistor *≥1: joint hyperextension *<1: natural joint flexion
Accelerometer
Vertical Upward acceleration/
Downward deceleration
Upward deceleration/
Downward acceleration
Horizontal Forward deceleration/
Backward acceleration
Forward acceleration/
Backward deceleration
Load
Transducer
Force Tension Compression
Moment Extension Flexion
31
2.3.1 Data Synchronization
Load transducer and ASPL-SS data were synchronized by a custom MATLAB function
(Appendix A.2) based on vertical force and acceleration data features at the start of each data
collection trial (i.e. stomp data) (Figure 15). Changes of at least 3000N/s in vertical force and
40g/s in vertical acceleration were used to define stomp start features and align the time stamps
for load transducer and ASPL-SS data. Some trials could not be synchronized confidently due to
gaps in data collection at the time of the stomp (i.e. > 0.01s between samples). These trials were
omitted from analysis (n=12).
Figure 15: Vertical force and acceleration stomp feature synchronization. Stomp start features were set at
0 seconds and subsequent timestamps were derived relative to them.
32
2.3.2 Engineering Validation
The second objective of this study was to verify that each of the ASPL-SS sensors met the design
goals in order to validate the developed system for use in achieving Objectives 3 and 4. To do so,
both bench test and walking trial (2.2.2) data were used to verify manufacturer specifications and
individual sensor calibration. Sensing system data were compared to those from a portable force
and torque transducer (2.1.2) and between individual sensors to test signal synchronization, and
that the ASPL-SS measurements were consistent with known relationships between
displacement, force, and acceleration. Data from both amputee and able-bodied participants were
used in the analysis of Objective 2 to confirm that the synchronization and function of the
sensors were comparable among participants and conditions. Table 6 provides an overview.
33
Table 6: Objective 2 test and analysis overview.
Parameter Test Analysis Acceptable Limits
Ind
uct
ive
Pro
xim
ity
Sen
sor
Precision; accuracy;
range; resolution
Bench Proximity measurements
compared to the calibrated
digital readout of a
machining table mill. [43]
Force/displacement
relationship; confirm
relevance of IPS
measurements
Bench Load applied to a linear
spring in contact with load
transducer. Spring
displacement measured by
IPS. [44]
Linear relationship between
measured force and
displacement.
Forc
e-S
ensi
ng
Res
isto
r
Sensitivity (i.e.
detection of small
applied loads at
terminal impact)
Walking Yes/No terminal impact
detection by FSR. Calculate
incidence of events
detected as a percentage of
total number of gait cycles.
Incidence of detection will
be compared among all AB
and AK conditions.
(2.3.2.1)
FSR values ≥ 1 will be
considered Yes. Suitability
of the sensor will be
determined by the overall
ratio although some
conditions may have a
higher incidence of
detection due to factors
outside of FSR sensitivity.
Acc
eler
om
eter
Calibration
Walking Confirm that vertical
acceleration ~ 1g and
horizontal acceleration ~ 0g
during stance-phase. The
mean value of
accelerometer readings
between initial stomp and
first gait cycle of each
NEUT trial will be
calculated. (2.3.2.2)
Values ± 0.25g will be
considered acceptable.
Identification of
characteristic gait
events
Walking Temporal offset (%GC)
will be calculated between
the selected ACC and load
transducer (FZ) data
features for heel strike and
toe-off events. For gait
cycles with FSR terminal
impact event detection,
temporal offset (%GC) will
be calculated between FSR
and ACC data features for
TI. Average temporal
offsets for each gait event
will be compared
statistically between AB
conditions. An analysis of
spatiotemporal parameters
(eg. stance time) will be
done and compared to
normative values. (2.3.2.2)
Temporal offset should not
be statistically significantly
different between
conditions and gait events
should occur at a clinically
appropriate point in the gait
cycle. The detection of gait
events may be more
challenging under certain
conditions.
34
Force-Sensing Resistor
The sensitivity of the force-sensing resistor was verified by checking for the detection of
terminal impact in late swing-phase of each gait cycle. By definition, terminal impact is the
instant at which the knee body contacts the thigh element, abruptly extending the limb in
preparation for weight-acceptance. This impact should register a contact force detectable by the
FSR and a corresponding feature in acceleration data measured by the accelerometer. Since
terminal impact occurs as an impulse in swing-phase, the contact force applied is substantially
less than that throughout weight-bearing. For this reason, terminal impact was used to verify the
sensitivity of the sensor. At baseline, without the application of any force, the FSR registered a
value of ~0.9 on a scale of 0-100. Increases in contact force, occurring in late swing-phase and
reaching impulse values ≥1 were considered terminal impact events. The overall ratio of events
detected to total number of gait cycles was calculated as a measure of sensor sensitivity across
conditions. Similar ratios were compared by condition and amongst participants to verify the
performance of the sensor even under conditions where the shank and foot may have experienced
less inertial force and smaller terminal impact forces, such as slow gait. Results of this analysis
are discussed in section 3.3.2.
35
Accelerometer
The accelerometer was calibrated prior to use in walking trials by placing it on a level surface
and setting the raw values (i.e. 0-1023) to the values listed in Table 7 in each orientation
(Appendix A.1). In order to assess this calibration, it was confirmed that vertical and horizontal
acceleration measured approximately 1g and 0g, respectively, for above-knee amputee and able-
bodied participants standing on level ground in the neutral alignment. The mean of accelerometer
values measured during the stationary period between the initial stomp and first gait cycle of
each NEUT condition trial was calculated for comparison (Figure 16). Values within ±0.25g
were considered acceptable given the curved profile of the All-Terrain Knee surface and
imperfect adhesion with the Velcro pads.
Table 7: Accelerometer calibration values.
Stationary Position Accelerometer axes (signed integer assignment)
AX AY AZ
Z down 0 0 -1 g
Z up 0 0 +1 g
Y down 0 -1 g 0
Y up 0 +1 g 0
X down -1 g 0 0
X up +1 g 0 0
Synchronization stomp Beginning of gait
Figure 16: Pale blue shading indicates the series of vertical and horizontal acceleration points between
initial stomp and first gait cycle used to calculate average accelerometer values while stationary on level
ground in the NEUT condition.
Vertical (z) acceleration
Horizontal (x) acceleration
Vertical (z) force
36
To verify the accurate detection of heel strike and toe-off by the accelerometer, acceleration data
features were compared temporally to those present in vertical force data. For terminal impact
detection, accelerometer data features were compared instead to force-sensing resistor data
features. Temporal offsets were evaluated separately for each gait cycle event to determine if the
selected thresholds identified the correct feature to represent each event, and whether or not the
accelerometer was equally effective at detecting heel strike, toe-off, and terminal impact.
Temporal offsets were also compared amongst participants and conditions under the assumption
that the ASPL-SS would detect features similarly, barring difficulty locating the defining
thresholds due to challenging terrain or undesirable gait characteristics (eg. toe-dragging) by less
experienced All-Terrain Knee users. The data features for comparison are shown in Figure 17
and were selected as follows. Their defining thresholds are listed in Table 8.
Heel Strike: A rapid upward acceleration reflects the inertial effect of ground contact and
should coincide with an increase in compressive force measured by the load transducer as the
limb accepts the weight of the body. The vertical force threshold is greater for the AK
participant since, on average, terminal impact occurs later in swing-phase (AK: 89.5±1.3
%GC, AB: 78.6±2.6 %GC), and vertical force immediately preceding heel strike is more
variable than for AB participants.
Toe-Off: Features surrounding toe-off were less distinct, and more challenging to identify
than those for heel strike and terminal impact. The following features were selected based on
biomechanical relevance and their temporal similarity to normative toe-off time (i.e.
~60%GC) however, future work should confirm this selection using gold standard force-
plates. Alternate features may need to be identified for participants exhibiting different gait
deviations. Forward acceleration reaches a peak as the foot leaves the ground and the user
initiates swing-phase, drawing the ipsilateral limb forward to progress gait. When the foot
leaves the ground, vertical force returns to a state of tension as the load transducer measures
the weight of the shank and foot components below it.
Terminal Impact: As the knee body makes contact with the thigh component, extending the
leg in preparation for weight-acceptance, the forward progression of the knee ends and inertia
causes a rapid forward deceleration while a force is applied to the FSR.
Results of this analysis are discussed in section 3.3.3.
37
Table 8: Thresholds selected to define accelerometer and vertical force data features for temporal
comparison (Figure 17). Increases/decreases in acceleration and force were measured between
consecutive data points (collected at ~100Hz).
Gait Event Acceleration Vertical Force
Heel Strike
(HS)
≥0.5g increase in vertical
acceleration
AB: ≥1N AK: ≥5N decrease in vertical
force
Toe-Off (TO) Negative peak in horizontal
acceleration Positive peak in vertical force
Terminal
Impact (TI)
≥0.5g increase in horizontal
acceleration
FSR reading ≥ 1
Figure 17: Acceleration, vertical force, and contact force data from one able-bodied participant gait cycle
in the NEUT condition. Features defined in Table 8 are indicated by the corresponding number and
colour.
TO HS TI
Acc
eler
atio
n (
g)
Forc
e (N
)
Vertical (z) acceleration
Horizontal (x) acceleration
Vertical (z) force
Contact Force Vertical (z) Force
Forc
e (N
)
Co
nta
ct F
orc
e
38
2.3.3 Measuring ASPL Function
The third objective of this study was to describe the relationships between moments applied at
the control and knee axes of the All-Terrain Knee and the resultant knee lock position and joint
flexion. While, theoretically, the function of the All-Terrain Knee and ASPL mechanism is well-
established, there has been a limited amount of research into quantifying the inner workings of
the device to confirm that it functions as intended. Objective 2 was to evaluate the ability of the
developed sensing system and protocol to detect known relationships so that they could
confidently be used to explore the relationships described in Table 9 for Objective 3.
Able-bodied participant data were used to accomplish Objective 3. Due to their inexperience
using a prosthetic simulator and All-Terrain Knee, able-bodied participants were likened to
above-knee amputees with weak gait or new to using the device, and since there were multiple
able-bodied participants with a comparable amount of experience, their results were more
generalizable than those of the above-knee amputee. The above-knee amputee participant data
were excluded from this portion of analysis; his experience and practice using the All-Terrain
Knee may have introduced confounding effects to the function of the ASPL mechanism.
Table 9: Objective 3 test and analysis overview.
Relationship Test Analysis (AB data)
Ind
uct
ive
Pro
xim
ity
Sen
sor
Knee lock
movement (IPS)
with respect to
changes in control
axis moment (load
transducer)
Walking Temporal offset (%GC) will be calculated by taking
the temporal difference between control axis
moment features and knee lock transitions in early
and mid-stance, and dividing it by 100% GC.
(2.3.3.1)
Forc
e-
Sen
sin
g
Res
isto
r Knee extension
(FSR) with respect
to changes in knee
axis moment (load
transducer)
Walking Temporal offsets (%GC) will be calculated by
taking the temporal difference between knee axis
moment features and knee extension-flexion
transitions in early and mid-stance, and dividing it
by 100% GC. (2.3.3.2)
Kn
ee L
ock
an
d
Kn
ee E
xte
nsi
on
Rel
ati
on
ship
Joint stability:
comparison of
knee extension
(FSR) and knee
lock behaviour
(IPS)
Walking Temporal offsets (%GC) will be calculated by
taking the temporal difference between extension-
flexion transitions at the control and knee axes in
mid- and late stance, as well as changes in lock
status and knee extension in mid- and late stance,
and following TI, and dividing it by 100% GC. Lock
position at the time of heel strike will also be
analyzed and compared to FSR data. (2.3.3.3)
39
Lock Displacement and Control Axis Moment
The inductive proximity sensor was used to detect and measure changes in lock position
throughout the gait cycle. In an effort to relate the function of the knee lock to applied control
axis moment measured by the load transducer, lock motion was compared temporally to moment
features at heel strike and mid-stance. Temporal offsets were divided by the time it took to
complete the gait cycle and represented as % gait cycle values for comparison amongst
conditions (Objective 4). The data features for comparison are shown in Figure 18 and were
selected as follows. Their defining thresholds are listed in Table 10.
Heel Strike: As described in section 1.2.4.1, the ASPL mechanism was designed such that
an early stance-phase flexion moment at the control axis should secure knee lock engagement
(i.e. drive the knee lock forward), ensuring full knee extension throughout weight-
acceptance.
Mid-Stance: Forefoot loading naturally creates an external extension moment about the
control axis; this moment should drive the knee lock backward, disengaging it.
Results of this analysis are discussed in section 3.4.1.
40
Table 10: Thresholds selected to define lock displacement and control axis moment data features for
temporal comparison (Figure 18). Increases/decreases in lock position and moment were measured
between consecutive data points (collected at ~100Hz).
Gait Event Lock Position (IPS) Control Axis Moment
Heel Strike (HS) ≥0.05mm forward displacement ≥0.1Nm decrease in moment
Mid-Stance (MS) ≥0.05mm backward
displacement
Flexion (-) to extension (+)
transition
HS MS
Figure 18: Lock position and control axis moment data from one able-bodied participant gait cycle in the
NEUT condition. Negative lock position values indicate forward displacement/lock engagement and
positive values indicate backward displacement/lock disengagement. Negative and positive moment
values indicate flexion and extension, respectively. Features defined in Table 10 are indicated by the
corresponding number and colour.
Lock Position Control Axis Moment
Vertical (z) force (not to scale)
TO TI
41
Knee Extension and Knee Axis Moment
Force-sensing resistor data were used to calculate temporal offsets between knee joint extension
and knee axis moment events measured by the load transducer (Table 11). Temporal offsets
were divided by the time it took to complete the gait cycle and represented as % gait cycle values
for comparison amongst conditions (Objective 4). The data features for comparison are shown in
Figure 19 and were selected as follows. Their defining thresholds are listed in Table 11.
Heel Strike: As described in section 1.2.4.1, an external flexion moment is applied to the
knee during the loading response, tending to flex the joint. This early stance-phase flexion
separates the thigh component and knee body at the contact interface, relieving the small
contact force previously applied to the force-sensing resistor following terminal impact when
the components came into contact.
Mid-Stance: When an external extension moment is applied at the knee axis, the knee
should become fully extended naturally (i.e. without the support of the knee lock) and apply
contact force to the force-sensing resistor.
Results of this analysis are discussed in section 3.4.2.
42
Table 11: Thresholds selected to define contact force and knee axis moment features for temporal
comparison (Figure 19). Increases/decreases in moment were measured between consecutive data points
(collected at ~100Hz).
Gait Event Contact Force (FSR) Knee Axis Moment
Heel Strike (HS) Contact force < 1 ≥0.1Nm decrease in moment
Mid-Stance (MS) Contact force ≥1 Flexion (-) to extension (+)
transition
Contact Force Knee Axis Moment
Vertical (z) force (not to scale)
HS MS
Figure 19: Contact force and knee axis moment data from one able-bodied participant gait cycle in the
NEUT condition. Negative and positive moment values indicate flexion and extension, respectively.
Features defined in Table 11 are indicated by the corresponding number and colour. Inset shows a
zoomed in view of the force and moment features at heel strike.
TO TI
43
Knee Stability
By comparing joint extension and knee lock behaviour, we can begin to understand the role
played by the knee lock in maintaining stability, determine when knee joint extension is naturally
sustained by an external extension moment, and ensure the lock is in place to provide stability
when appropriate. Temporal offsets between joint extension and knee lock events were divided
by the time it took to complete the gait cycle and represented as % gait cycle values for
comparison amongst conditions (Objective 4). The data features for comparison are shown in
Figure 20 and were selected as follows. Their defining thresholds are listed in Table 12.
Terminal Impact/Heel Strike: Since external flexion moment at the knee axis tends to
cause knee joint flexion at heel strike, the knee lock must be engaged to prevent buckling and
maintain knee stability. To confirm that the knee lock was engaged when the thigh
component and knee body were drawn apart at heel strike, the time of lock engagement
preceding heel strike was compared to that of contact force loss (see inset of Figure 20).
Mid-Stance: At mid-stance, as the body moves from weight-acceptance to propulsion, the
ground reaction force vector passes anterior to the control and knee axes for a time,
generating an external extension moment at each. To determine whether the knee became
hyperextended prior to lock disengagement, ensuring knee stability, the application of
contact force and lock transition from engaged to disengaged were compared temporally.
Toe-Off: To confirm that the knee lock was disengaged in time to permit knee flexion in
preparation for toe-off, the removal of contact force was compared temporally to the forward
displacement of the lock in terminal stance.
The control and knee axis moment transitions between flexion and extension were also
compared temporally at mid-stance and toe-off.
Results of this analysis are discussed in section 3.4.3.
44
Table 12: Thresholds selected to define features for temporal comparison of knee stabilizing events
(Figure 20).
Gait Event Lock Position (IPS) Contact Force (FSR)
Terminal Impact (TI)/
Heel Strike (HS)
Disengaged (+) to engaged
(-) transition Contact force < 1
Mid-Stance (MS) Engaged (-) to disengaged
(+) transition Contact force ≥1
Toe-Off (TO) Disengaged (+) to engaged
(-) transition Contact force < 1
Control Axis Moment Knee Axis Moment
Mid-Stance (MS) Flexion (-) to extension (+)
transition
Flexion (-) to extension (+)
transition
Toe-Off (TO) Extension (+) to flexion (-)
transition
Extension (+) to flexion (-)
transition
45
Figure 20: Lock position, contact force, knee and control axis moment data from one able-bodied
participant gait cycle in the NEUT condition. Negative lock position values indicate forward
displacement/lock engagement and positive values indicate backward displacement/lock disengagement.
Negative and positive moment values indicate flexion and extension, respectively. Features defined in
Table 12 are indicated by the corresponding number and colour. Inset shows a zoomed in view of the
contact force event at heel strike.
HS TI TO MS HS TI
Contact Force Knee Axis Moment
Vertical (z) force (not to scale)
Lock Position Control Axis Moment
Vertical (z) force (not to scale)
46
2.3.4 Comparing Conditions
The fourth objective of this study was to use collected data to confirm that the ASPL-SS can
detect differences in ASPL function between a variety of alignment, gait, terrain, and All-Terrain
Knee setting conditions, and to conduct a preliminary exploration of the effects of these
conditions on the function of the ASPL mechanism.
In order to examine the effects of variable conditions on the results of Objective 3, as well as on
lock engagement duration, walking trial data were compared between able-bodied and above-
knee amputee participant groups where possible, and amongst conditions (Table 13). The tested
conditions and their theoretical effects are described in Table 3 and Table 14, respectively.
Table 15 defines the features analyzed for the comparison of lock engagement duration and
shown in Figure 21.
Results of these analyses are discussed throughout section 3.4.
Table 13: Objective 4 analysis overview.
Related
Objective Relationship
Gait
Events
Conditions Compared
(Table 14) Statistical Analysis
3
Lock Displacement
and Control Axis
Moment
HS, MS AB: NEUT, ANT, POST
Single-factor ANOVA
will be performed between
neutral (NEUT), anterior
foot translation (ANT),
and posterior foot
translation (POST)
conditions across able-
bodied participants. Paired
t-tests will be performed
where statistically
significant differences are
found. No statistical
analysis will be performed
on AK results since data
were collected from only
one individual. Qualitative
comparisons will be done
for the AK relationships
listed here.
3 Knee Extension and
Knee Axis Moment HS, MS AB: NEUT, ANT, POST
3
Knee Stability:
Lock Displacement
and Knee Extension
MS, TO AB: NEUT, ANT, POST
Lock Engagement
Duration (% stance-
phase)
HS-MS
AB: NEUT, ANT, POST
AK: NEUT, ANT, POST,
PFLX, DFLX, TTLS,
LSLS, INC, DEC, FST,
SLW, LZY, RCK, GRS
47
Table 14: Test conditions, their clinical relevance, and theoretically expected effects. Clinical relevance Condition Expected effect
Baseline for comparison NEUT
Pro
sth
etic
ali
gnm
ent
The relative alignment of prosthetic
components plays a role in controlling
stance-phase stability. Understanding
how alignment affects the function of
the All-Terrain Knee will inform both
manufacturer-recommended settings
and potential design changes. By
communicating the function of the
ASPL following alignment changes to
clinicians in real time, the ASPL-SS
will facilitate efficient and proper
alignment for the comfort and safety
of each patient.
ANT
Shifting the foot anteriorly with respect to
the knee would be expected to generate
external extension moments at the control
and knee axes sooner in stance-phase,
hyperextending the joint naturally and
driving the knee lock to the disengaged
position. A posterior shift should make
the knee more dependent on the knee lock
for stability and keep the lock engaged for
more of stance-phase.
POST
PFLX
Like an anterior translation of the foot,
plantar flexion would be expected to
generate external extension moments at
the control and knee axes earlier in
stance-phase, stabilizing the knee
naturally and disengaging the knee lock
sooner. The opposite should be true for
dorsiflexion.
DFLX
All
-Ter
rain
Knee
set
tings
The All-Terrain Knee has a number of
adjustable features to accommodate
the mobility needs of different users.
While this study only explored the
function of altering lock spring
tightness, future large-scale trials
exploring more conditions would also
inform manufacturer-recommended
settings and facilitate effective patient-
specific setup of the joint.
TTLS
Tightening the lock spring will increase
the force it exerts on the knee lock,
thereby driving the lock forward and
biasing it toward the engaged position.
The joint should remain locked longer
with less applied external flexion moment
at the control axis. Maintaining lock
engagement ensures knee extension rather
than depending solely on naturally
applied external extension moment at the
knee axis in late stance. This may also
make knee flexion more challenging
preceding toe-off. When the lock spring is
loosened, one would expect more sudden
lock movements largely dependent on
changes in externally applied moment,
and for the knee lock to be in the neutral,
unlocked position for more of the gait
cycle.
LSLS
SFSS
The addition of the stance-phase flexion
mechanism would be expected to increase
flexion moments acting on the joint
keeping the prosthesis locked and
extended while still permitting a small
amount of flexion in stance-phase as the
name of the mechanism suggests.
48
Wal
kin
g s
pee
d/c
on
fid
ence
It is important to understand the effect
of walking speed on the function of
the joint from both the research and
clinical perspectives. If the joint
performs poorly under certain
conditions, design changes may be
necessary, or it may be possible to
apply alignment or knee setting
changes to improve mechanism
function for patients expressing
certain characteristic gait patterns or
wishing to use the prosthesis for
activities such as running.
FST
If a strong ambulator is walking quickly,
the ASPL should not impede flexion at
the initiation of swing-phase. For those
with less experience or weaker
musculature, fast gait may pose a larger
physical challenge, demanding more
support from the knee lock in stance-
phase, this was not tested in this study.
SLW
In slow gait, more time is spent in each
phase of the gait cycle, potentially
resulting in more exaggerated lock
motion.
LZY
The ‘lazy’ gait condition was intended to
mimic the gait of weak ambulators who
would depend heavily on stance-phase
control to maintain extension while
weight-bearing. In this condition, the lock
should remain securely engaged as long
as the knee axis is experiencing an
external flexion moment. This condition
should produce results similar to those of
the able-bodied participants.
Ter
rain
/Mobil
ity c
ondit
ion
Similar to walking speed, exploring
the effects of terrain and mobility
conditions on ASPL function will help
define recommended alignment and
knee setting changes for users in
diverse environments
RCK Rough terrains are expected to introduce
variability within trials. The ASPL
mechanism should be sensitive and
reactive to these changes. GRS
INC
Like anterior foot translation and
plantarflexion, walking on an incline
would be expected to generate an external
extension moment at the control axis and
unlock the mechanism sooner. The
opposite is true for descending a slope,
which is theoretically equivalent
dorsiflexion.
DEC
49
Table 15: Thresholds selected to define data features for comparison of lock engagement duration
(Figure 21). Increases/decreases in lock position and force were measured between consecutive data
points (collected at ~100Hz).
Beginning End
Lock Engagement
Duration (%
Stance-Phase)
≥0.05mm forward lock
displacement
≥0.05mm backward lock
displacement
Figure 21: Lock position and vertical force data from one able-bodied participant gait cycle in the NEUT
condition. Negative lock position values indicate forward displacement/lock engagement and positive
values indicate backward displacement/lock disengagement. Features defined in Table 15 are indicated
by the corresponding number and colour.
Lock Position Vertical (z) force
HS MS TO
TI
50
Results and Discussion
This section present and interprets the results of the data analysis described in section 2.3, their
relevance to the field, and their clinical implications.
3.1 Sensitivity Analysis
Toe-dragging in early swing-phase was common among able-bodied participants (63% of AB
gait cycles analyzed). This additional and irregular contact between prosthetic foot and ground
had resultant effects on applied force and ASPL events recorded by the ASPL-SS. The features
used to identify toe-drag cycles were selected based on their temporal relationship to terminal
stance-phase and their absence from any above-knee amputee participant gait cycles since lack
of toe-drag was confirmed visually at the time of AK data collection (Figure 22). In order to
determine if and how these additional forces affected the spatiotemporal parameters of able-
bodied participant gait and the ability to identify relevant gait events from sensor data, a
sensitivity analysis was performed to statistically compare the following parameters between gait
cycles with and without toe-drag in the NEUT condition (Table 16) .
Stride Time (s): The length of time between two consecutive heel strikes of the same foot,
i.e. one gait cycle, was compared between drag and no-drag cycles since most temporal
parameters were reported relative to the duration of a gait cycle.
Stance Time/Offset (%GC): The end of stance-phase was defined as a peak in vertical force
following a sustained period of compression measured by the load transducer or a negative
peak in horizontal acceleration measured by the accelerometer. Comparing stance time
between drag and no-drag cycles should indicate whether or not toe-drag influenced the
detection of these features by either sensor.
Terminal Impact Time/Offset (%GC): Since toe-drag occurs during swing-phase, terminal
impact timing measured by the accelerometer and force-sensing resistor was compared
between drag and no-drag cycles to determine if the progression of swing-phase was affected
temporally by toe-drag and whether or not toe-drag influenced the detection of terminal
impact features by either sensor.
51
Figure 22: Data from one above-knee amputee gait cycle (top), one able-bodied gait cycle with no toe-
drag (middle), and one able-bodied gait cycle with toe drag (bottom). Graphs on the left depict lock
position and control axis moment, graphs on the right depict contact force and knee axis moment. Shaded
grey regions highlight the period between toe-off and terminal impact events. Negative peaks in lock
position and moment which occur between toe-off and terminal impact (indicated by red arrows) are not
present in above-knee amputee gait cycles and suggest that there was an external force applied to the limb
during swing-phase.
Lock position and control axis moment
Lock Position Control Axis Moment Vertical (z) force (not to scale)
Contact Force Knee Axis Moment Vertical (z) force (not to scale)
AK
A
B n
o-d
rag
AB
dra
g Contact force and knee axis moment
52
Table 16: All able-bodied participant1 gait cycles in the NEUT condition were identified as either drag or
no-drag based on the presence or absence of the toe-drag features described above. The results of paired t-
tests used to compare drag and no-drag cycles for each parameter are shown below. Statistically
significant differences (p<0.05) are indicated by *.
Mean SD p-value
Stride Time (s) Drag 1.88 0.22
0.52 No drag 1.86 0.28
Stance Time (FZ) (%GC) Drag 62.00 3.37
0.20 No drag 63.99 3.07
Stance Time (ACC) (%GC) Drag 59.12 3.35
0.39 No drag 60.40 3.72
Toe-Off Offset (%GC) Drag -2.88 1.26
0.27 No drag -3.60 2.27
Terminal Impact Time (FSR) (%GC) Drag 79.04 2.67
0.004* No drag 75.95 2.47
Terminal Impact Time (ACC) (%GC) Drag 79.38 2.74
0.004* No drag 76.29 2.39
Terminal Impact Offset (%GC) Drag 0.35 0.19
0.93 No drag 0.34 0.23
The only statistically significant difference detected was in terminal impact time measured by
both the force-sensing resistor and accelerometer. These results suggest that toe-dragging slowed
the progression of the leg through swing-phase thereby delaying terminal impact in gait cycles
where toe-drag occurred. Stride and stance time were not significantly affected by toe-drag, nor
were the temporal offsets measured between load transducer and accelerometer events for toe-off
and between FSR and accelerometer at terminal impact. Based on these results, both drag and
no-drag cycles were included for the remainder of the analyses.
1 AB4 and AB8 were excluded from this comparison. Both participants exhibited toe-drag in each of the analyzed
NEUT gait cycles, thus drag values could not be paired for statistical comparison with no-drag.
53
3.2 Spatiotemporal Parameters
Participant stride and stance time were compared to values reported in the literature for
normative, above-knee amputee, and prosthetic simulator gait to give the collected data clinical
context and relevance (Table 17). Average stride time for the able-bodied participants was
abnormally long compared to literature values, including the mean stride time reported by
Lemaire, et al. for prosthetic simulator gait without the use of a cane. This may be attributable to
the limited simulator-walking experience of able-bodied participants in this study, who received
one 45-60 minute gait training session prior to data collection while Lemaire’s received two [40].
The average stride time for the above-knee amputee participant was within the normative range,
and shorter than the literature average for above-knee amputees. In healthy gait, stance-phase
usually lasts about 60% of the gait cycle, as was observed in both able-bodied and above-knee
amputee participant data. This also supports the selection of the negative peak in horizontal
acceleration as the identifying feature for toe-off and the end of stance-phase in the subsequent
analyses.
Table 17: A comparison of measured and literature averages for stride time (seconds) and stance time (%
gait cycle) in the NEUT condition.
Stride Time (s) Stance Time (ACC) (%GC)
AB AK AB AK
Mea
n Experimental 1.93 1.13 60.41 59.19
Normative 0.87-1.32 [34] 60 [34]
Above-Knee Amputee 1.38 [40] 57.25 [45]
Prosthetic Simulator 1.56 [40] Reference not found
54
3.3 Engineering Validation
3.3.1 Inductive Proximity Sensor: Bench Tests
Prior to beginning data collection trials with the full sensing system, preliminary bench
experiments were carried out to quantitatively verify the calibration curves provided by the
manufacturer of the inductive proximity sensor and assess its performance in a simulation of its
final application. Results showed adequate resolution, a high degree of accuracy and precision in
sensor output, as well as a linear position-voltage correlation for the applicable range of knee
lock positions [43].
In a second bench test, the ability to record and relate data from the load transducer and inductive
proximity sensor was assessed. Force was applied to a spring, which in turn applied a
compressive force to the load transducer in the vertical direction. The inductive proximity sensor
simultaneously measured the distance the spring was compressed. As expected, the results
showed a linear relationship between displacement and force applied (force applied = spring
constant*displacement).
55
3.3.2 Force-Sensing Resistor
A distinct terminal impact event (i.e. FSR value transition from <1 to ≥1) was detectable by the
FSR in almost every gait cycle (97.3%). The one exception was amputee gait up a 10° ramp,
likely due to coincidental terminal impact and heel strike on an inclined surface (Table 18).
These results indicate that the FSR is sensitive enough to detect small applied loads, however
inconsistencies in FSR signal magnitude between gait cycles may reflect issues with sensor
fixation at the contact interface (Figure 23).
Table 18: Terminal impact events detected as a percentage of total gait cycles analyzed, listed by
condition.
Condition AK (%) AB (%)
NEUT Neutral 100 98.3
ANT Anterior Foot Translation 100 100
POST Posterior Foot Translation 100 100
PFLX Plantarflexion 100
DFLX Dorsiflexion 100
TTLS Tight lock spring 100
LSLS Loose lock spring 100
FST Fast 100
SLW Slow 100
LZY Lazy gait 100
RCK Rocks 100
GRS Grass 100
INC Incline 0
DEC Decline 100
SFSS Stance-Phase Flexion 100
TI HS TO HS TO TI
Contact Force Vertical (z) force
Figure 23: An example of two consecutive above-knee amputee gait cycles under the same condition
(fast gait) with significantly different contact forces at terminal impact (red circles).
56
3.3.3 Accelerometer
Accelerometer calibration was confirmed as described in section 2.3.2.2. The values for the
above-knee amputee participant deviated by up to ±0.25g from the desired values, likely due to
poor accelerometer fixation which was rectified before beginning AB trials. Average values
across able-bodied participants were within ±0.05g of the desired values (Table 19).
The results of temporal offset analysis for gait event detection are shown in Table 20. For both
able-bodied and above-knee amputee participants in the NEUT condition, all mean temporal
offsets between gait cycle events measured by different sensors were within 4% gait cycle; with
the exception of AB toe-off detection, values were well-within 2%. Past studies evaluating
inertial sensors against normative and foot switch values for the detection of gait events accepted
similar results [46-47].
While mean toe-off temporal offset was less than 1% for the above-knee amputee, able-bodied
participant results exceeded 3% gait cycle. Excluding able-bodied toe-dragging trials did not
significantly change this outcome (Table 16). Possible explanations for the larger offsets in AB
toe-off event data include variable mechanics of the prosthetic simulator compared to a true,
well-fitting prosthesis and challenges initiating flexion or swing-phase for inexperienced users,
creating variable and less natural acceleration and vertical force profiles. The defining data
features of toe-off were also less distinct than those present at heel strike and terminal impact,
especially in AB data (Figure 17). This may have produced some uncertainty in toe-off
detection.
Based on a repeated measures ANOVA, changes in translational alignment did not significantly
affect the detection of any gait events. Heel strike: F (2, 14) = 3.07, p = 0.08; toe-off: F (2, 14) =
3.73, p = 0.05; terminal impact: F (2, 14) = 0.16, p = 0.85 (Table 21). These results suggest that
the selected gait event data features are present, and that the ASPL-SS is capable of detecting
them, under variable conditions.
Table 19: Average horizontal and vertical acceleration values in stationary NEUT condition.
AB AK
Horizontal Acceleration 0.952g 0.887g
Vertical Acceleration 0.012g -0.250g
57
Table 20: Average temporal offset in gait event detection by the accelerometer and load transducer (heel
strike and toe-off) and accelerometer and force-sensing resistor (terminal impact). Values represent gait
cycles in the NEUT condition and are shown as %GC. Negative values indicate that acceleration events
preceded vertical or contact force events.
Gait Event AB AK
Mean SD Mean SD
Heel Strike -0.37 0.15 1.01 0.76
Toe-Off -3.28 1.38 -0.30 0.41
Terminal Impact 0.35 0.17 -0.91 0.54
Table 21: Mean and standard deviation of temporal offset in gait event detection under different
translational alignment conditions for able-bodied participants. Values are shown as % gait cycle.
Negative values indicate that acceleration events preceded vertical and contact force events. Statistically
significant (p < 0.05) results of repeated measures ANOVA indicated by *.
Gait Event NEUT ANT POST
p-value Mean SD Mean SD Mean SD
Heel Strike -0.37 0.15 -0.42 0.21 -0.29 0.14 0.08
Toe-Off -3.28 1.38 -2.61 1.09 -2.91 0.97 0.05
Terminal Impact 0.35 0.17 0.38 0.13 0.36 0.17 0.85
58
3.4 ASPL Function
The results of engineering validation tests indicate that the inductive proximity sensor, force-
sensing resistor, and accelerometer measurements are valid and that experimental ASPL-SS
measurements can therefore be used to determine whether the empirical performance of the
ASPL mechanism is consistent with its theoretical function. This section describes the
relationships between moments applied at the control and knee axes of the All-Terrain Knee and
the resultant knee lock position and joint flexion, and offers a preliminary comparison of those
relationships, as well as an exploration of ASPL stance-phase control function, under relevant
variable conditions.
59
3.4.1 Lock Displacement and Control Axis Moment
The lock displacement events at heel strike and mid-stance described in section 2.3.3.1 occurred
on average within 2% gait cycle of the corresponding control axis moment events (Table 22). At
heel strike in the NEUT condition, average temporal offset (-0.42 %GC ~ 8ms) is less than 1%
gait cycle and within the sampling frequency of the ASPL-SS and load transducer (100Hz = 1
sample/10ms), therefore we can conclude that the forward displacement of the knee lock and
application of a flexion moment at the control axis were effectively simultaneous. At mid-stance,
however, the larger negative average offset (-1.28 %GC) implies that the backward displacement
of the lock begins before the transition from flexion to extension moment at the control axis,
likely corresponding more closely to an earlier change in moment magnitude instead.
Based on repeated measures ANOVA, the effect of translational alignment on temporal offsets of
the lock and moment events at heel strike is not statistically significant, F(2, 7) = 0.05, p = 0.95
(Table 22). A repeated measures ANOVA did indicate a statistically significant effect of
translational alignment on temporal offsets at mid-stance, F(2, 7) = 4.10, p = 0.04, however post
hoc comparisons using a paired t-test with Bonferroni correction for multiple (3) comparisons
indicated that none of the mean values were significantly different (p ≤ 0.017):
Neutral (M = -1.28, SD = 0.76) and anterior (M = -1.52, SD = 1.08) alignment; t(7) = 2.36, p
= 0.24.
Neutral (M = -1.28, SD = 0.76) and posterior (M = -1.06, SD = 0.83) alignment; t(7) = 2.36, p
= 0.05.
Anterior (M = -1.52, SD = 1.08) and posterior (M = -1.06, SD = 0.83) alignment; t(7) = 2.36,
p = 0.04.
Table 22: Average temporal offset between lock displacement and control axis moment events at heel
strike and mid-stance. Values are for able-bodied participant gait cycles in each condition and are shown
as %GC. Negative values indicate that lock displacement events preceded moment events. Statistically
significant (p < 0.05) results of repeated measures ANOVA indicated by *.
NEUT ANT POST p-value
Gait Event Mean SD Mean SD Mean SD
Heel strike -0.42 0.32 -0.42 0.31 -0.39 0.40 0.95
Mid-stance -1.28 0.76 -1.52 1.08 -1.06 0.83 0.04*
60
3.4.2 Knee Extension and Knee Axis Moment
The contact force events at heel strike and mid-stance described in 2.3.3.2 occurred on average
within 1% gait cycle of the corresponding knee axis moment events in the NEUT condition.
Based on repeated measures ANOVA, differences were not statistically significant (p < 0.05)
between the various translational alignment conditions (Table 23):
Heel strike: F(2, 4) = 1.81, p = 0.232
Mid-stance: F(2, 7) = 0.60, p = 0.56
These results suggest that the occurrence of natural knee joint flexion and extension is closely
related to the application of external flexion and extension moments at the knee axis, and that
weak or inexperienced users rely on knee lock engagement to maintain extension and stability in
early stance-phase.
Table 23: Average temporal offset between contact force and knee axis moment events at heel strike and
mid-stance. Values are for able-bodied participant gait cycles in each condition and are shown as %GC.
Negative values indicate that contact force events preceded moment events. Statistically significant (p <
0.05) results of repeated measures ANOVA indicated by *.
2 AB3, AB6, and AB8 were excluded from this comparison. These participants stopped applying contact force at the
FSR prior to heel strike, thus the relevant events could not be compared temporally between conditions.
NEUT ANT POST p-value
Gait Event Mean SD Mean SD Mean SD
Heel strike -0.96 0.55 -0.97 1.16 -0.40 0.17 0.23
Mid-stance 0.44 1.15 0.21 0.72 0.31 0.88 0.56
61
3.4.3 Knee Stability
In all but one of the able-bodied participant gait cycles analyzed, the knee joint became extended
at terminal impact and exerted a measurable force on the FSR (Table 18). In each of these
cycles, the knee lock then moved forward and remained engaged, quantitatively confirming that
joint extension was maintained throughout weight-acceptance, despite external flexion moments
at the knee axis and loss of contact force at the contact interface caused by the natural flexion
tendency during the loading response.
At mid-stance, the transition from weight-acceptance to propulsion caused external knee and
control axes moments to transition from flexion to extension in all able-bodied participant gait
cycles. In the neutral (NEUT) condition, the knee became hyperextended (i.e. contact force was
applied at the contact interface and detected by the force-sensing resistor) an average of 0.45%
gait cycle before the knee lock began to displace backward into the unlocked position, thereby
ensuring continuous stability. Similarly, the transition from flexion to extension moment at the
knee axis occurred an average of 2.1% gait cycle sooner than at the control axis. The temporal
offsets between knee hyperextension and lock displacement, and between knee and control axes
moment transitions were compared statistically between able-bodied conditions with a repeated
measures ANOVA as a measure of relative alignment stability (Table 24). The results of these
analyses indicated statistically significant (p < 0.05) differences:
Knee hyperextension/backward lock displacement: F(2, 7) = 5.05, p = 0.02
Knee axis/control axis flexion-to-extension transition: F(2, 7) = 5.51, p = 0.02
62
Table 24: Average temporal offset between contact force and knee lock, and between knee axis (KA) and
control axis (CA) moment events at mid-stance and toe-off. Values are for able-bodied participant gait
cycles in each condition and are shown as %GC. Negative values indicate that lock or control axis events
preceded contact force or knee moment events. Statistically significant (p < 0.05) results of repeated
measures ANOVA indicated by *.
NEUT ANT POST p-value
Gait Event Mean SD Mean SD Mean SD
Mid
-sta
nce
Knee hyperextension/
backward lock
displacement
0.45 2.25 0.71 2.71 -1.35 0.97 0.02*
KA/CA flexion-to-
extension 2.10 2.31 2.44 3.20 0.02 0.93 0.02*
Toe-
off
Natural knee flexion/
forward lock displacement 8.84 4.79 10.61 5.10 7.74 5.85 0.02*
KA/CA extension-to-
flexion 9.10 2.49 9.92 3.72 9.08 2.87 0.41
Post hoc paired t-tests with Bonferroni correction for multiple (3) comparisons showed no
statistically significant (p ≤ 0.017) differences between conditions for hyperextension and lock
disengagement offsets:
Neutral (M = 0.45, SD = 2.25) and anterior (M = 0.71, SD = 2.71) alignment; t(7) = 2.36, p =
0.60.
Neutral (M = 0.45, SD = 2.25) and posterior (M = -1.35, SD = 0.97) alignment; t(7) = 2.36, p
= 0.04.
Anterior (M = 0.71, SD = 2.71) and posterior (M = -1.35, SD = 0.97) alignment; t(7) = 2.36, p
= 0.05.
Similarly, post hoc paired t-tests with Bonferroni correction for multiple (3) comparisons showed
no statistically significant (p ≤ 0.017) differences between conditions for knee and control axis
moment offsets:
Neutral (M = 2.10, SD = 2.31) and anterior (M = 2.44, SD = 3.20) alignment; t(7) = 2.36, p =
0.61.
Neutral (M = 2.10, SD = 2.31) and posterior (M = 0.02, SD = 0.93) alignment; t(7) = 2.36, p =
0.02.
Anterior (M = 2.44, SD = 3.20) and posterior (M = 0.02, SD = 0.93) alignment; t(7) = 2.36, p
= 0.04.
63
Although there were no statistically significant differences to report, likely due to the amount of
variability between able-bodied participants, the results of the alignment comparisons did follow
a pattern. At mid-stance, the average temporal offset between knee hyperextension and backward
lock displacement in the posterior alignment condition (POST) is a negative value (-1.35 %GC)
which indicates that the knee lock actually began to disengage prior to knee hyperextension. This
implies that a posterior foot offset relative to the knee is more likely to become unstable in mid-
stance compared to the neutral condition. In the anterior alignment condition (ANT), average
knee lock disengagement occurs later relative to knee hyperextension when compared to the
neutral and posterior conditions, suggesting that, at mid-stance, this is the most stable
translational alignment of the three. The same pattern is present for knee and control axis
moment transitions from flexion to extension at mid-stance.
A similar analysis was done for terminal stance-phase to confirm that the knee lock did not
impede knee flexion at the initiation of swing-phase. In every gait cycle analyzed, contact force
was removed, indicating knee flexion, prior to forward lock displacement in terminal stance-
phase, suggesting that the knee lock did not interfere with the progression of gait. Knee axis
moment transition from extension to flexion also preceded control axis moment transition in
each analyzed gait cycle. These results were also compared statistically between able-bodied
conditions with a repeated measures ANOVA (Table 24). The results of these analyses indicated
statistically significant (p < 0.05) differences only in the knee flexion/lock displacement offset:
Knee flexion/forward lock displacement: F(2, 7) = 5.06, p = 0.02
Knee axis/control axis extension-to-flexion transition: F(2, 7) = 0.95, p = 0.41
Post hoc paired t-tests with Bonferroni correction for multiple (3) comparisons showed no
statistically significant (p ≤ 0.017) differences between conditions for knee flexion and lock
engagement offsets:
Neutral (M = 8.84, SD = 4.79) and anterior (M = 10.61, SD = 5.10) alignment; t(7) = 2.36, p
= 0.07.
Neutral (M = 8.84, SD = 4.79) and posterior (M = 7.74, SD = 5.85) alignment; t(7) = 2.36, p =
0.26.
Anterior (M = 10.61, SD = 5.10) and posterior (M = 7.74, SD = 5.85) alignment; t(7) = 2.36,
p = 0.02.
64
Unlike able-bodied participants, the above-knee amputee participant maintained contact force
throughout stance-phase in every condition analyzed except for slow (SLW) and lazy (LZY) gait,
indicating that he was not depending on the lock for knee stability in most conditions. Flexion
moments were rarely measured at either axis in above-knee amputee gait, therefore transitions
between flexion and extension were not available for analysis. For these reasons, above-knee
amputee participant data were not considered in the analysis of Objective 3 and were not
included in the above comparisons.
As discussed in section 1.2.3.2, stance-phase stability is affected by prosthetic alignment and
settings as well as by gait and terrain conditions. The duration of knee lock engagement under
different conditions can be used to compare the relative stance-phase control demands of
different conditions, and how manipulating alignment and All-Terrain Knee settings may help
address them.
A repeated measures ANOVA was conducted to compare the effect of foot translational
alignment on knee lock engagement duration for able-bodied participants. An analysis of
variance showed that the effect of prosthetic alignment on lock engagement duration was
significant (F(2, 7) = 34.42, p < 0.001). Post hoc paired t-tests with Bonferroni correction for
multiple (3) comparisons were conducted to compare lock engagement duration in each of the
alignment conditions (Table 25). Lock engagement duration in the anterior condition was
significantly less than in both the neutral and posterior conditions. The lock remained engaged
longest in the posterior alignment condition, and the least in the anterior condition. While the
results of the previous stability analyses (Table 24), show that anterior foot translation stabilizes
the ASPL mechanism by maintaining lock engagement throughout the mid-stance transition, the
present comparison of lock engagement duration illustrates the intrinsic stability of the anterior
foot translation alignment based on the large proportion of stance-phase spent in hyperextension,
rather than depending on the knee lock for stability.
Above-knee amputee data followed the same trend: ANT (0% stance-phase) < NEUT
(8.83±2.36% SP) < POST (35.52±2.40% SP). Average lock engagement durations are shown in
Figure 24.
65
Table 25: The results of paired t-tests used to compare knee lock engagement duration between each set
of conditions. Mean and standard deviation values are shown as % stance-phase. Statistically significant
(p ≤ 0.017) results of paired t-tests indicated by *.
Condition Mean SD df t p-value
NEUT 32.55 7.32 7 2.36 0.001*
ANT 21.80 5.83
NEUT 32.55 7.32 7 2.36 0.19
POST 34.74 7.71
ANT 21.80 5.83 7 2.36 <0.001*
POST 34.74 7.71
Figure 24: Average and standard deviation lock engagement duration for able-bodied participants (n=8)
under each condition (shown in black). Individual able-bodied participant means shown in grey. Above-
knee amputee means and standard deviations shown in dark grey and broken lines for comparison.
0
10
20
30
40
50
Lock
En
gage
men
t D
ura
tio
n (
% s
tan
ce-p
has
e)
Condition
Lock Engagement Duration
NEUT ANT POST
66
Statistical comparisons were not conducted on the single above-knee amputee participant’s data.
Lock engagement duration was compared among different conditions as a relative measure of the
stability requirement of each condition but again, it should be noted that the above-knee amputee
participant maintained contact force throughout stance-phase in every condition analyzed except
for slow (SLW) and lazy (LZY) gait, indicating that he was not depending on the lock for knee
stability in most conditions. The results are shown in Figure 25.
As predicted in Table 14, posterior foot translation (POST), dorsiflexion (DFLX), tight lock
spring (TTLS), and walking down a slope (DEC), resulted in longer lock engagement than
anterior foot translation (ANT), plantarflexion (PFLX), loose lock spring (LSLS), and walking
up a slope (INC), respectively (Figure 25).
The participant appears to have been more dependent on the lock for maintaining knee stability
in the lazy (LZY) condition while mimicking the gait of a weak ambulator, and maintained lock
engagement longer in both the slow (SLW) and lazy (LZY) gait conditions, than in fast (FST)
walking (Figure 25).
While it was expected that walking on uneven terrain like rocks and grass would generate greater
variability in stability requirements, the knee lock became engaged only once in all of the rock
(RCK) and grass (GRS) gait cycles analyzed, therefore variability could not be measured
67
0
5
10
15
20
25
30
35
40
Lock
En
gage
men
t D
ura
tio
n (
% s
tan
ce-p
has
e)
Condition
Lock Engagement Duration
NEU
T
AN
T
PO
ST
PFL
X
DFL
X
LSLS
TTLS
INC
DEC
FST
SLW
LZY
Figure 25: Average lock engagement duration under different prosthetic alignment, All-Terrain Knee
setting, terrain, and walking speed conditions tested by the above-knee amputee participant.
68
3.5 Limitations
There were a number of limitations in this study which may have affected the quality of data
collected and its analysis. There was a very small pool of eligible above-knee amputees for
participation in this study (n=4) and the successful recruitment of only one limited the
generalizability of his results. The recruitment of able-bodied participants was convenient and
the use of a prosthetic simulator has been shown to generate results representative of a weak or
inexperienced population of prosthesis users, however the comparison is not perfect. The knee
joint on a simulator is necessarily lower than the contralateral anatomical joint creating
additional asymmetry in simulated prosthetic gait, and below-knee translational alignment
changes are not commonly used in clinical practice. The prosthetic simulator in this study was fit
and aligned by an amateur with limited experience in prosthetic alignment and gait training, and
the design of the simulator provided only a limited fit on the legs of different participants, both
of which may have contributed to variability in able-bodied participant data. The prevalence of
toe-drag among able-bodied participants may have been a result of improper alignment or a
deficiency in training and also had a significant effect on some spatiotemporal parameters of
simulated gait. Identification of gait cycles involving toe-drag was done retrospectively and,
despite relatively clear indications of toe-dragging in the collected data, may have been
incomplete. Future data collection may involve video recording to facilitate post-acquisition
analysis. Additional limitations in data collection were issues with the fixation of the force-
sensing resistor and accelerometer which were identified in pilot trials but still required some
iterations to solve. The flexibility of the force-sensing resistor made it susceptible to bending
which may have introduced noise to the collected data. Since participant-generated signals were
used to synchronize data from the load transducer and developed sensor system, variable
strength, stomp technique, or terrain may have affected data feature identification; efforts were
made to select the most appropriate thresholds to address this limitation and minimize its effect
on temporal data comparisons. The selection of thresholds used to define data features, like
synchronization stomp onset, throughout this study was done by trial and error and thus, may not
have been optimized. Similarly, data extraction was done manually which was less sensitive to
some data artifacts than an automated system would have been but may have been a source of
error.
69
3.6 Future Work
With the development of a more robust and user-friendly iteration of the ASPL-SS evaluated in
this study, the potential to conduct a more thorough exploration of ASPL function is promising.
The final ASPL-SS mount and container should enclose all components of the sensing system
securely. Its weight and size should not affect gait biomechanics. Using rechargeable and more
compact power alternatives where possible would reduce maintenance and streamline the design.
The inductive proximity sensor should be mounted in such a way that its relative position to the
front surface of the All-Terrain Knee is fixed, thereby creating an absolute reference point for
knee lock position measurements and ensuring that the knee lock always moves within the linear
sensing region of the sensor. The mount should ensure minimal relative motion between its
sensors and the All-Terrain Knee. The SD card and Arduino USB plug should be easily
accessible without disassembling the sensing system and the system should provide a visual
indication of its ON/OFF and data recording status and battery life.
Following the completion of this study, I also have several recommendations for the future
directions of related clinical research. Including a larger and more varied sample of the relevant
user population would generate a database from which comparisons and generalizable
conclusions could be made regarding the function of the ASPL under variable conditions.
Conducting longer walking trials would indicate whether dependence on the knee lock for
stability increases with fatigue, and repeating trials throughout the lifespan of an All-Terrain
Knee may help identify component deterioration. Future trials should be conducted under more
clinically relevant alignment conditions, in particular, with translational adjustments applied
above the knee rather than below it, as well as with a wider range of All-Terrain Knee setting
permutations including lock spring and friction shim tightness, and different extension spring
stiffness. In addition to walking trials, a study of ASPL function in various activities of daily
living and more challenging mobility conditions would contribute to the optimization of the
mechanism and quality of life for users with different needs. Alignment and All-Terrain Knee
set-up conditions should be combined with variable terrain and gait conditions in order to
compare the compensatory stability provided by each in challenging scenarios. The developed
sensor system could also be used in conjunction with traditional gait analysis instruments for
more in-depth biomechanical analyses of the kinetics and kinematics of All-Terrain Knee gait, or
70
with qualitative measures of participant-reported balance, stability, and comfort in order to more
closely replicate the clinical alignment process.
Systematic sensor placement and function checks should be incorporated in the experimental
protocol to minimize physical or calibration drift. Should the prosthetic simulator be used in the
future, efforts should be made to ascertain comparable initial alignment among participants.
Employing the help of a certified prosthetist would contribute to consistency.
For any future ASPL-SS trials involving the load transducer or another data collection system,
auto-synchronization of the time stamps from each system would alleviate the uncertainty
described in sections 2.3.1 and 3.5. Kinetic or kinematic features selected for gait event detection
should be validated against gold standard force plates or motion capture software, especially for
populations demonstrating pathological gait. The development of an automated algorithm for
data analysis would accelerate the interpretation of results, allow for feature-identification
threshold optimization, and eliminate human error. In larger studies, factorial ANOVA may be
useful to measure the effect of multiple independent variables on the function of the ASPL.
71
Conclusions
The development of the Automatic Stance-Phase Lock-Sensing System addresses the previously
unmet need to monitor the internal function of the All-Terrain Knee wirelessly and without the
limitations of an instrumented gait laboratory. The results of this study indicate that the selected
set of sensors which comprise the system are sensitive to knee lock position changes, transitions
between natural knee extension and flexion, and gait events including heel strike, toe-off, and
terminal impact. Using data collected by the developed system, the fundamental relationships
between applied moments and knee lock engagement, which are central to All-Terrain Knee
stance-phase control, were confirmed. The results of comparing data across various relevant
conditions alluded to the responsiveness of the Automatic Stance-Phase Lock mechanism to
changes in gait speed, terrain, and mobility conditions, and demonstrated how joint stability may
be augmented by adjustments to prosthetic alignment or All-Terrain Knee settings.
With the addition of a robust mounting system and more user-friendly software interface, the
developed system has the proven potential for use in larger biomechanical studies to inform All-
Terrain Knee design iterations, and clinically, to optimize prosthetic alignment and set-up in
real-time based on individual user stability requirements.
The following is are my perceived contributions to the field:
1. Development and engineering validation of a stand-alone, portable sensor system capable of
monitoring the stance-phase control function of the ASPL mechanism continuously
throughout gait in indoor and outdoor environments, on flat ground or over obstacles, and
under variable prosthetic conditions.
2. Development of a wireless data logging and transmission system and software to record and
relay collected sensor data to researchers in real-time.
3. Collection and synthesis of data from walking trials with multiple participants under variable
conditions using the developed system and an additional load transducer.
4. Confirmation of control and knee axis moment relationships with lock position and natural
knee extension and flexion which define the unique ASPL stance-phase control technique.
5. Preliminary exploration of the effects of prosthetic alignment, All-Terrain Knee settings, gait
speed and technique, and terrain on the stability requirements and provision by the ASPL.
72
6. Generated feasibility data and recommendations for the design of a clinic-ready ASPL-SS
and for conducting future research on the biomechanical and clinical performance of the All-
Terrain Knee.
73
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79
Abbreviations and Glossary AB Able-bodied
ACC Accelerometer
ACCx Horizontal acceleration
ACCz Vertical acceleration
AK Above-knee
ANOVA Analysis of variance
ASPL Automatic Stance-Phase Lock
ASPL-SS Automatic Stance-Phase Lock-Sensing System
FSR Force-sensing resistor
FZ Vertical force
GC Gait cycle
HS Heel strike
IPS Inductive proximity sensor
MS Mid-stance
SD Standard deviation
SP Stance-phase
TI Terminal impact
TO Toe-off
The following glossary lists the definitions of terms as they are used throughout this report.
Accelerometer (ACC): 3-axis accelerometer mounted on the All-Terrain Knee to measure
horizontal acceleration (ACCx) in the anterior-posterior direction and vertical acceleration
(ACCz) in the superior-inferior direction. Accelerometer data was used to detect gait events.
All-Terrain Knee: Mechanical prosthetic knee joint for transfemoral amputees. All-Terrain Knee
may be used interchangeably with “knee” or “joint”. The components shown in Figure 2 are also
referred to throughout this report. All-Terrain Knee flexion occurs about the knee axis and the
knee lock rotates about the control axis.
Automatic Stance-Phase Lock (ASPL): Internal mechanism of the All-Terrain Knee which
provides stance-phase control. Comprised of the knee lock, lock spring, and control axis.
“ASPL” may be used interchangeably with “mechanism” or “stance-phase control mechanism”.
Automatic Stance-Phase Lock-Sensing System (ASPL-SS): Developed system of sensors
including the accelerometer, force-sensing resistor, and inductive proximity sensor, as well as the
data logger and transmitter and their associated circuitry and power supplies. “ASPL-SS” may be
used interchangeably with “system” or “sensing system”.
80
Contact Interface: Surfaces of the thigh component and knee body which make contact when the
All-Terrain Knee is fully extended. In Figure 2, the extension bumper is at the contact interface.
The Knee Body surface of the Contact Interface is where the force-sensing resistor rests.
Engagement/engaged: The state of the ASPL wherein the knee lock is rotated anteriorly and is
securing the thigh component such that the All-Terrain Knee cannot flex. The knee lock may be
rotated anteriorly while the All-Terrain Knee is flexed, however, in this state it is not
contributing to stability and is not considered “engaged”.
Engineering validation test: Used on prototypes to verify that the design meets pre-determined
specifications and design goals. Includes functional tests, signal quality tests, conformance tests,
and specification verification. All references to validation and validity are with regards to
engineering validation.
Extension moment: At the knee axis: applied moments which tend to extend the All-Terrain
Knee. At the control axis: applied moments which tend to disengage the knee lock (Figure 3).
External moment: Moment applied due to gravitational forces rather than “internally” by
muscular contraction, bone-on-bone forces, or tension in soft tissues.
Flexion moment: At the knee axis: applied moments which tend to flex the All-Terrain Knee. At
the control axis: applied moments which tend to engage the knee lock (Figure 3).
Force-sensing resistor (FSR): Sensor mounted at the contact interface of the All-Terrain Knee
to detect full extension of the joint and onset of flexion of the joint.
Heel strike (HS): The instant of contact between the foot and the ground. Defines the beginning
and end of a gait cycle. Defined throughout this report by a rapid increase in compressive force
measured by the load transducer following a relatively stagnant period of tension (i.e. swing-
phase).
Inductive proximity sensor (IPS): Sensor mounted on the anterior surface of the All-Terrain
Knee to measure the relative anterior-posterior position of the knee lock.
81
Knee extension: Increasing the angle between the thigh and shank segments of the leg. Full
extension refers to the time when the thigh component and knee body of the All-Terrain Knee are
in contact (i.e. natural or hyper- extension) or when the knee lock is engaged and preventing knee
flexion.
Knee/joint flexion: Decreasing the angle between the thigh and shank segments of the leg. When
the All-Terrain Knee is not hyperextended and the knee lock is not engaged, the All-Terrain
Knee is free to flex, or bend.
Load transducer: 6 degree-of-freedom portable force and torque transducer used to validate the
function and use of the developed ASPL-SS. Primarily used for measurement of vertical force,
and moment at the control and knee axes of the All-Terrain Knee in this study. Mounted below
the All-Terrain Knee in the prosthetic simulator and in the prosthetic leg of the above-knee
amputee participant, the sagittal axis of the transducer is parallel to the control and knee axes.
Mid-stance (MS): The increase in extension moment that occurs in stance-phase as the ground
reaction force vector transitions from posterior to the knee axis to anterior. Generally, the
transition between weight-acceptance and propulsion. Usually marks the beginning of knee
hyperextension.
Stability: The ability of the leg to resist flexion and remain supportive throughout weight-
bearing. May occur naturally when an external extension moment is applied at the knee axis
and/or be supplemented by knee lock engagement.
Stance-phase control: Controlling flexion in weight-bearing.
Terminal impact (TI): The instant in late swing-phase when the leg becomes fully extended in
preparation for weight-acceptance.
Toe-off (TO): The instant when the trailing foot leaves the ground, ending stance and initiating
swing-phase. Defined throughout this report as a peak in tension force following a period of
large compressive forces (i.e. stance-phase) and preceding a return to a relatively stagnant period
of tension (i.e. swing-phase).
82
Appendix A: Code
This appendix contains the Arduino and MATLAB code used to collect and analyze walking trial
data.
A.1 Arduino Data Logger //Reads, displays, and records ASPL-SS sensor values.
#include <SPI.h>
#include <SD.h>
#include <DS3231.h>
#include <Wire.h>
//SD card attached to the SPI bus as follows:
// MOSI 11
// MISO 12
// CLK 13
// CS 4
//chip select pin
const int chipSelect = 4;
////AnalogRead pins
//const int xPin = 0; //Vertical acceleration
//const int zPin = 1; //Horizontal acceleration
//const int proxPin = 2; //Lock position
//const int presPin = 3; //Contact force
// SDA = 4; //Real time clock
// SCL = 5; //Real time cock
//digitalRead/Write pins
const int switchPin = 3;
const int LEDPin = 5;
//The min and max values from accelerometer while standing still
int minVal = 257;
int maxVal = 393;
//The min and max values from IPS
int minPVal = 0;
int maxPVal = 1023;
//Button state (SD write on or off)
int switchState = 0;
//Set up RTC parameters and variables
DS3231 Clock;
bool Century=false;
bool h12;
bool PM;
byte ADay, AHour, AMinute, ASecond, ABits;
bool ADy, A12h, Apm;
byte year, month, date, DoW, hour, minute, second;
83
void setup() {
pinMode(switchPin, INPUT); //initializes switch pin as an input
pinMode(LEDPin, OUTPUT); //initializes LED pin as output
Wire.begin(); //Start I2C interface
Serial.begin(57600); //initialize serial communication
Serial.print("Initializing SD card...");
if (!SD.begin(chipSelect)){ //check chipSelect pin for SD communication,
status notification
Serial.println("Card failed, or not present.");
return;
}
Serial.println("Card initialized.");
}
void loop() {
switchState = digitalRead(switchPin); //read switch input pin
if (switchState == HIGH) { //if switch is ON
File myFile = SD.open("test.csv",FILE_WRITE); //open file to write to it
if(myFile){//if file opens,
digitalWrite(LEDPin, HIGH); //turn LED on
myFile.println ("New Test"); //separate data between switches
int date = Clock.getDate(); //prints date from RTC
int month = Clock.getMonth(Century);
int year = Clock.getYear();
myFile.print("20");
myFile.print(year, DEC);
myFile.print("-");
myFile.print(month, DEC);
myFile.print("-");
myFile.println(date, DEC);
while(switchState == HIGH) { //while the switch is ON
String dataString = "";
int second = Clock.getSecond(); //prints time from RTC, once/sensor
analog read
int minute = Clock.getMinute();
int hour = Clock.getHour(h12, PM);
dataString += String(hour, DEC);
dataString += ",";
dataString += String(minute, DEC);
dataString += ",";
dataString += String(second, DEC);
dataString += ",";
dataString += String(millis());
dataString += ",";
84
for(int analogPin = 0;analogPin < 4;analogPin ++){ //read sensor
analog pins
int sensor = analogRead(analogPin);
if(analogPin < 2){
sensor = map(sensor, minVal, maxVal, -100, 100); //convert
accelerometer values to -1g to 1g
}
else if(analogPin == 2){
sensor = map(sensor, minPVal, maxPVal, 0, 400); //convert IPS
values to mxe^-5 0-400 (0.00-4.00mm)
}
dataString += String(sensor);
dataString += ",";
}
myFile.println(dataString); //print dataString in file
Serial.println(dataString); //print dataString (time and sensor
values) to serial monitor
switchState = digitalRead(switchPin); //check switch input pin
}
}
else{ //if file does not open
Serial.println("Error in opening the file.");
return;
}
myFile.close(); //close the file when switch is OFF
digitalWrite(LEDPin,LOW); //turn LED OFF
}
else {//if switch is OFF,
String dataString = "";
for(int analogPin = 0;analogPin < 4;analogPin ++){
int sensor = analogRead(analogPin);
if(analogPin < 2){
sensor = map(sensor, minVal, maxVal, -100, 100);//convert
accelerometer values to -1g to 1g
}
else if(analogPin == 2){
sensor = map(sensor, minPVal, maxPVal, 0, 400);//convert IPS
values to mxe^-5 0-400 (0.00-4.00mm)
}
dataString += String(sensor);
dataString += ",";
}
Serial.println(dataString); //print sensor values to serial monitor, no
time
}
}
85
A.2 MATLAB Synchronization and Analysis function [] = asplssStompSyncedits close all %Processes and plots IPS, FSR, Accelerometer, and load transducer data %Synchronizes ASPL-SS data with ATI data based on slope change within first %500 readings for AccX and FZ which correspond to stomp action
%%Load all system data %Load ATI .csv file [atiFile,atiPath] = uigetfile('*.csv','Select the ATI data file.'); atiData = xlsread(fullfile(atiPath,atiFile));
%Load ASPLSS .csv file [asplssFile,asplssPath] = uigetfile('*.csv','Select the ASPLSS data file.'); asplssData = xlsread(fullfile(asplssPath,asplssFile));
%User-entered neutral lock position (in mmx10^2) prompt = {'Neutral Proximity:'}; dlg_title = 'Input Neutral Proximity'; num_lines = 1; def = {'0'}; answer = inputdlg(prompt,dlg_title,num_lines,def); neutralProx = str2double(answer{1});
%% Time stamps
% Calculate seconds since epoch day (00:00:00, 1st January 1990) to date % and time for ATI data based on file name and Excel timestamps dateA= regexp(atiFile, '(?<=\()[^\(]+(?=\))', 'match'); % regexp extracts %date and time string from ATI filename %Process date and time string to a format readable by MATLAB dateA{1}(1:3)=''; %replaces day (columns 1-3) with '' dateB=strrep(dateA{1},'EDT',''); dateC=strrep(dateB,'-',':'); ref=datevec('01-01-1900, 00:00:00'); %Transform epoch day into date vector cur=datevec(dateC); %Transform current time into date vector NTPTime=etime(cur,ref); %Elapsed secs since epoch day/time to current
day/time cur2=cur; cur2(1,4:6)=0; NTPTime2=etime(cur2,ref); %Elapsed secs since epoch day/time to current day
%ATI ModNTPTime = floor(NTPTime/(2^20))*(2^20); %Masks lower 20 bits of NTP Time CorNTPTime = ModNTPTime-10800; %Corrects for EDT Timezone ROTime = atiData(:,1)./4096; %Seconds since last rollover ROTime(ROTime<0) = ROTime(ROTime<0) + 2^20; % Make -ve timestamps +ve atiTime = ROTime+CorNTPTime; %col vec = # secs since epoch time for each row atiData(:,1) = atiTime; %substitute col vec for original ATI time stamps
%% Synchronization
%ATI Fz atiDataTrunc = atiData(1:500,9); %create a truncated matrix of ATI data
86
fzDiff = 0; i = 1; while (fzDiff) > -30000; %search for first instance of deltaFz > -30000 i = i+1; fzDiff = atiDataTrunc(i)-atiDataTrunc(i-1); end fzStartInd = i-1; atiTime = atiData(:,1)-atiData(fzStartInd,1); %set sync timestamp = 0
%ASPLSS AccX asplssDataTrunc = asplssData(1:1000,5); %create a truncated matrix of ASPL-SS
data accXDiff = 0; j = 1; while (accXDiff) < 40; %search for first instance of deltaAccX > 40 j = j+1; accXDiff = asplssDataTrunc(j)-asplssDataTrunc(j-1); end accXStartInd = j-1; asplssTime = asplssData(:,4)-asplssData(accXStartInd,4); %set sync timestamp
= 0 asplssTime = asplssTime/1000; %convert to seconds from ms
%% ATI data origin asdjustment
FX = atiData(:,7)/1000; % /1000 to correct for ATI errors FY = atiData(:,8)/1000; % /1000 to correct for ATI errors FZ = atiData(:,9)/1000; % /1000 to correct for ATI errors TY = atiData (:,11)/1000; % /1000 to correct for ATI errors DX = 0.0235; DX2 = 0.0165; DZ = 0.06433; DZ2 = 0.1953;
CAMoment = TY + FX*DZ + FZ*DX; %control axis KAMoment = TY + FX*DZ2 + FZ*DX2; %knee axis
%% Data set variables
AccX = asplssData(:,5)/100; %Vertical acceleration AccZ = asplssData(:,6)/100; %Horixontal acceleration Prox = (asplssData(:,7)-neutralProx)/100; %Lock position (mm) Pres = asplssData(:,8); %FSR force
%maxima, for plotting normalized graphs MAccX = abs(max(abs(AccX))); MAccZ = abs(max(abs(AccZ))); MProx = abs(max(abs(Prox))); MPres = abs(max(abs(Pres))); MCAMoment = abs(max(abs(CAMoment))); MKAMoment = abs(max(abs(KAMoment))); MFZ = abs(max(abs(FZ)));
%normalize normAccX = AccX/MAccX;
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normAccZ = AccZ/MAccZ; normProx = Prox/MProx; normPres = Pres/MPres; normCAMoment = CAMoment/MCAMoment; normKAMoment = KAMoment/MKAMoment; normFZ = FZ/MFZ;
zero = zeros(size(CAMoment)); %to plot a horizontal line at 0 x1=asplssTime; x2=atiTime;
%% Plots
figure %ACCFZ [h1,~,~]=plotyy(x1,[AccX AccZ],x2,[FZ]); title(['Knee Acceleration and Vertical Force']); ylabel(h1(1),'Acceleration (g)'); ylabel(h1(2),'Force (N)'); xlabel(h1(1),'% Gait Cycle'); set(h1(1),'YLim',[-5 5]); set(h1(1),'YTick',[-5:1:5]); set(h1(2),'YLim',[-1000 1000]); set(h1(2),'YTick',[-1000:250:1000]); legend('Acceleration_{Vertical}','Acceleration_{Horizontal}','Force_{Vertical
}');
figure %MCIPSFZ [h2,~,~] = plotyy(x1,[Prox],x2,[CAMoment,FZ/10,zero]); title(['Lock Position and Control Axis Moment']); ylabel(h2(1),'Lock Position (mm)'); ylabel(h2(2),'Moment (Nm)'); xlabel(h2(1),'% Gait Cycle'); set(h2(1),'YLim',[-3 3]); set(h2(1),'Box','off'); set(h2(1),'YTick',[-3:0.5:3]); set(h2(2),'YLim',[-100 100]); set(h2(2),'YTick',[-100:25:100]); legend('Lock Position','Moment_{Control Axis}');
figure %MKFSRFZ [h3,~,~] = plotyy(x1,[Pres/10],x2,[KAMoment,FZ/10,zero]); title(['Knee Extension and Knee Axis Moment']); ylabel(h3(1),'Knee Extension'); ylabel(h3(2),'Moment (Nm)'); xlabel(h3(1),'% Gait Cycle'); set(h3(1),'YLim',[-100 100]); set(h3(1),'Box','off'); set(h3(1),'YTick',[-100:25:100]); set(h3(2),'YLim',[-100 100]); set(h3(2),'YTick',[-100:25:100]); legend('Force_{FSR}','Moment_{Knee Axis}');
figure %AllSensors [h4,~,~] =
plotyy(x1,[normProx,normPres],x2,[normCAMoment,normKAMoment,normFZ,zero]); xlabel(h4(1),'% Gait Cycle');
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legend('Lock Position','Force_{FSR}','Moment_{Control Axis}','Moment_{Knee
Axis}','Force_{Vertical}');
end