Development of a Lower Extremity Exoskeleton Robot with aQuasi-anthropomorphic Design Approach for Load Carriage

Donghwan Lim1, Wansoo Kim1, Heedon Lee1, Hojun Kim1,Kyoosik Shin2, Taejoon Park2, JiYeong Lee2, and Changsoo Han2∗

Abstract— This study developed the Hanyang ExoskeletonAssistive Robot (HEXAR)-CR50 aimed at improving musclestrength of the wearer while transporting a load. The developedexoskeleton robot HEXAR-CR50 has 7 degrees of freedom(DOF) for one foot, 3-DOF for the hip joints, 1-DOF for theknee joints, and 3-DOF for the ankle joints. Through functionalanalysis of each human joint, two DOFs were composed ofactive joints using an electric motor developed in an under-actuated form with heightened efficiency. The rest of the DOFswere composed of passive or quasi-passive joints to imitatehuman joints. The control of the exoskeleton robot was basedon the physical human-robot interaction. In order to verify theperformance of the developed HEXAR-CR50, muscle activitywas measured using electromyography, vGRF was measuredusing F-Scan sensor. The experimental results showed that%MVIC was reduced against the external load applied, whileGRF had a decrement rate, compared with the external loadwhen the exoskeleton was worn, which verified the performancein accordance with the development objective of load carrying.A muscle strength augment effect from the developed wearablerobot was verified.

I. INTRODUCTION

An exoskeleton robot system is a mechanical structurethat is attached to the exterior of a human body to improvethe muscular power of the wearer. Exoskeleton robots havebeen actively studied in the US, Japan, and Europe since the1990s , and research is currently being conducted to identifytheir application in various industries, including the military,medicine, and rehabilitation. The exoskeleton is dividedinto either a power assistance or a power augmentationsystem, depending on its purpose. And, depending on itsdesign structure of the alignment to the rotation axis of therobot joint and the rotation axis of the human joint, theexoskeleton may be classified as anthropomorphic, quasi-anthropomorphic, or non-anthropomorphic [1].

The power assistance system adds power to the wearer’sexternal skeleton because it supports the human body viathe exoskeleton. The MIT, HAL, ALEX, and HARO ex-oskeleton systems have been developed as supplementarysupport for the elderly and disabled persons to enable normaldaily living. As mentioned, these exoskeletons are designedanthropomorphically. In addition, the EKSO exoskeletonsystems is designed quasi-anthropomorphically[2]-[6]. Onthe other hand, the power augmentation system amplifiesthe power of the wearer through the exoskeleton, thereby

1Department of Mechanical Engineering, Hanyang University, Seoul,South Korea blackfire857@gmail.com

2∗Department of Robot Engineering, Hanayng University ERICA, Ansan,South Korea cshan@hanyang.ac.kr

enabling high-load carrying motions, which would otherwisebe difficult. The BLEEX, HULC, and XOS exoskeletonsystems were developed with the aim for use in industrialand military campaign environments, where heavy loads arecarried or manipulated. These exoskeletons are designedquasi-anthropomorphically. In addition, the WAD exoskele-ton systems was designed non-anthropomorphically[7]-[10].

In this study, we propose a lower exoskeleton systemdesigned to carry a heavy load. We analyzed the humanjoint function to design the joint modules of the exoskeletonrobot. We designed an exoskeleton robot through a quasi-anthropomorphic architecture. This paper focuses on thesetup of the mechanical structure and the design of theHEXAR-CR50 through the functions of each joint. In ad-dition, we confirm the effectiveness of the exoskeleton robotthrough experiments.

II. MECHANICAL DESIGN OF EXOSKELETON ROBOT

A. Structure of Exoskeleton Robot

The exoskeleton structure is placed on the human wearer,and it generates a moment according to the payload. In orderfor the generated moment not to affect the wearer, it isnecessary to maintain a constant rigid state in the stancephase. To maintain a constant rigid state, some robot jointsshould be misaligned with the human joints, thereby deliver-ing payloads to the ground. Thus, a quasi-anthropomorphicarchitecture was selected for efficiency and human motionsynchronized controllability to achieve the load carrying goalof the HEXAR-CR50. Fig. 1 shows the arrangement of theDOF selected for each joint and joint type. We developedan under-actuated exoskeleton consisting of 7 degrees offreedom(DOF) per leg. Based on the gait power analysis,the hip and knee extension/flexion joints were selected tobe active. And Ankle dorsiflexion/plantarflexion joints wereselected to be quasi-passive mechanisms.

The remaining joints were selected to be passive withmechanical limitations because range of motion(ROM), mo-ment, and power were low.

The HEXAR-CR50 is of an under-actuated type andapplies a quasi-anthropomorphic mechanism in which someof the robot’s DOFs are different from human DOFs in orderto achieve the goal of load carrying. In addition, the goalweight of the design was 23 kg, and the system was designedto carry a payload of 20-30kg. Before designing a joint oractuator, information on the desired range of motion for eachjoint is needed. To avoid potential harm to the operator, therange of motion in the exoskeleton joint should be limited

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Fig. 1: Exoskeleton structure design and Joint Modules

to less than that of human flexibility. Table I shows a ROMwhich is set to the design of the HEXAR-CR50.

TABLE I: Ranges of motion for the joints of HEXAR-CR50

Joint axis Angle(deg)

Min. Max.

HipInternal/external rotation -10 10Extension/flexion -40 150Adduction/Abduction -7 35

Knee Extension/flexion 0 127

AnkleInternal/external rotation -10 10Plantarflexion/dorsiflexion -30 30Eversion/inversion rotation -10 10

B. Design of Joint Modules

1) Active Joint Module: Several assumptions were madein order to select and size the actuator. The first assumptionwas that the size, mass, and inertial properties of the ex-oskeleton were equivalent to those of a human. The secondwas that the human and exoskeleton joints move along thesame axis. The third was that the maximum external loadspermitted were 30 kg.

The joint speed required for hip joints was a maximum of4.6 rad/sec, and -2Nm/BW of torque was required for levelwalking at 6 km/h with 30kg. The joint speed required forknee joints in the same environment was a maximum of 7rad/sec. However, most motions are done within 4 rad/sec,and motions faster than 5 rad/sec are done momentarily.

The active joint module refers to a module consisting ofan electrical motor and harmonic gear. It is applied to thehip and knee extension/flexion joint. The selected motorsare pancake-shaped, and a component-type harmonic drivewas selected in order to minimize the weight and widthof the joint module. The BLDC’s motor specifications canbe selected according to the characteristics of current andtorque. The motor produces the following torque, which isproportional to the motor current i, such that τ = kt · i,where kt is the motor torque constant. For the steady-state

Joint Velocity [rad/sec]0 50 100 150 200 250 300 350 400 450 500

Join

t Mom

ent [

Nm

]

0

0.5

1

1.5

2

2.5Level WalkStair AscentStair DescentWinding LineCurrent Line

(a) Hip joint velocity vs. moment

Joint Velocity [rad/sec]0 100 200 300 400 500 600 700

Join

t Mom

ent [

Nm

]

0

0.5

1

1.5

2

2.5Level WalkStair AscentStair DescentWinding LineCurrent Line

(b) Knee joint velocity vs. moment

Fig. 2: The required motor torque for the hip and knee ascalculated from the gait analysis data for walking

torque-speed relationship, the voltages around the loop aresummed and rearranged for i by the following equation:

i =V − vbR

=V − kvω

R(1)

where vb is the motor back-EMF in the armature, V is thevoltage applied to the motor, kv is the back-EMF constant,R is the motor’s resistance, and ω is the angular velocity.Substitute this value of i into the following equation for themaximum theoretical torque:

τmax =ktR

(Vmax − kvω) (2)

Therefore, the required torque must be less than τmax

from the motor winding [7]. The final limit of the motor‘scapabilities is the current limit; Ip is based on the maximumcurrent density in the winding and is available for a maxi-mum duration of 10 seconds. If the lowest of these is Ip andthe maximum torque τp, the winding can produce the torqueat any motor speed. It is calculated as follows :

τp = ktIp (3)

The required joint torque τreq and speed ωreq are trans-lated to the absolute values. With a harmonic drive gear ratioof N , the required motor torque T and ω are calculated by

T =TreqN

(4)

ω = ωreqN (5)

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Joint Angle [°]-30 -20 -10 0 10 20 30 40

Join

t Mom

ent [

Nm

/BW

]

-0.5

0

0.5

1

1.5

2

Level WalkStair AscentStair DescentLv-SpringStAsc-SpringStDesc-Spring

Fig. 3: Ankle dorsiflexion/plantar flexion motion analyses forselecting the spring constant kh during level walking, stairascent, and stair descent

Fig.2 shows that the required motor torque-speed charac-teristics for the HEXAR-CR50’s joints were calculated andplotted for various walking environments. Stair ascent anddescent, with the winding torque line, were lower than thecurrent torque line. However, the required torque-velocitycurve of level walking for the knee joint exceeded the short-term operation limit of the actuator in the transition fromextension to flexion. In order to satisfy the required torque,only upgrading of the motor should be needed.

However, this research did not consider this aspect becauseof the short timeframe available for actuating the motion.The parameters of the electrical motors that were selected topower the flexion joints of the HEXAR-CR50 are shown inTable II.

2) Quasi-passive Joint Module: Ankle dorsiflex-ion/plantarflexion joint absorbs energy during the firsthalf of the stance phase and releases energy just beforetoe-off. Power is generated by the moment and jointvelocity, whereas the ankle joint moment is not generatedsignificantly. That is, changes in the joint moment aresmall, but large joint velocity is required. Furthermore,the region where ankle power is applied is about 10%of the gait cycle, which is inefficient since instant poweris generated to control the active joints. To satisfy thisfunction, quasi-passive mechanisms were designed with

TABLE II: Electric motor and harmonic gear selected forHEXAR-CR50

Parts Parameters Value Unit

Motor

Motor torque constant,kt

0.061 Nm/A

Input voltage, V 29.4 VNominal speed 4601 RPMNominal torque 0.429 NmNominal current 7.39 ABack-EMF constant, vb 6.39 V/kRPMMotor resistance, R 0.615 ΩCurrent limit, Ip 30.2 A

Harmonic gear

Transmission ratio, N 100 -Continuous torque 34 NmStall torque 57 NmEfficiency 90 %

Fig. 4: Design of the quasi-passive mechanism for the ankledorsiflexion/plantarflexion joint

compliance material to support the weight of external loadsin the stance phase and to absorb or release the energy.

The joint used for dorsiflexion and plantarflexion demandsa large moment during the gait. Power absorption andgeneration occur repetitively, according to the gait phase[13].Fig.3 shows the results of the gait analysis for the angle-moment relationship during level walking and stair climbing.

The stance phase may be divided into the dorsiflex-ion section, where power absorption takes place, and theplantarflexion section, where power generation occurs. Thisstudy proposes a mechanism in which the power storedduring the dorsiflexion motion by using a spring is usedduring plantarflexion. The approximate spring constant canbe known from the relationship between angle and momentin the results of the gait analysis. When the body weightof the exoskeleton is 23 kg, the mean value of the springconstant kh is about 0.068 Nm/BW·deg with three differentwalking conditions. The quasi-passive module refers to apassive joint that can generate a force by applying springelements; it is applied to the dorsi-plantarflexion motion ofthe ankle joint. The spring constant kh to generate the anklepropulsion torque is analyzed in Table III.

The spring constant (ks) of the spring used for thismechanism may be calculated using (6). Here, τh is theconversion of the results of the gait analysis to the bodyweight of the exoskeleton. It can be expressed as τh = kh ·θr.Consequently, the spring constant was determined to be69.6N/mm. Fig.4 shows the design of an ankle mechanismfor the dorsiflexion and plantarflexion motions.

ks =khθr∆xll2

(6)

where, ∆xs = l2tan θr (0 ≤ θr ≤ 30)

τh = τS , khθr = ks∆xsl2

III. EXOSKELETON ROBOT CONTROLLER

A. Controller Architecture

The control law relies on the real-time model of thesystem, and kinematic and dynamic data must be contin-uously monitored. Kinematic variables require joint angles,

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TABLE III: Ankle spring constant kh for the plantarflexionmotion while walking

Spring constant (Nm/BW·deg) Mean

Level walking Stair ascent Stair descent0.0758 0.0537 0.0747 0.068

velocities, and angular acceleration that can be measured andestimated by the rotary encoders. The absolute foot angle isalso needed to measure the orientation of the upper bodywith respect to gravity, using an attitude heading referencesystem (AHRS). To select the gait control at each phase,an insole sensor was developed as well as to detect thegait phase through the center of pressure (COP). In thestance phase control, to measure the human-robot interactionforce, a multi axis force/torque sensor was mounted in theharness module between the human and robot. Finally, thetorque applied by the actuator was measured by the torquesensor, and it is used for both joint force feedback controland human-robot interaction force estimation. The controllerstructure is composed of a main controller, a sensor controlunit (SCU), and a motor control unit (MCU). To implementthe control algorithm, the main controller communicates withthe other control units to obtain the sensor signal and send thecommand. The SCU is responsible for acquiring the sensorsignal and transmitting it to the main controller by convertingit to a communication signal. The MCU performs the jointfeedback control and sends commands to the actuator.

B. Control Law with Human-Robot Interaction Force

As discussed in the previous section, each leg has twoactuated DOF: the hip and knee joints in the sagittal plane. Inorder for the leg to walk with a human motion, we considertwo distinct gait phases and select two different controlmodels. Fig.5 shows the block diagram of the control lawfor the exoskeleton leg. The human-robot interaction force,Tint, affects the exoskeleton by distinguishing the positionwith impedance, kh. The main tracking objective of thiscontrol law that was Tint going to zero. The gait phasecould be detected by the insole sensors and by switchingthe control laws for the stance or swing phase. In the stancephase, the control law uses a sensor that measures the human-robot interaction force, FintST = [FxFyFz]

T , at the harnessmodule. The joint torques caused by the human, T

′

intST , arecomputed in (7).

~T′

intST = JT ~FintST (7)

In swing, the inertial reference frame can be arbitrarilyselected to be the harness module. The estimate of theTintSW through the virtual control law is thus as follows(8).

~T′

intSW = B (q) q + C (q, q) q +G (q) − ~Ta (8)

These joint torques can now be used in the force controlscheme, and they can added to the gravity compensation

Fig. 5: HRI Based Control law for the Exoskeleton Leg.

torque, Tg . These desired torque, Td, was used in joint torquefeedback control with actual torque, Ta.

IV. EXPERIMENT AND DISCUSSION

The exoskeleton HEXAR-CR50 developed in this studywas designed to achieve power augmentation by reducing theload bearing on the human body during load-carrying mo-tions. That means to develop a exoskeleton system that canassist individuals who must carry loads for extended periodtime. During human walking, a substantial amount of powerconsumption can measure by the metabolic power rate. Manyresearchers have attempted to reduce the metabolic burdenon load carriage by developing exoskeletons [11], [12]. Toverify the metabolic power effect, the forces generated bythe human or the forces directly applied to the human toachieve motions were measured and analyzed. However, thedata that can be measured through human are extremelylimited. Among them, EMG sensors are generally used tomeasure muscle activity during movement. They are alsoused to represent muscle usage [13]. If muscle activity duringmovement is measured using EMG, it can be utilized inthe verification of the exoskeleton effect because it candemonstrate muscle usage. In addition, the forces applied toexternal load objects are transmitted to the ground throughthe human body so that the forces caused by the load canbe analyzed by measuring the vGRF between the humanand the ground. Force sensors of the insole type can beused to measure vGRF. The effect on the exoskeleton can beobserved separately because the sensors are located betweenthe human and the shoes while the forces are measured. Thus,

Fig. 6: Experimental setup of the power augmentation effectin level walking.

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two sensors were used to verify the metabolic power effecton load carrying by the exoskeleton.

A. Experimental Setup

The experiment was conducted in one subject (age: 29,weight: 75 kg). One subject was healthy and had no gaitimpairment. The EMG sensor used is a four-channel surfaceEMG (LAXTHA). The sensor used for the vGRF mea-surement was F-Scan in the shoe sensor of EMG. Theperformance of the robot was verified through the gaitexperiments using the above two sensors. The experimentalresults indicate the exoskeleton wearing and one without ex-oskeleton wearing by measuring sensor signals and comparedwith two experiment results. The experimental conditionswere external load conditions were 10, and 20 kg whilewalking on level ground at a constant speed (3 km/h).All experiments repeated five sets in which the externalload condition was changed sequentially. The experimentalconditions and procedures are shown in Fig. 6.

EMG electrodes were attached to the four main musclesused while walking: vastus lateralis and medialis, gastrocne-mius, and soleus. The sampling frequency of the EMG sensormeasurement and F-scan sensor with Tekscan devices wereat 512 Hz. The frequencies of two devices were matchedthrough a trigger because they were measured by their ownsoftware. To express EMG sensor signals as quantitativemuscle strength, the normalized technique, percent maxi-mum voluntary isometric contraction (%MVIC) was used.Equation(9) shows the calculation of %MVIC [14].

%MV IC =RMSmuscle

MV IC(9)

(a) Attachment positions of the EMG electrodes in themuscle

(b) Attachment positions of the F-scan sensor be-tween the human foot and foot module

Fig. 7: Sensors positions of EMG and F-scan sensor

where RMSmuscle is the root mean square (RMS) valueof the measured signal, and %MVIC is the maximum musclestrength. %MVIC is represented for each muscle, so %MVICwas measured accordingly. The method of measurementdiffered according to the locations of the muscles. Becausethe maximum muscle was measured, the measurement wasconducted immediately after the electrodes were attachedand re-measured when the electrodes were changed. Fig.7(a)shows the sensor settings in the experiment, the muscles at-tached to electrodes, and the %MVIC measurement method.Fig.7(b) shows the F-scan sensor position setting for measurethe vGRF between the human and ground.

B. Experimental Result

Fig.8(a) shows the %MVIC of comparing the EMGsmeasuring with or without wearing the exoskeleton. Whenthe exoskeleton was worn and the external load applied, most%MVIC values were lower than when the exoskeleton wasnot worn. In particular, %MVIC values increased linearlywhen the exoskeleton was not worn yet were constantwhen the exoskeleton was worn. In the case of the soleusmuscle, %MVIC decreased when the exoskeleton was worncompared to when it was not worn. This verified that theankle joint module, the quasi-passive joints developed in thisstudy, had normal function. The experimental data result ofthe %MVIC while not wearing and wearing the exoskeletonare summarized in Table IV.

Fig.8(b) shows the vGRF of comparing the TEKSCANmeasuring with or without wearing the exoskeleton. Mea-

10kg

20kg

HumanCR50HumanCR50

HumanCR50HumanCR50

0

20

40

60

80

100

120

SOL

GM

LevelWalking

VM

VL

%M

VIC

0

20

40

60

80

100

120

(a) The %MVIC during level walking

(b) The vGRF during level walking

Fig. 8: Comparison of the experimental result of the %MVICand vGRF while not wearing and wearing the exoskeleton

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TABLE IV: Experimental data result of the %MVIC whilenot wearing and wearing the exoskeleton

Experiment conditions Human without Exoskeleton %MVIC (%)

VL∗ VM† GM‡ SOL§

Level 10kg 8.97 17.00 16.19 50.53walking 20kg 9.74 27.08 17.39 55.08

Experiment conditions Human with Exoskeleton %MVIC (%)

VL∗ VM† GM‡ SOL§

Level 10kg 11.38 10.95 18.04 44.70walking 20kg 9.89 8.46 20.40 47.60

∗ Vastus lateralis muscle† Vastus medialis muscle‡ Gastrocmemius medialis muscle§ Soleus muscle

TABLE V: Decrement rate of vGRF impulse for stance phaseunder the experiment conditions

Experiment conditions vGRF impulse (×103) Decrement rate(%)

WithoutExo

WithExo

Level 10kg 3.93 3.59 9.46walking 20kg 4.25 3.95 7.67

sured vGFR values were represented by the gait cycleof the stance phase. The measured vGRF data can beanalyzed through the gait phase from the initial contactof 0% to the toe-off of 100%. The experimental resultswere expressed as mean (SD) with respect to the repeatedexperiments.Compared to the experimental result, the appli-cation of the worn the exoskeleton condition significantlyaffects vGRF parameters. The first vGRF peak value is muchhigher, the second vGRF peak value is lower. The increasein the early vGRF peak value therefore suggests that theexoskeleton did not absorb the impact of the heel strike. Thismight be because the ankle quasi-passive joint was actingon the plantar flexion only. vGRFs were lower when theexoskeleton was worn while at 30∼70% stance phase. Thisfact relates to external loads not acting on the human, butwhich transmitted to the ground through the exoskeleton.That results of the decrement rate indicates to the impulsefor 30∼70% of gait phase, as shown in Table V.

V. CONCLUSIONS

Through this study, a lower-extremity exoskeleton, calledHEXAR-CR50, was developed for industry and militarypurposes, particularly to achieve power augmentation dur-ing load carrying. To develop the exoskeleton, the designspecifications of the system were based on biomechanicalconsiderations and gait analysis. The design goal was asystem within a 23 kg weight, and the specification wasselected accordingly. Based on the design specification, anexoskeleton of a semi-anthropomorphic architecture wasdesigned, and joint shapes and layouts were studied. In

the design, each leg had 7-DOF and was configured witha combination of active and passive joints. An electricalmotor and harmonic gears were applied to the extension andflexion joints of hips and knees, and springs were applied tothe ankle and toe joints, which resulted in a quasi-passivemechanism. The experimental results showed that %MVICwas reduced against the external load applied, while vGRFhad a decrement rate, compared with the external load whenthe exoskeleton was worn, which verified the performance inaccordance with the development objective of load carrying.In the future, simplified control algorithms, a lightweightmechanical structure, and miniaturization will be used toimprove the efficiency of the current exoskeleton.

ACKNOWLEDGMENTThis work was supported by the Duel-Use Technol-

ogy Program of MOTIE/DAPA/CMTC.[13-DU-MC-16],and the National Research Foundation of Korea(NRF)grant funded by the Korea government(MEST).(No.NRF-2015R1A2A2A01002887)

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