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

M.FarshiTaghavi@email.kntu.ac.ir.

© 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.

[2] J. Perry and J. M. Burnfield, Gait analysis : normal and pathological function, 2nd ed.. Thorofare, NJ: SLACK, 2010.

[3] Moltedo, M., Baček, T., Verstraten, T., Rodriguez-Guerrero, C., Vanderborght, B., & Lefeber, D. (2018). Powered ankle-foot orthoses: the effects of the assistance on healthy and impaired users while walking. Journal of neuroengineering and rehabilitation, 15(1), 86.

[4] Pritham, C. H. (1994). Biomechanical basis of orthotic management. JPO: Journal of Prosthetics and Orthotics, 6(2), 25A.

[5] Cullell, A., Moreno, J. C., Rocon, E., Forner-Cordero, A., & Pons, J. L. (2009). Biologically based design of an actuator system for a knee–ankle–foot orthosis. Mechanism and Machine Theory, 44(4), 860-872.

[6] Veneman, J. F., Ekkelenkamp, R., Kruidhof, R., van der Helm, F. C., & van der Kooij, H. (2006). A series elastic-and bowden-cable-based actuation system for use as torque actuator in exoskeleton-type robots. The international journal of robotics research, 25(3), 261-281.

[7] Moreno, J. C., de Lima, E. R., Ruíz, A. F., Brunetti, F. J., & Pons, J. L. (2006). Design and implementation of an inertial measurement unit for control of artificial limbs: application on leg orthoses. Sensors and Actuators B: Chemical, 118(1-2), 333-337.