2017 IEEE International Conference on Systems, Man, and Cybernetics (SMC)
Banff Center, Banff, Canada, October 5-8, 2017

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Development of a String-Driven Walking Assist
Device Powered by Upper Body Muscles
Koichiro OHASHI, Yasuhiro AKIYAMA, Shogo OKAMOTO, Yoji YAMADA
Department of Mechanical Systems Engineering
Nagoya University, Nagoya, Japanoohashi.kouichirou@c.mbox.nagoya-u.ac.jp
Abstract—The elderly sometimes fall owing to tripping caused
by decreased dorsal flexion during the swing phase. A wearable
assist device that was able to increase dorsal flexion to reduce therisk of tripping was proposed in this study. This device assists in
dorsal flexion during the swing phase by transferring the tension
force from the upper body to the foot by connecting them using
strings. To improve the efficiency of the assist, the string path was
optimized to maximize the assist torque at the time of minimum
foot clearance (MFC). In this study, two string-path patterns were
tested. Based on the walking experiment results, we confirmedthat the wearable assist device that we developed successfullyincreased both the dorsal flexion and minimum foot clearance
during the swing phase.

Table. 1 Type of walking assist device
Type

Actuator Type [9]Supporter Type

Assist Power○High Power×

Assist Pattern○Variable Assist Pattern

×Constant

Weight×○×○×Limited○

Simplicity ofAttaching
OperatingTime

Keywords— gait motion, assistive device, minimum footclearanceI.

flexion angle and restore the foot clearance in order to avoid
tripping [8].

INTRODUCTION

The walking ability of the elderly gradually declines owing
to the decrease in muscle strength associated with aging.
Recently, wearable gait assist devices (hereinafter referred to as
wearable assist devices) have been proposed for supporting thewalking motion of the elderly. Such wearable assist devices wasable to be used to improve the working efficiency of aged labors.
In addition, increasing demand of rehabilitation associated withincreasing population of the elderly required such assist devices
for several decades. Active leg exoskeleton (ALEX) [1], which
is a rehabilitation wearable assist device, assists in walking
rehabilitation for stroke patients. It modifies the gait motion of
the wearer by applying torque to the hip and knee joints. Hybrid
assistive limb (HAL) [2] utilizes electromyography, which canbe used to estimate the intention of the muscle contraction to
apply torque to joints at the appropriate time. The rhythm assist
device [3] can be synchronized with the gait timing of the wearer
using a central pattern generator. This light device is designed to
fit the wearer.

Several wearable assist devices include an actuator thatprovides torque to each joint to assist in the wearer’s joint
motion [9]. Thus, although it is possible to adjust the assist force
and pattern to the wearer, the usability of such devices in dailyliving environments is limited because of its weight, and limited
operating time owing to the battery, the attach and detachprocesses, and other mechanical limitations. The other type of
wearable assist device—the supporter type—can also support
dorsal flexion by continuously supplying elastic force. Thus,
such devices are usually lighter than the actuated ones and canbe easily attached/detached. However, the assist pattern should
be constant, and thus, it potentially restricts the motion of theankle joint. The characteristics of such wearable devices arepresented for comparison in [Table 1].

The decrease in gait ability is determined based on the
combination of a reduction in balance sensation, muscle strength,and other functions. These typically result in a reduction in stride
length, range of joint motion, and foot clearance [4-6]. The
reduction in foot clearance and increase in the variation of thefoot clearance increases the risk of tripping. Such changes in gait
motion occur in the elderly owing to the decrease in the dorsal
flexion angle [7]. Thus, it is important to increase the dorsal

978-1-5386-1644-4/17/$31.00 ©2017 IEEE

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Owing to achieve profits of actuator type and supporter typedevices at the same time, we developed a muscle-forcetransmission type dorsal flexion assist device using string, which
utilizes a strong wearer's muscle as the power source instead of
actuators [10]. The string-type wearable assist device applies
torque to lower limb joints through the string fixed along the leg
portion [11, 12]. The tension was generated by pulling the string
using the upper body muscles. By optimizing the string path, theefficiency of the torque transfer was able to be improved, whichmeant that the torque applied to the ankle joint was maximizedin the swing phase although the collateral torque applied to theother joints was minimized.

Fig. 2 Theory of generating resultant force and torqueforce of the string 𝑷 and the distance vector from the ankle joints
to the string guide 𝒍 (shown in Fig. 2).
𝑻𝟏 = 𝒍 × 𝑷(1)The extension torque 𝑻𝟐 around the knee joints was thencalculated. The tension vector acted along the folded strings 𝒇𝒂
and 𝒇𝒃 , and the resultant force 𝑭, vectors 𝒓𝒂 and 𝒓𝒃 from thestring guide to adjoining guides, and the distance 𝒓 from theknee joints to the string guides were defined. These definitions
are shown in Fig. 2. Using these values, the extension torque 𝑻𝟐
is obtained from the following equation.

Fig. 1 Design of wearable assist device𝒓II.

𝒇𝒂 = 𝑷 × |𝒓𝒂 |DESIGN OF WEARABLE ASSIST DEVICE𝒂𝒓

A. Structure of the wearable assist device
The wearable assist device developed in our study consistedof cuffs and belts that were fixed to the subject's body and limbs.
Strings were used to connect the toe to the arm or body throughstring guides fixed on the cuffs and belts. The string transferredtension force from the upper body, which generates tension
force synchronously with the gait cycle, to the toe for assisting
the dorsal flexion during the swing phase. In this study, motionsof the upper body, such as the shoulder twist and arm swing,
were used for assistance in walking instead of an actuator. The
wearable device that was capable of realizing these functions
was manufactured as shown in Fig. 1. It consisted of four
components: toe supporter, knee supporter, harness, and arm
supporter. String guides, which is highlighted in Fig. 1, were
fixed at the appropriate position of the appliances using Velcro.
Because a part of the string consisted of rubber, the tension in
the string was able to be measured from the expansion of therubber portion.
B. Calculation of torque applied by the device
First, the normal gait motion of a healthy young adult on thetreadmill was recorded. Following vectors was calculated as astandard coordinate, the zero point was located on theexperimental floor. The torque applied to joints during walking
was then calculated. The dorsal torque 𝑻𝟏 acting around theankle joint was obtained from the cross product of the pulling

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(2)

𝒇𝒃 = 𝑷 × |𝒓𝒃 |(3)

𝑭 = 𝒇𝒂 + 𝒇𝒃

(4)

𝑻𝟐 = 𝒓 × 𝑭(5)𝒃

C. Design of string path
In order to avoid tripping due to decreased dorsal flexion
during the swing phase, the assist of the dorsal flexion iseffective in maintaining foot clearance. Usually, the footclearance is the minimum when the swing leg passes by the
stance leg. This minimum foot clearance (MFC) appeared in therange of 80% to 85% of the gait cycle (Fig. 3), as reported
previously [7]. MFC and dorsal flexion are illustrated in Fig. 4.
Thus, the dorsal flexion during this phase should be assisted toincrease the foot clearance. In addition, the collateral force
generated by the string tension should not disturb the natural gait
motion. Because the string guides were placed along the links ofthe entire leg, the string tension possibly also applies torque tothe other joints such as the knee.
Thus, the string path on the lower limb was optimized by
maximizing the dorsal flexion torque at the ankle joint at thetime of MFC and minimizing the maximum torque around theknee joint. The positions of the string guide were defined as thevariable parameters. The evaluation function 𝐽 was defined asshown in eq. (6). By minimizing 𝐽, the dorsal flexion torque atthe ankle joint 𝑇𝑎𝑛𝑘𝑙𝑒 was maximized and the torque at the knee
extension torque 𝑇𝑘𝑛𝑒𝑒 was minimized. In this simulation, the(a) Arm swing path(b) Body twist path
Fig. 5 String path candidates

process, the string path, along which the torque was generated at
the appropriate time, was identified heuristically.

Fig. 3 Joint angle and gait motion

As a result of the optimization process mentioned above, twostring paths were determined. The first string path was the arm
swing path (Fig. 5 (a)), which generated the force from the swingof the arm. This string path was designed to increase theelongation of the string greatly and to have the ability to adjust
the timing when the dorsal torque is generated by controlling thearm swing. The second string path was the body twist path (Fig.
5 (b)), which made use of the twist of the torso. This string path
was designed to generate the dorsal torque stably using the twistof the body, which generates the strongest force of the upper
body. Figs. 6 and 7 show the estimated torques generated by therespective string paths. It seems that the torque was generated
around the ankle joints in the range of 60% to 80% of the gait
cycles performing dorsal flexion. Although torque was
generated around the knee joint, the magnitude of the torque wasminimized.

Fig. 4 MFC and dorsal flexion

tension force was maintained at a constant value throughout the
gait cycle.
𝐽 = −𝑇𝑎𝑛𝑘𝑙𝑒 + 𝑇𝑘𝑛𝑒𝑒(6)III. WALKING EXPERIMENTThe string path from the upper limb, which should bedesigned to pull the string at the appropriate time, wasdetermined by simulating the changing pattern of the string-path
length during the gait cycle because the tension probably
increased in proportion to the path length. In the preliminary
experiment, markers were attached at the position where thestring guide was possibly fixed in order to record the relative
motion between markers during the gait cycle. By connecting
the positions of such markers, the string path, along which thestring was expected to stretch at the time of assist, wastentatively determined. The sum of the distances between themarkers selected as paths was defined as the string-path length.
The tension generated in the string was calculated from theelongation, which was defined as the extension of the string path
length, and the elastic modulus of the string. The elastic
coefficient of the rubber was 2.8 N/mm.

A. Apparatus
The motion of the subject was measured using a motion
capture system (MAC 3D System, Motion Analysis). The
position of the markers, which were fixed at the center of eachjoint of the subject, was projected in the sagittal plane. Each joint
angle was then calculated. The angles of the hip, knee, and ankle
joints were determined to be 0 deg in the upright posture. The
direction of the joint flexion or dorsal flexion was determined to
be in the positive direction. The elastic coefficient of the stringwas 2.8 N/mm. The tension applied to the string was calculated
from the extension of the string in a manner similar to the design
process of the string path. A treadmill was used to continue thewalking for several minutes.

Using these assumptions, the tension occurring in the string
was simulated, and the torque acting on the joints via the string
path was calculated using equations (1) and (5). Through thisB. Protocol
The experiment was performed with the permission of theinstitutional review board of the Nagoya University. Two1413

dorsal flexion, which confirmed that the change in dorsal flexion
was caused by the assist device. Furthermore, the stroke of thestring measured in the experiment was much smaller than that in
the simulation because the tension of the string not only affected
the ankle but also the upper body motion. Thus, the effective
torque acting on the lower limb joints was smaller in theexperiment.
Fig. 8 (a) shows the joint angle pattern of the assisted
walking with the arm swing path and normal walking. The
torque acting on the knee and ankle joints was also drawn in Fig.
8(b). As compared to normal walking, the dorsal flexion angleof the assisted walking became approximately 2 deg greater inthe range of 70% to 80% of the gait cycle, which was the timeof the assistance.
Fig. 8 (b) is compared with the simulation result shown in
Fig. 6. The torque pattern differed greatly in the range of 25% to
65% of the gait cycle; Fig. 8 (b) shows that the torque becamezero in the experiment during this phase, which means that thestring was sagging. Because the arm swing path generated force
from the swing of the arm, the wearer was able to adjust thetiming for generating the dorsal torque. Thus, the pattern of thestring extension, which generated tension, was probably
controlled by the subject.

Fig. 6 Simulated joint torque (arm swing path)

Fig. 9 (a) shows the comparison of the joint angle pattern of
the assisted walking with the body twist path and normal
walking. According to this graph, the assist increased the dorsal
flexion angle by approximately 10 deg during the assisted timein the range of 60% to 80% of the gait cycle. This result
indicated a decrease in the planter flexion, which meant a weakpush-off at the timing of the toe-off. This weak push-off
probably caused the approximately 8 deg of reduction in theknee flexion angle, which was larger than that of the arm swing
path.

Fig. 7 Simulated joint torque (body twist path)

healthy young male students participated in the experiment. The
mean ages of them was 22 years old. To evaluate the influenceof the wearable assist device on the gait motion, the gait motionsof the assisted trial, which consisted of arm swing path and body
twist path, were compared to the normal gait. At first, normal
walking was recorded to avoid the learning effect caused byassist devices. Then, assisted trials were recorded sequentially.
The order of the assisted trials was randomized. To eliminate thepossibility that subjects intentionally changed their gait, theelectromyograms of the tibialis anterior, gastrocnemius,
quadriceps femoris, and hamstrings were obtained. The walking
speed was determined for each subject and fixed through theexperiment. The time of the single trial was 3 min, of which thefirst 1 min was used for gait stabilization and the subsequent 2
min were analyzed. The measured data was normalized by the
gait cycle—which was determined as the period between
successive heel contacts—and statistically processed.
C. Experimental result
Because a similar trend was observed between the two
subjects, the result of one subject is mentioned in the following
analysis. In addition, it should be mentioned that there was nodifference in the electromyogram of the tibialis anterior during
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According to Fig. 9 (b), the extension torque was exerted at
the knee joint, which probably affected the joint angle. However,the value of the knee extension torque did not differ from that of
the arm swing path. Thus, the weak push-off probably decreased
the knee flexion angle. On comparing the estimated joint torqueshown in Fig. 9 (b) and the simulation result shown in Fig. 7, thepattern was found to differ in the range of 80% to 100% of the
gait cycle. The change in the gait motion, which meant a weakpush-off, resulted in the change in the time at which the torquewas generated.
The MFC was then compared for various path in Fig. 10. The
MFC of both the assisted trials became larger than that of normal
walking, which meant that the assist device was able to decrease
the risk of tripping. In addition, the assist of the arm swing pathworked more effectively to increase the MFC than the assist of
the body twist path. On the other hand, in the other subject, theMFC of the assist with the body twist path was smaller than thatof normal walking as shown in Fig. 11. Because this greatly
decreased the knee flexion angle, the height of the foot probably
decreased despite the increased dorsal flexion in the body twist
path. This result suggested that in order to maintain the knee
flexion, which was probably realized by adjusting the string path
not to prevent push-off, it is important to increase the MFC to
avoid tripping. Thus, such individual differences have to beconsidered to achieve stable assist performance.

(a) Comparison of joint angles(a) Comparison of joint angles(b) Generated joint torque(b) Generated joint torqueFig. 8 Comparison of normal walking versus arm swing path

Fig. 9 Comparison of normal walking versus body twist path

Fig. 10 Comparison of foot clearance

Fig. 11 Comparison of foot clearance of other subject

IV.

CONCLUSIONSIn order to prevent falls induced by tripping in the elderly,
we developed a non-actuated dorsal flexion assist device, whichapplied dorsal flexion torque to the foot from the upper bodyusing string. Using a simulation that estimated the torqueapplied to the lower limb joints via the designed string path, twostring-path patterns were determined. The first string path was
the arm swing path, which had a long stroke during the gait cycle
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and allowed the adjustment of the torque pattern by controllingthe arm swing. The second string path was the body twist path,
which was able to generate the dorsal torque stably using thestrong muscle. The assist performance was then evaluated bycomparing the gait motion between assisted walking and normal
walking. Although the two string paths exhibited different
characteristics, both paths successfully increased the dorsal
flexion and MFC.

However, the result obtained from one of the subjects
suggested that the weak push-off, which was caused by thecollateral torque owing to the string path, decreased the MFC in
the case of the body twist path through the reduction in the knee
flexion. Thus, it is necessary to improve the string path to reducethe collateral torque and fit individual differences. Furthermore,
the experiment will be conducted to many subjects, thedifference of the result between subjects will be analyzed.

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