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1

Knee Pain Patient Simulation for Recommendation of
Sit-to-Stand Handrail Positions
Kohei Yamakawa, Shogo Okamoto , Member, IEEE, Ryu Kubo, Naomi Yamada, Yasuhiro Akiyama ,and Yoji Yamada, Member, IEEE

Abstract—Estimating the physical burdens of people with motor dysfunction is helpful for developing care equipment and facilities. We propose a patient simulation where a trained healthy
subject simulates impaired movement with instruments to guide
his/her body motion with less possible incurred pain. Electric cutaneous stimuli are used to intuitively indicate possible pain of an
injured body part. We used this simulator to investigate a suitable handrail position for knee pain holders during sit-down andstand-up motions. Electric stimuli were presented in proportion to
the moment at the supposedly impaired knee. The results suggestthat a handrail on either the healthy or impaired side significantly
reduces the experienced physical burden (i.e., internal knee flexionextension moment). Specifically, the maximum knee moment was
smaller when the handrail was on the impaired side than it was on
the healthy side.
Index Terms—Knee osteoarthritis (OA), knee moment, motion
simulation.

I. INTRODUCTION
HE physical burden experienced by people with motion
impairment is fundamental information for designing careequipment and facilities. However, there are issues with performing experiments to collect data on such burdens undervarious conditions and with an adequate amount of trials forstatistical analysis, because of the painful experiences the participants may be subjected to. To solve this problem, we proposea patient simulation where the motions of a healthy person arecontrolled or guided to simulate the typical motions of patientsand estimate the burden.1
To date, there have been some attempts to use healthy people to simulate motion dysfunction. Aging simulation kits use
braces or weights to limit joint motions and mimic the inconvenience of the elderly with motion impairments in their daily

T

Manuscript received May 25, 2018; revised October 21, 2018 and December
17, 2018; accepted January 20, 2019. This work was supported by the basic
research fund of Nagoya University. This paper was recommended by Associate
Editor Kang Li. (Corresponding author: Shogo Okamoto.)
K. Yamakawa, S. Okamoto, R. Kubo, Y. Akiyama, and Y. Yamada are with the
Department of Mechanical Systems Engineering, Nagoya University, Nagoya
464-8601, Japan (e-mail:, yamakawa0510k@gmail.com; okamoto-shogo@mech.nagoya-u.ac.jp; bokuryu0825@gmail.com; akiyama-yasuhiro@mech.
nagoya-u.ac.jp; yamada-yoji@mech.nagoya-u.ac.jp).
N. Yamada is with the Department of Mechanical Systems Engineering,
Nagoya University, Nagoya 464-8601, Japan, and also with the Aichi Medical
College, Nagakute 480-1195, Japan (e-mail:,yamada-nao@yuai.ac.jp).
Color versions of one or more of the figures in this paper are available online
at http://ieeexplore.ieee.org.
Digital Object Identifier 10.1109/THMS.2019.2896487
1 See also discussion part in terms of necessity and possibility of patient
simulation.

lives. Pacala et al. [1] and Wood [2] reported the educational
values of such simulations. Their objectives were to let participants experience the inconvenience, rather than estimation of
the physical burden. Kubo et al. [3], [4] introduced a wearable
robot as a programmable joint brace. The brace interfered withthe motion of the wearer to resemble that of patients with knee
disabilities. Ullauri et al. [5] restricted the contraction of theanterior tibialis muscle by taping the lower thigh and inhibiting the dorsal flexion of the ankle to simulate the gait of theelderly. Joint splints have been used to investigate the effects
of a limited range of joint motion [6]–[8]. For example, Spierset al. [8] used an elbow splint to observe the compensative
motions caused by an elbow disability in daily lives. These approaches disturb the intended motion of healthy people or guidethem to unintended motion by using mechanical restriction or
external forces. However, they are not suitable for simulating
the voluntary and preventive motions of patients to avoid pain
during motion.
Unlike previous methods, we did not use braces or joint splints
to simulate patient motions. Instead, we used electric cutaneous
stimuli on the body part that was supposed to be diseased. The
electric stimuli were not at a painful level; however, they served
as warning stimuli and allowed healthy participants to voluntarily take preventive motions against possible pain. This approach
suits the simulation of the motion of knee osteoarthritis (OA)
patients. Knee OA is a major disease, especially for the elderly,
and approximately 10% of who suffer from this degenerative
disease [9], [10]. Knee OA is caused by various factors that comprise aging and causes articular inflammation and pain. Given
the pain during motion, knee OA patients move differently fromhealthy people. The characteristic motions of knee OA patientsare not due to limited joint motions but are voluntary motions
to prevent possible pain as well as muscle weakness and joint
deformation. Hence, braces or joint splints to mechanically limit
the motion are not suitable for simulating the motions of knee
OA patients.
We consider stand-up and sit-down motions as typical daily
activities during which knee OA patients may feel pain. Thus far,the influences of the height and location of chairs and handrails
during the sit-to-stand motion have been studied. For example, high chairs are known to reduce the knee moment during
this motion [11], [12]. Kinoshita et al. [13], [14] investigatedthe handrail height in experiments involving healthy adults andthe elderly and reported that higher handrails are more effective in terms of the reduction of lower limb joint moments and2168-2291 © 2019 IEEE. Personal use is permitted, but republication/redistribution requires IEEE permission.
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required time to complete the sit-to-stand motion where the useof handrails influenced lower limb joint and trunk angles as well
as resultant center of gravity. A bilateral handrail of which rightand left bars were at different heights was potent for healthyand disabled elderly people [13]. For healthy people, using a
unilateral handrail reduces the burden on the ipsilateral lower
body [15]. Although most of these studies involved healthy participants and focused on the location or usage of handrails, few
studies have targeted people with motion disabilities. Ishiguro
et al. [16] analyzed the motions of paraplegia patients who sat
down and stood up with the support of handrails. However, thepreventive motions of knee OA patients are distinct from thoseof paraplegia patients. As mentioned before, limited work has
been performed in understanding the configuration and role ofhandrails to aid in reducing symptoms and improving functionof unilateral knee OA patients; therefore, this requires further
investigation.
In general, the use of sticks on the healthy side during walking
improves the stability and preserves the impaired side. However,there have been no unified reports about the best handrail position for unilateral knee pain holders when standing up andsitting down. A handrail on the impaired side may properly balance the loads on the healthy and impaired sides and facilitatea stable sit-to-stand motion. On the other hand, the use of ahandrail on the healthy side may enable patients to fully lean on
the healthy side and decrease the burden on the impaired side.
Hence, placement of the handrail on either side may effectively
reduce the burdens on the symptomatic knee in patients withknee OA. In this study, we used a patient simulation with painwarning electric cutaneous stimuli to investigate which side of
the handrail lowers the physical burden on the supposedly affected knee for unilateral knee OA patients to sit down andstand up. Note that the objective of this study was not clinicalvalidation but rather the proposal of a new method of patient
simulation. Direct comparison between the simulated and actual
patients’ motions is necessary before fully accepting the resultsof the patient simulation.
II. APPARATUS
As shown in Fig. 1, the handrails and chair were at heights
of 725 and 450 mm, respectively. These are within the range ofJapanese Industrial Standards [17], accessibility standards [18],and those reported in [19] where data for human-friendly design based on experiments involving Japanese subjects were
publicized. It should be noted that chair height was adjusted forindividuals in some studies [20], [21], and such settings reducethe variation of initial conditions of lower limb joint angles. The
distance between the participant and handrail was 350 mm.
We used ground reaction force sensors for each foot and goniometers (K800, Biometrics, Ltd., Cwmfelinfach, U.K.) for the
ankles and knees on both sides. The ground reaction force sensors were embedded in shoes, and each shoe contained threeload cells (FX1901-0001-0050-L, TE Connectivity Measurement Specialties, Schaffhausen, Switzerland) located at the heel,
thenar eminence, and antithenar beneath a sponge sheet. The
force and joint angle data were collected by using a microcomputer at 1 kHz.

Fig. 1. Experimental setup. Goniometers and electric pads were attached toboth sides.

Fig. 2. Electrical stimulation to the skin at a supposedly impaired knee. Painwarning stimuli are controlled by the pulse interval of the electric stimuli.
(a) Circuit. (b) Pulse operation. (c) Electrodes at the knee. (d) Pulse interval asa function of the knee moment on the supposedly impaired side.

Fig. 2(a) shows a schematic circuit of the electric stimulator.
Three electrodes (two cathodes and one anode) were attachednear the knee with adhesive silicone pads. The anode was connected to a voltage amplifier (MHV12-470S06P, Bellnix Co.,
Ltd., Saitama, Japan, maximum current: 2 mA) by a photo relay (AQY210EH, Panasonic Corporation, Kadoma, Japan). The
relay was operated by a microcomputer at 1 kHz. When therelay was switched ON, a pulse current flowed to the skin to deliver the electric cutaneous stimulus. As shown in Fig. 2(b), we
controlled the pulse frequency while the maximum voltage remained constant, which easily prevented unintended activationof nocireceptors and painful experiences [22]. The pulsewidth
was constant at 100 ms and the interval between pulses wasvariable. The small pulse intervals led to frequent pulse stimuli

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YAMAKAWA et al.: KNEE PAIN PATIENT SIMULATION FOR RECOMMENDATION OF SIT-TO-STAND HANDRAIL POSITIONS

and transient current flows. In other words, smaller intervals led
to more prominent cutaneous stimuli [22].
As shown in Fig. 2(c), the stimuli were presented to the medial part of the knee and this position did not vary substantiallyacross the participants. This position was determined for individual participants such that it clearly evoked the sense of skinstimulation but did not cause muscle activation. Some may feelthat the part just above patella is a suitable position; however,electrical stimulation was rarely felt at this part.
III. EXPERIMENT: SIT-TO-STAND MOTION WITH A HANDRAIL
In this experiment, we simulated the motion of people impaired by unilateral knee pain, which clearly differs from thatof healthy people.
A. Participants
Five healthy university students (mean height 174.0 ± 2.8 cm,
weight 60.7 ± 3.8 kg) took part in the experiments with written
informed consent. All participants were right handed and declared that they had no recognizable symptoms or disabilities intheir lower limbs. They had no experience with our apparatus in
the past.
B. Tasks
The participants practiced the typical motions of knee OA
patients [23], [24] before the main tasks. These motions weresupervised by an experienced physical therapist and included
the following: slow motion due to a fear of pain, body inclination
to the healthy side, concentration of the foot load on the healthyside, limited knee joint flexion (up to ∼ 100o ), and avoidanceof full knee extension. Furthermore, participants were informedthat large burden on the supposedly impaired knee would cause
the electrical cutaneous stimuli as pain-warning signals andthey should avoid such states. This practice session was only
4–5 min; nonetheless, it is indispensable to simulate theimpaired motion. Otherwise, participants would not recognize
the meanings of the warning electric stimuli and rectify theirmotions.
Note that the participants had to be familiar with the abovementioned typical impaired motions for simulation. In addition,
electric stimuli guided their motions based on the physical statesof their body postures and loads. This combination of practiceand computer-controlled guidance was to ensure that the motion
was simulated as designed. In the later data analysis, outlier trials
that did not match the above-mentioned typical motions were
excluded.
Participants repeated the simulated sit-to-stand motion ten
times under each of three conditions: no handrail, handrail on
the impaired side, and handrail on the healthy side. In order
to investigate the laterality of knee impairment, each participant
alternatively simulated impairment on the right and left sides. In
total, each participant performed 60 trials (three handrail conditions × two lateral conditions of impairment × 10 repetitions).
The order of the handrail condition and side of impairment was
randomized. The participants grabbed the handrail at a position
just beside their bodies at handrail conditions. They lay their3

feet at their comfortable positions shoulder width apart. They
were not allowed to move foot positions during a single trial.
For each participant, the entire experiment, including the explanation, practice, and recess took approximately 3 h and was
conducted on the same day. They could take a rest when theywanted.
C. Electrical Stimulation
Some previous studies have considered the adduction moment of the knee joint as a predictor of the pain experienced byknee OA patients [25], [26], whereas others showed inconclusive results [27], [28]. The adduction moment is prominent forpatients with a deformed knee joint and could not be applied tothe healthy participants of this study. Instead, we used the internal flexion-extension moment of the knee in the sagittal plane,
which has been used as an index of the physical burden during sit-to-stand motions and stair climbing [13], [29], [30]. We
varied the interval of the electric pulses inversely in accordancewith the knee flexion-extension moment, as shown in Fig. 2(d).
As a result, more intense cutaneous stimuli were presented witha greater knee moment.
To leverage the range of nonpainful electric stimuli, we set
the minimum interval of the pulse stimuli to be 50 ms above
the knee moment normalized by the weight of the participant:
0.6 Nm/kg. We determined this value referring to one study
where the maximum knee moment of knee OA patients during
sit-to-stand motion was 0.51 Nm/kg on average [23]. The knee
moment of the supposedly affected side rarely exceeded thisvalue in our later experiment. To simulate the pain upon motion,
electrical stimuli were presented only during motions with aknee angular velocity that was greater than 0.1 rad/s. Hence,the stimuli were not produced when a participant rested on achair or stood still. Each participant adjusted the voltage of theamplifier to ensure that the electrical stimuli were not painful
before the experiment.
D. Data Analysis
Based on the recorded ground reaction forces, knee moments,and videos, we excluded invalid trials that did not reflect typicalmotions of knee OA patients. Body weight and knee moment
concentration on the healthy side were used for screening out
trials. According to Turcot et al. [23], during sit-to-stand motions of knee OA patients without handrails, the maximal footload on the impaired side was 10% (±14% standard deviation)
smaller than that on the healthy side. Another study reported asimilar value of asymmetry: 13% [31]. Also, the maximal knee
flexion-extension moment of the patients was approximately
70% of that of the healthy [23]. Considering the natural deviation in these values and the difference in settings between ourstudy and the studies of Turcot et al. [23] and Christiansen andStevens-Lapsley [31], we excluded trials with a weight ratio
beyond 130% – 50% (mean ± 3 times the standard deviation
reported in [23]) or where the ratio of knee torques was out of
the range 100% – 40% (mean ± 1.0 – 1.5 times the standard
deviation reported in [23]). Both ratios were calculated by dividing the value of the impaired side by that of the healthy side.

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IV. RESULTS

Fig. 3. Example of knee angles, foot loads, and knee internal flexion-extension
moments of both legs when right-side impairment was simulated. The knee
angle was 0o at the full extension. The bottom figure shows the change in pulseinterval for the electrical stimulus.

On average, three to four trials (5% – 7% of all trials) were
judged invalid for each participant.2
As an example of simulated motions, Fig. 3 shows the knee
angles, ground reaction forces, flexion-extension moments, andpulse interval when a participant simulated the sit-to-stand motion without using a handrail. The change in the pulse interval
has a zigzag shape because the pulse interval was updated only
when the pulse stimulus was generated. The figure shows the
typical lateral imbalance of knee OA patients. The impaired
knee did not fully extend, and the foot load and knee momenton the impaired side were smaller than those on the healthy side.
The pulse interval decreased during the stand-up and sit-down
motions, which led to a clear cutaneous stimulus.
As indicators of physical burdens, we computed the maximum and integral values of the moment of the impaired kneefor each of the sit-down and stand-up phases. The integral values
were computed while the knee angle of the supposedly healthyside was between 5◦ –95◦ . Both values were normalized by themass of the participant. We then applied three-way analysis ofvariance (ANOVA) on each of the maximum and integral values
with the following factors: the side of the simulated knee (rightor left), motion (sit-down or stand-up), and handrail condition(no handrail, healthy side, or impaired side). For this analysis,
the maximum and integral moment values were averaged withineach individual for each condition.
2 We also conducted another outlier analysis based on Mahalanobis’ distanceson the two-dimensional space of the ratios of maximum knee moments and footloads for each condition. The 4% of trials for which distances from the centroids
were greater than those of the others were fully included in the trials excluded
from the methods presented in the main sentence.

With respect to the maximum knee moment, the side of
the simulated knee (F (1, 48) = 10.6, p < 0.01) and handrail(F (1, 48) = 42.6, p < 0.001) had significant effects, whereas
the motion (F (1, 48) = 0.44, p > 0.05) and interaction effects were insignificant. The use of handrails significantly decreased the maximum knee moment. We then performed post
hoc three-way ANOVA to compare the two handrail conditions (healthy side or impaired side). The side of the simulated knee (F (1, 32) = 19.1, p < 0.001) and handrail condition(F (1, 32) = 4.3, p < 0.05) also had significant effects, whereas
the other factors had insignificant effects, indicating that the
maximum knee moment values were lower when the handrail
at the impaired side was used. The results of the three-way
ANOVA are summarized in Fig. 4 (upper row) with the means
and standard errors across the participants for the maximum values of the knee moment on the supposedly impaired side underindividual conditions.
With respect to the integral values of the knee moment, theside of the simulated knee (F (1, 48) = 31.2, p < 0.001) andhandrail condition (F (1, 48) = 33.6, p < 0.001) had significant effects, whereas the motion (F (1, 48) = 0.14, p > 0.05)
and interaction effects were insignificant. The use of handrails
significantly decreased the integral value of the knee moment.
A post hoc three-way ANOVA was performed to compare thehandrail sides; the side of the simulated knee had a significant
effect (F (1, 32) = 26.1, p < 0.001), but the handrail position
did not (F (1, 32) = 0.25, p > 0.05). These results are summarized in Fig. 4 (lower row) with the means and standard errorsof the integral moment values.
Collectively, the maximum and integral knee moments on the
impaired side decreased when a handrail was used on either thehealthy or impaired side. The maximum values were smallest
when the handrail was placed on the impaired side, whereas theintegral values were not significantly influenced by the handrail
position. Furthermore, the handrail was equally effective for thesit-down and stand-up motions in terms of reducing the knee
moment.
The variation of knee moments was small within individual
participants. The ratio of standard deviation to the mean of
the maximal knee moments were 10.0%–26.4% for all of the
combinations of participants and conditions. For example, themean and standard deviation of ten trials was 21.2 ± 5.6 Nmfor one participant whose variation of the knee moments waslargest during stand-up motion with a right handrail and left
simulated knee. Similarly, for the integral value of the knee
moment, the ratios of the mean and standard deviation ranged
13.0%–30.6%.
The asymmetry of the maximum foot loads was 82% ± 8%
(standard error) on average across the participants for nohandrail condition. The load on the supposedly impaired sidewas 18% smaller than that on the contralateral side. These
values were 75% ± 9% and 70% ± 9% when the handrail
was placed on the healthy and impaired sides, respectively.
Similarly, the asymmetry of the maximum knee flexionextension moments was 77% ± 10% for no handrail condition.

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Fig. 4. Maximum and integral values for the flexion-extension moments of the impaired knee. The means and standard errors among participants are presented.
Asterisks * and *** indicate significance levels of 0.05 and 0.001 (ANOVA), respectively.

Fig. 5. Handrail function during sit-to-stand motions for a unilateral knee-pain
holder. Upward arrows indicate ground reaction forces.

Furthermore, they were 73% ± 10% and 68% ± 11% for thehandrail on the healthy and impaired sides, respectively.
V. DISCUSSION
Using a handrail naturally reduced the knee flexion-extension
moment because the knee moment is largely determined by thelower limb joint angles in the sagittal plane and ground reaction
force to the foot, and the handrail partially supported the weightof the participants. Our main interest was in determining whichhandrail side is more effective for unilateral knee pain holders.
The experimental results suggest that either the healthy or impaired side is effective. In terms of the maximum knee moment,
having the handrail on the impaired side more effectively reduced the burden. Ahead we speculate a possible explanation of
the results.
When the handrail was on the healthy side, the center ofgravity approached the edge of the main supporting area, as
shown in Fig. 5(a), which was an unstable state. The participants may have tried to avoid such an unstable state and loaded
the impaired side, which prevented the knee moment on the
supposedly impaired side from decreasing. In contrast, when
the handrail was on the impaired side, a greater supporting area
was formed by the healthy foot and handrail. The center of gravity might have been located near the central part of the area, as

shown in Fig. 5(b). This enabled the participants to stably sit
down and stand up. The participants then did not need to depend on the impaired foot, and the knee moment on the impairedside remained small. Observation of the center of gravity during
trials, which was not within the scope of this study, will help
validate these explanations.
Unlike the maximum knee moments, we did not find a significant difference in the integral moment values between two
handrail conditions. The results do not deny their potential differences, and the difference between the two handrail conditions might have been diminished because the time requiredfor completing sit-down and stand-up motions was short when
the handrail was used. The integral values are sensitive to theinterval for computation.
The side of the simulated impairment was found to have asignificant influence on the knee moment. The knee moment
was greater with a left-side impairment than a right-side impairment. This may be related to the foot dominance. Some studies
have reported that dominant or preferred legs are suitable formanipulative tasks, and contralateral legs play a major role insupporting legs during balancing or stabilization tasks [32], [33].
Supporting legs are not necessarily stronger than nondominant
ones in terms of muscle forces but tend to mainly be used forsupporting or weight-bearing tasks [32], [33]. Given that all participants reported that they use their right legs to kick a ball,3
they may have relied on their left legs during the sit-to-standtasks in our experiments. Hence, even when their left legs were
designated as impaired, they tended to load the left leg. However, Velotta et al. [33] reported that dominance by the right andleft legs of 22 participants was equally observed for balancing
tasks, whereas most of them declared that their dominant leg
was on the right. Although leg dominance may be a possible
factor, it cannot be concluded that leg dominance explains theimbalance effect of the simulated side in our experiment.
3 This is not a complete test but a concise questionnaire for determining thedominant leg.

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IEEE TRANSACTIONS ON HUMAN-MACHINE SYSTEMSIt should be noted that we focused on limited features ofcharacteristic motions of knee OA patients. For instance, their
upper trunk motions were also distinct from those of the healthyin earlier studies [23], [34]. Patients leaned their trunks forward
such that their center of mass was close to the base of support
formed by feet. In this study, such upper trunk motions were notmeasured, and we did not instruct the participants to take suchmotions.
Although the same phenomena are likely to hold for actual
patients, we cannot fully validate the patient simulation before
comparing the results of the simulated patients and actual patients. Of course, only the minimum data should be collectedfrom the latter. When collecting the motion data from actual
knee OA patients, they should be asked to make comfortablemotions without pain. Hence, they do not provide motion data
under exhaustive conditions, which makes it difficult to compare differences among the conditions. Hence, experimental
validation of patient simulation by comparing clinical and simulated data remains to be studied. Computational generation of
motion [35], where human motions are optimally computed interms of, for example, the energy criteria, can be an alternative toclinical experiments involving actual patients. Nevertheless, sitto-down motions of knee OA patients are not biomechanically
optimal [34]. Furthermore, in some studies involving healthy
participants, the grabbing position such as height and rear-front
position of the handrail was a factor for the physical burden
during the motion [14], [16], [29]. In this study, the participants
grabbed the handrail at the point just beside their bodies. Further
experiments are required to generalize the effects of handrails.
VI. CONCLUSION
In this study, healthy participants simulated the typical sitto-stand motions of people with unilateral knee OA in a controlled environment. Unlike previous studies, wherein braces
or splints were used to mechanically limit human motions, thisstudy guided participants’ motions using a pain-warning indicator, which was an electrical cutaneous stimulus proportional tothe physical burden on the supposedly impaired knee. Hence, themethod was suitable for simulating patients’ preventive motions
to avoid the motions to cause pain. By using this method, weinvestigated the effective side of the handrail for the sit-to-standmotion of knee pain holders. Both the impaired and healthysides were equally effective at reducing the maximum and integral knee moments of the supposedly impaired side during the
motion. For reducing the maximum moment, it was suggested
that a handrail on the impaired side is more effective than on thehealthy side. To be conclusive, these observations through patient simulation should be compared with experiments involving
actual knee OA patients.
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Kohei Yamakawa received the B.S. degree in engineering from from Nagoya University, Nagoya,Japan, in 2018.
He is a mechanical engineer.

Shogo Okamoto (M’07) received the Ph.D. degree in information sciences from Tohoku University, Sendai, Japan, in 2010.
He joined Nagoya University as an Assistant Professor in 2010, and is currently an Associate Professor with the Department of Mechanical Systems
Engineering. His research interests include haptics,
rehabilitation engineering, affective engineering, andassistive robotics.

7

Ryu Kubo received the B.S. and M.S. degrees inengineering from Nagoya University, Nagoya, Japan,
in 2015 and 2017, respectively.
He is a Mechanical Engineer. His research interest
includes medical engineering.

Naomi Yamada received the Master’s degree in rehabilitation science, in 2016, from Nagoya University, Nagoya, Japan, where she is currently working
toward the Ph.D. degree at the Department of Mechanical Systems Engineering.
She is also a physical therapist and an Assistant Professor with Aichi Medical College, Nagakute,
Japan.

Yasuhiro Akiyama received the B.E. degree in engineering from the Tokyo Institute of Technology,
Tokyo, Japan, in 2006, and the M.S. and Ph.D.
degrees in engineering from the University of Tokyo,
Tokyo, Japan, in 2008 and 2011, respectively.
He has been an Assistant Professor with the
Graduate School of Engineering, Nagoya University,
Nagoya, Japan since 2016. His main areas of researchinterest are mechanical safety, gait assist, and manned
space mission.

Yoji Yamada (M’85) received the Ph.D. degree fromthe Tokyo Institute of Technology, Tokyo, Japan, in
1990.
He has been with Toyota Technological Institutesince 1983. In 2004, he joined Intelligent Systems
Research Institute of AIST as a group leader. In 2009,
he moved to Nagoya University as a Professor. He iscurrently studying safety and intelligence technology
in human/machine systems and robotic sensing andcontrol for them.