iomimetic design of a ontrollable knee actuator – a tutorial...iomimetic design of a ontrollable...
TRANSCRIPT
Biomimetic Design of A Controllable Knee Actuator – A Tutorial
Mahsa Farshi Taghavi
The material in this tutorial is based in part
on Wearable Robots: Biomechatronic
Exoskeletons (1st Ed., 2008) by Jl Pons and
my own research. For more information,
please write to
© 2019 K. N. Toosi University of Technology
nee–ankle–foot orthoses are
prescribed as a partial solution
for joint disorders to provide
stability and keep joints in their functional
positions. Conventional systems provide
stability during walking by maintaining the
knee in a fixed position, but this produces
unnatural gait patterns.
This tutorial has been developed to help
you understand what Knee ankle foot
orthoses(KAFOs) is, why it is important, and
how to present computational modeling for
it. First, the kinematics of normal knee joint
and normal knee stiffness during the gait
cycle are presented, which is critical for
understanding knee behavior and
developing better KAFOs. Components of
KAFOs are then described including their
Actuator, Sensor, and controller.
Introduction
Knee ankle foot orthoses (KAFOs) are
frequently prescribed to cerebrovascular
accident, poliomyelitis or cerebral palsy
patients with leg muscle weakness, in order
to provide knee stability, reduce falling risk
and enable a certain degree of mobility. The
concept of automatic compensation of
human walking consists of providing
dynamical adaptation of the artificial
joint/skeleton, (e.g. allowing knee flexion
during the swing phase of gait, controlling it
K
during the stance phase at a certain range)
mimicking an average normal profile. The
control strategies applied at the level of the
joints for functional compensation of gait in
wearable devices can be classified in: (1)
position control, (2) impedance control, or
(2) intermittent joint control strategies. By
means of error position adjustment in a
closed control loop, human joint rotation
can be tracked and defined with an active
exoskeleton. In a basic scheme, the control
loop is treated as a black box that provides
the demanded torque [1].
During locomotion, the knee provides
shock absorption, maintains stability during
stance, and contributes to limb progression
during swing. The knee extensor muscles
resist knee flexion during stance, absorb
forces due to body weight, and facilitate
limb progression. When an external knee
flexion moment acts on the knee, the knee
extensor muscle group generates an
opposing extension moment. If the knee
extensors are not sufficiently strong, the
knee will collapse due to the external knee
flexion moment and the person will fall [1].
A knee actuator system based on energy
storage release has been designed for a
KAFO to apply functional compensation to
the knee throughout the gait cycle. This is
done by means of two elastic actuators
whose elastic constants adapt to the
different phases of the cycle so as to
approach a normal profile. The approach
adopted is a biological one, and the
principle underlying the design consists in
using biomechanical data from the leg to
determine the configuration of the
actuators and actions that are applied at
joint level. The absence of the necessary
muscle control in the leg segments can
affect locomotion in a variety of ways, from
an undesired gait pattern to bodily collapse.
For designers of actuator systems, it is
helpful to analyses the possible situation in
each joint when these problems occur [1].
Knee extensor weakness
The knee extensor muscles maintain
knee joint stability during stance phase.
These muscles, primarily the quadriceps
muscle group, resist knee flexion during the
early stance to absorb shock and facilitate
limb progression. If the knee extensors are
not sufficiently strong, the knee will
collapse due to the external knee flexion
moment [2]. Muscular weakness can result
from peripheral neurological diseases (e.g.,
poliomyelitis), muscular diseases (e.g.,
Duchene muscular dystrophy), central
neurological diseases (e.g., multiple
sclerosis), spinal cord injury, osteoarthritis,
and severe injury. The effects of knee
extensor weakness can range from the
individual having an abnormal gait pattern
to complete instability. Even if the extensor
muscles are only slightly weakened, walking
with an abnormal gait pattern is highly
energy consuming and can cause further
soft tissue damage [4].
Pathological gait
Individuals with knee extensor
weakness can range from having a slightly
unnatural gait pattern to complete
instability while walking, depending on the
severity of their condition. The knee
extensor muscles are active in both stance
and swing phases since these muscles
control stance and the flexion rate during
early stance. Individuals with knee extensor
weakness are at risk of collapsing when an
external knee flexion moment is acting on
the knee joint (figure 2 and figure 3).
To avoid having the GRF vector pass
behind the knee joint center, certain
techniques are adopted. In particular,
people increase hip extensor muscle
activity and anterior trunk flexion to shift
the body center of gravity forward. Long
term effects of this strategy may include
knee joint hyperextension since the joint is
constantly being loaded to a fully extended
position.
Individuals will have trouble negotiating
stair descent if they have moderate to
severe knee extensor weakness, due to the
large external knee flexion moments [4].
Knee joint function during
locomotion
The knee is the largest synovial joint in
the body, connecting the lower limb (shank)
to the upper limb (thigh). Knee movement
is primarily in the sagittal plane. Fourteen
muscles control the knee during gait, and
can be divided into extensor and flexor
groups. Knee extensor muscles decelerate
knee flexion during stance and contribute
to limb progression during swing. During
the swing phase, the knee extensor muscles
also contribute to limb progression [2].
Normal level ground walking
Gait cycle
The gait cycle is composed of two
phases, stance and swing (Figure 1). Stance
phase is when the limb is in contact with
the ground (foot contact with ground until
foot leaves the ground). Stance sub-phases
are initial contact, loading response, mid
stance, and terminal stance. Swing phase is
when the limb is in the air (foot off until
foot contacts the ground again). Swing sub-
phases include pre-swing, initial swing, mid
swing, and terminal swing [2].
Figure 1:Phases and sub-phases of the gait cycle [3].
The gait cycle can also be divided into
three tasks:
weight acceptance, single limb support,
and swing limb advancement. Weight
acceptance is the first task of stance phase,
and is comprised of initial contact and
loading response. Weight acceptance
includes shock absorption when the foot
contacts the floor, limb stability, and
preservation of forward limb progression.
Single limb support is the second task and
involves mid stance and terminal stance
sub-phases, where body weight is
supported and the opposite limb is in the
air.
Swing limb advancement is when the
limb is lifted in the air and prepares for the
next stance phase. Pre-swing, initial swing,
mid swing, and terminal swing sub-phases
are part of the limb advancement task.
Table 1 defines each gait cycle sub-phases
[2].
Knee function during walking
During stance phase, the knee provides
limb stability during weight bearing and
prepares the limb for swing phase. The
ground reaction force (GRF) vector with
respect to the knee joint center determines
the external moment’s direction. When the
GRF vector is anterior to the knee joint
center of rotation, an external extension
moment acts on the knee and a GRF vector
posterior to the knee joint center of
rotation produces an external flexion
moment at the knee (Figure 2) [4].
Table 1:Gait cycle sub-phases [2].
Phase Task Sub
phase
% gait
cycle Description Function
Stance
Weight
acceptance
Initial
contact 0-2
Foot contact
with ground,
knee is extended
Stable weight bearing
Loading
response 2-12
Active knee
flexion (20˚
flexion)
Shock absorption,
stability, anterior knee
joint movement
Single limb
support
Mid
stance 12-31
Active knee
extension (15˚
flexion)
Stable weight bearing,
advance femur over
tibia
Terminal
stance 31-50
Active maximum
knee extension
(5˚ flexion)
Stable weight bearing,
femur advances over
tibia to max knee
extension
Swing limb
advancement
Pre swing 50-62
Passive knee
flexion (40˚
flexion)
Knee prepares for
swing phase and toe
clearance
Swing
Initial
swing 62-75
Active knee
flexion (60˚
flexion)
Knee flexes to advance
limb and foot
clearance
Mid
swing 75-87
Passive knee
extension
Knee extends passively
(hip flexor moment) to
advance limb
Terminal
swing 87-100
Active knee
extension
Knee extends to
prepare for stance
phase
Figure 2: External moments acting at the
knee: a) external extension moment when
the GRF vector is anterior to the knee joint
axis, b) external flexion moment when the
GRF vector is posterior to the knee joint axis
[4].
Figure 3 illustrates GRF vectors during
stance phase (initial contact to pre-swing).
When an external extension moment acts
on the knee, the knee is in a stable state.
When an external flexion moment acts on
the knee, the joint is in an unstable position
and an extension moment is required to
oppose flexion and maintain stability, or
the knee will collapse. This extension
moment is generated by the knee extensor
muscles (quadriceps), which are the
dominant muscle group at the knee. During
stance phase for a normal gait cycle, the
quadriceps create an extension moment
about the knee to decelerate joint flexion
due to external flexion moments.
Figure 3: Ground reaction force vector
magnitudes and directions for initial contact,
loading response, mid stance, terminal
stance, and pre-swing [4].
The GRF vector moves posterior to the
knee joint axis during loading response,
mid-stance, and pre-swing. As seen in
Figure 3, the GRF magnitude between
loading response and terminal stance is
large and posterior to the knee joint axis,
requiring substantial extension moments
from the knee extensor muscles.
The knee joint flexes twice during a
normal gait cycle. The first is during loading
response and mid-stance, to absorb the
impact from foot contact and advance the
knee anteriorly. The second is during early
swing phase, for foot clearance and
following step preparation [2].
Knee kinematics during normal
gait
The knee joint has a large range of
motion (ROM) and moves in all three
anatomical planes. Coronal plane motion
(adduction/abduction) maintains vertical
balance over the limb during single support
and transverse plane rotation (internal
rotation) accommodates for alignment
changes. The knee joint’s motion occurs in
the sagittal plane (flexion/extension) for
progression and advancement during swing
and stance, as well as foot clearance (Figure
4) [2].
Figure 4:Movements about the knee joint [2]
Figure 5 shows the knee angle for able-
bodied level ground natural cadence
walking in the sagittal plane. The vertical
line at 60 percent of the stride is the
transition between the stance and swing
phases of the gait cycle.
The knee ROM in the sagittal plane is
approximately 0-70˚during the gait cycle,
with two flexion movements (Figure 5).
The first flexion peaks at approximately
20˚ during the transition between the
loading response and mid stance, where
the knee absorbs the impact from foot
contact. The knee then extends to almost
full extension (8˚) and then begins to flex
again to prepare for the swing phase. Toe-
off occurs at 60 percent of the stride cycle,
at a knee angle of approximately 40˚.
Flexion peaks at 70˚ in the initial swing
phase and then extends to full extension
for foot contact [2], [4].
Figure 5 :Knee angle for able-bodied, natural
cadence, level ground walking with standard
deviation shown with a dotted line (flexion
is positive) [4].
Functional analysis of gait as
inspiration
Imitation of muscle design can be
envisaged as a first step towards a new
type of functionally diverse and robust
actuators.
But future actuators (e.g. polymer-
based) might not be able to include the
underlying mechanical principles of
biological muscles, and that is reflected in
undergoing research on imitation of muscle
functionality based on alternative means.
Moreover, a biogenetic approach for
functional compensation of pathological
gait ought to be versatile and customizable.
Critical construction requirements for such
an actuation system are low volume and
size, low energetic consumption, low heat
dissipation and a high torque [1].
In order to design the compensator for a
pathological gait profile, our approach is
based on the identification of mechanical
performance ranges of knee and ankle
joints during normal performance.
Mechanical functions at specific ranges are
identified from normal gait data profiles of
average, and low speeds, given by Winter.
It can be seen that approximately between
5 and 15% of the gait cycle (at stance),
when the joint is absorbing the impact
(through the action of quadriceps), the
performance of the joint, by comparison of
the angular position against the torque at
the knee—load-displacement
relationship—, can be associated with an
elastic performance. Although theoretically
it is accepted that in this period there is a
damping effect at the knee, our findings
suggest that an elastic behavior can be
appropriated to apply compensation which
includes recovery [5].
During this interval the musculoskeletal
activity is dedicated to power absorption
and then immediately, corresponding to
the extension trajectory, this power is
recovered until approximately 30% of the
phase. This roughly elastic relation starts at
the beginning of the gait cycle (around
1.5%) and is maintained after the energy
generation, until approximately 50% of the
cycle. Then, the knee extension is
completed and the knee is prepared to
swing [5].
The evolution of the ratio between
torque and angle for the knee joint is
illustrated in Figure 6 (normalized data).
Various different modes of operation can
be identified by analyzing how the
rotational displacement of the joint relates
to its torque in the gait cycle [5].
Figure 6: Normal gait average biomechanical
data. Knee angle versus torque in the
sagittal plane during a complete gait cycle.
Identification of elastic constants (dashed
lines) [1].
According to the intervals defined in Fig.
6, the shock absorption area and the
recovery of extension just before the swing
phase is characterized by the segment A–
B–C. The flexion phase during swing is
characterized by the segment C–D and the
extension phase during swing is
characterized by the segment D–E [5].
Actuator design
The actuator system presented in this
tutorial is inspired by the function of the
real muscles of the leg, and it is conceived
to mimic its behavior. Consequently, the
starting point in the design process will be
the actions in the joints presented before.
From an engineering perspective an
analogy of the operation of the human
musculoskeletal system of the lower limb
with a mechanical system can be
established and its functionality can be
considered as actuation functionality or as
a combination of them. In this section,
actuation functions for each joint are
identified during each phase, and thus, a
first approach of a mechanical adjustment
of gait by elastic means is described.
Finally, a design of an actuator system is
proposed.
This prototype was adapted to the GAIT
orthosis (Figure 7), a novel knee, ankle and
foot orthosis (KAFO). It has been designed
to be modular and adaptable to subjects
with different anthropometric
characteristics. The mechanical structure is
formed by a single sided frame with two
joints, knee and ankle. Knee hinge is
performed by a four-bar mechanism to
follow the displacement of the helicoidal
instant axis of the knee in the lateral side.
Ankle hinge is performed by a single hinge
placed on the malleoli [5].
The main difference with other systems
is the knee joint concept used. A four-bar
mechanism was designed to follow the
movement of the human knee, simulating
the movement of the knee cruciate
ligaments. The performance of this joint
can be considered as a rotation around an
instant helicoidal axis, which changes its
position during the displacement. The four-
bar hinge therefore follows the
displacement of said helicoidal instant axis
of the knee in the lateral side, to get a
movement similar to the physiological
movement of the knee [5].
Figure 7:Controllable KAFO fitted to patient.
The actuators attached to the structure
Linear solenoid
Sensor set
Ankle passive actuator and
carbon fiber insole
Controller
Knee actuator
apply selectively levels of stiffness at the
joints [1].
Based on the functions of the muscles in
the leg during gait cycle, functions of an
actuator system can be defined as follow:
Stance phase: Shock absorption and
knee assistance back to full extension.
Swing phase: Free knee flexion in early
swing and assisted knee extension in late
swing to prepare for the next foot contact.
The lower leg system consisted of the
thigh, shank and foot segments (see Fig. 8)
[5].
In such system, the group of weak
quadriceps provide partial or null torque at
the knee. The system was restricted to
transmit torque in the sagittal plane by the
kinematics of the hinge. In the KAFO the
ankle actuator was designed as a passive
compensator, by means of two springs
applying different stiff nesses according to
the direction of rotation of the foot. The
knee actuator applies in the stance
phase during a period of time to provide
joint stability and during swing phase the
actuator applies ( ), to store
and recover spring energy to assist leg
extension prior to heel contact. Transition
from to provides a free knee joint,
passing from a restricted and stiff hinge to
a flexing leg [1].
Figure 8 : Partial control objectives of the
knee in gait cycle: a Heel contact (HC) with
foot fall control followed by stabilised knee
extension through K1 at terminal swing; b
Heel off (HO) during stance phase after
controlled flexion of the knee; c During pre-
swing rate of turn of the shank change of
sign. K1 to K2 transition releases the knee; d
shank rotation during knee flexion. K2—
partially charged by inertia—provides
assistance to extension at the end of cycle
[1].
The concept for the knee actuator during
the gait cycle is described in Fig 9.
During the start of the shock absorption
(A) and weight bearing (B), the stance
control spring (green) is active and
compresses while the leg is in the ground.
At start of the swing phase (C), the swing
control spring (blue) is enabled and the
stance control spring is disabled. The
swing phase spring compresses and
recovers the energy completing the swing
phase.
An electromagnetic solenoid is used to
switch between springs as a function of the
gait cycle and to lock the knee flexion if
required. Both springs will provide the
required action for each part of the cycle
when compressed (by the user's weight
during weight bearing, and by a
combination of inertia and push off action
of the ankle during swing phase). The
change between springs will be done with
the leg completely extended [5].
Figure 9: Knee actuator concept and functionalities during the gait cycle. Stacked discs
constitute the stance control spring (top) and one compression spring (below) for the swing
control [5].
Two elastic springs were included in the
attachable knee actuator. Therefore, the
theoretical torsional elastic constant of an
action based on the joint angle needs to be
calculated, for a given body weight. Let us
consider the intervals:
• Interval 1 (Shock absorption- Recovery
of extension):
It can be considered as an interval
between the beginning and 50% of the
stance phase as an approximately linear
relation between applied torque and flexion
angle. This corresponds to the interval
between points A and C in the torque-gait
percentage diagram in Fig. 6. The interval 1
can be approximated as a constant ratio,
similar to the elastic constant defined by
Hook’s law that holds proportionality
between mechanical stress and strain, so a
line equation can be adjusted to this
interval. The selected solution (dashed line
in Figure 6) was to adjust a line from point
0–0 to the maximum torque value, point
22–0.62 in the angle-torque diagram. From
the equation of this line a torsional elastic
constant can be obtained [1]:
Equation 1
• Intervals 2–3 (Flexion and extension
during the swing phase): Interval 2 presents
a nonlinearity in the torque displacement
curve that can be referred to as pseudo
elasticity. This behavior is present when,
after reaching a given loading stress, the
deformation strain augments considerably
with minimum applied stress. Super
elasticity phenomena is known for its
nonlinearity during unloading. In the case of
the knee joint, the torque-angle
relationship only holds pseudo elasticity
behavior during loading. A second interval
featuring an approximately linear relation
between torque and angular displacement
can be approximated between points D and
E, interval 3. The optimal adjustment for
both intervals is a line between A and D,
but the need of starting in the same neutral
position as in stance imposes point 0–0 in
the torque-angle diagram as starting point.
A torsional elastic constant can be
estimated for the swing phase from the
expression [1]:
Equation 2
Multiplying these constants by the angle
in the knee during the complete gait cycle
we can obtain the action of two theoretical
knee actuators featuring this elastic
performance. Figure 01 shows the elastic
adjustment for both constants and the
theoretical action of the actuator with the
configuration explained before, in
comparison with average data of healthy
subjects. This action is a combination of
both adjustments. It uses for stance
angles and for swing angles [5].
With the definitions of the separate
actions per each gait cycle, the next step
was to develop an actuator prototype to
test the concept. It was decided to use a
linear actuator placed in the sagittal plane.
This actuator will apply force on the
orthosis, generating the needed moment
during gait cycle.
joint torque
(NM/kg)
torsional elastic constant
(NM/kg degree)
joint angle for stance phase
(degree)
joint torque
(NM/kg)
torsional elastic constant
(NM/kg degree)
joint angle for swing phase
(degree)
The requirements for the design for the
knee are:
• Stance range of movement 0–22°
• Swing range of movement 0–65°
• Maximum range of flexion 95°
•Minimum torque values for 90° of knee
flexion (sitting position).
The solution adopted for the last
requirement was to align the force applied
by the actuator with the center of rotation
of the orthotic joint, making minimum the
applied moment. From the rest of the
requirements and after considering
different geometric configurations, the
solution adopted for the knee actuator is
formed by two telescopic cylinders, one
containing the stance spring and the other
containing the swing spring, following the
concept explained in Figure 9 [5].
Two types of springs were used in the
prototype. Swing springs are compression
springs, made of stainless steel and with
right hand direction of the helix. Different
constants were available for selection. An
elastic element (8 mm length) was used for
the stance control, applying at the joint
with a maximum longitudinal stroke,
corresponding with a flexion limit. A
compression elastic element (57 mm
length) is provided for the swing phase.
Compression springs made of Stainless
Steel Type 302, were used.
Selection of stance and swing springs in
the knee actuator is done as follows:
assuming no residual actions on the joints,
from equations of elastic adjustment done
previously and with patient weight data,
maximum torque values (stance and swing)
were obtained:
Equation 3
Torque provided by the actuator comes
from the expression:
Equation 4
Applied at maximum stance and swing
flexion angles. While the orthotic knee joint
is a four-bar mechanism it is necessary to
calculate the instant center of rotation at
those maximums in order to obtain x and y
distances. This has been done finding the
intersecting point of both central bars of
the mechanism. The expression of a linear
spring gives us the linear elastic constant of
the springs:
Equation 5
The action provided by the actuator with
the obtained constants for both springs is
different to that calculated in the elastic
Figure 10 : Elastic adjustment for both elastic constants during the complete gait cycle
and final actuator action on the knee [5]. adjustment. This is due to the relative
displacement of the actuator in the
orthosis, but the differences are not
significant [5].
Algorithm required to calculate variation of the spring length to rotate the
desired amount of Knee ankle foot orthoses (KAFOs)
Step 1: calculating the torque (for unit weight ) by using
the torque- degrees graph or equation 1 (or equation 2)
Step 2 : Calculation of the torque based on patient weight
by using equation 3
Step 3 : Calculation the force required to create this
torque by using equation 4
Step 3 : Calculation the force required to create this
torque by using equation 4
Example:
consider the Knee ankle foot orthoses that the stiffness of the spring in actuator is
KN/m[6] and the x and Y position of the KAFO are 10 cm and 15 cm respectively. if the
joint angle for stance phase is 20 degrees and the patient's weight is 60 kg, calculate the values
of torque that comes to the KAFO and also, calculate how much change should be made during
the spring to create this torque.
Step 1:
Step 2:
Step 3:
Required force:
Step 4:
Variation of the spring length:
0.58
Sensor system
For detection of human knee angle
and estimation of human segments’
orientation, combining rate gyroscopes
and accelerometers signals, have been
applied. The sensor setup for control
consisted of a first inertial measurement
unit (IMU) at the foot element inside
the shoe (below the orthotic ankle joint)
and a second unit for the lower bar of
the KAFO. Each IMU is composed by (a)
a single miniature MEMs rate
gyroscope, sensing Coriolis force during
angular rate by measuring capacitance
(Analog Devices ADXRS300, volume less
0.15cm3, weight 0.5g) with a maximum
sensitivity ±300◦ s−1 and (b) a complete
dual-axis (surface micro machined)
200mV/g accelerometer (ADXL202
5mm×4.5mm×1.78mm) [7].
Each unit is housed in a box at foot
and shank orthotic bars, as depicted in
Figure 00. The unit suited at the ankle
bar of the orthosis senses rotational
motion, tilt, tangential and radial
segment accelerations in orthogonal
directions (X and Y), while the majority
of orthosis rotations at the level of
joints and bars take place in the plane of
locomotion progression (sagittal)
considering mechanical constraints
imposed by common orthotic and
prosthetic hinges.
The same rotations and motions (in
sagittal plane) are sensed for the shank
by a second unit, suited at the lateral
aspect of the lower bar. Having human
pathologic gait, characterized by muscle
force absence to control the knee, as
the motion of interest in possible
applications at average (2.6km/h) and
low (2km/h) speeds, signals outside the
band frequency related to gait
kinematics (0.3–20Hz), are rejected
from the sensor outputs with −3dB low
Figure 11: Sensors setup in unilateral
knee ankle foot orthosis prototype [7].
pass filters, while lowering noise floor
by bandwidth limiting. A current
follower drives accelerometers signals
after filtering and amplification; circuits
and microelectromechanical sensors are
disposed in a small two-layer PCB
(Figure 01) inside the housing, to be
feed to an on-board control and
processing unit.
Figure 12:IMU circuit (top view) [7].
Gyroscopes are pre calibrated in a
test bench equipped with an encoder,
by generating known rotations and
measuring angle in repetitive trials. Each
accelerometer separately is calibrated
by measuring the nominal demodulated
analog output while the sensing axis is
placed in line with gravity [7].
IMUs signals
Shank and foot IMUs signals are
digitized through a 10bit A/D converter
at 100Hz, with a 3.3V reference voltage
and a resolution of 2.92mV/bit.
Assuming a robust fixation and shared
motion, data can be used as inputs for
calculation, within a local reference, of a
number of biomechanical parameters.
Accelerometer output signals can be
represented, along axis as
Equation 6
being the sensor linear
acceleration, gravity and white
noise. Combining signals from the
tangentially mounted accelerometers, a
measure of segment angular
acceleration can be obtained [7].
Stance control detection
The embedded control system of a
controlle orthotic joint must be able to
shift safely between stance and swing
knee hinge status. Mechanically driven
available KAFOs rely on a cable driven
mechanism to control knee hinge. This
mechanism is tuned in a way it allows
free swing only when a certain amount
of ankle dorsiflexion has been
overcome. We define the shifting
moment detection of a tuned
mechanism working under the same
principal, installed on a novel orthosis
prototype, as ascertaining reference for
experimentation (Figure 13) [7].
Figure 13: Average knee angle and knee
actuator moments at natural cadence
(sagittal plane). Right: Instrumented
orthosis prototype [7].
Walking controller
From the point of view of the safety
of the patient with unilateral partial or
complete absence of control of the
knee, it is desired that the human motor
system is capable of providing input to
the system with priority with respect to
the response of the controller. A
reactive controller of the knee mode
has been designed, appropriated for
testing in patients with risk of a
collapsing knee. The control system is
presented in Figure14 [1].
Figure 14 :Gait cyclical controller [1].
Dynamic activity detector Detection of dynamical activity can
be performed by the tangential uniaxial
accelerometer attached at the foot
element. Accelerometer signal is a
combination of segment acceleration
and the earth’s gravity. During static
activities the piezoelectric
accelerometer yields signals within a 2g
range. Filtering the accelerometer signal
can indicate the beginning and end of a
dynamical activity, applying a threshold
to the segment linear (tangential)
acceleration, which is overcome at gait
initiation with foot rise [1].
Refrences
[1] Moreno, J. C., Brunetti, F., Rocon, E., & Pons, J. L. (2008). Immediate effects of a controllable knee ankle foot orthosis for functional compensation of gait in patients with proximal leg weakness. Medical & biological engineering & computing, 46(1), 43-53.
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