2017 26th IEEE International Symposium on
Robot and Human Interactive Communication (RO-MAN)
Lisbon, Portugal, Aug 28 - Sept 1, 2017.

Analysis of Upper Extremity Motion during Trip-Induced Falls
Saeed Abdolshah, Member, IEEE, Yasuhiro Akiyama, Member, IEEE,
Kento Mitsuoka, Yoji Yamada, Member, IEEE and Shogo Okamoto
Abstract— Forward fall is one of the most common causes ofupper extremity fractures. Significant factors influencing
impact force and injuries were widely studied; however, it isnecessary to investigate the natural reactions of humans during
a forward fall to obtain a realistic evaluation of injuries. The
purpose of this study was to analyze the natural motion of theupper extremity during an induced trip. We carried out atripping experiment using an obstacle colliding with one leg;
while recovery step was prevented to produce a forward fall.
Results showed that the elbow extension had a slight ascending
trend during the forward fall and elbow angle at the moment ofhand-ground contact was appropriate to reduce the peak force.
Landing on the obstacle-side hand was more likely due to bodyrotation towards the obstacle-side. To prevent injuries, subjects
were connected to a safety harness not to strike the ground withhigh impact velocity. Thus, the fall motion was simulated using
a 12 DOF model to obtain a realistic evaluation of the impact
velocity and the related impact force caused by the forward fall
was estimated using a sagittal 3-segment model. Results of thisstudy can be useful in human-robot collaboration, where acollision between human and robot may cause a forward fall.
I.

INTRODUCTION

Forward falls are one of the most common types of falls
especially among older people [1], [2]. Some balancerecovery strategies have been reported such as elevating,lowering and skip motion [3]. Unsuccessful balance recovery
causes injuries such as wrist fracture, which is one the mostprevalent body fractures [4] due to high impact loading
during the fall. Conducting controlled human fall experimentin a laboratory can be hazardous and imposes high risks onparticipants. Thus, some researchers carried out experiments
to reproduce a part of fall motion (e.g. simulation of the realcondition with less intensity) and provide fall motion datasuch as impact velocity and force. However, such data may
strongly depend on experimental conditions such as fallheight, initial velocity and etc.
Tan et al [5] designed a tether-release falling experiment
to compare the impact velocity in forward and backward falls.
The subjects were released while leaning forward andsupported by a tether. They were instructed to land on theirknees followed by the hands. The average wrist velocity at
the time of landing was 1.33 m/s; however, the maximum
velocity was 3.57 m/s.
DeGoede et al. [6] designed a similar experiment. The
subjects were released while leaning forward and attached toa tether supporting about 30% of their body weight. Theirresults showed that the impact velocity was 2.6 ± 0. 3 m/sec;
and it could be reduced using a different falling arrest
strategy.
S. Abdolshah, Y. Akiyama, K. Mitsuoka, Y. Yamada and S.
Okamoto are with Department of Mechanical Systems Engineering,
Nagoya University, Nagoya, Japan (saeed.abdolshah@mech.nagoyau.ac.jp).
978-1-5386-3517-9/17/$31.00 ©2017 IEEE

There are a few studies that proposed a model to calculate
the impact velocity and the peak force. Some of themvalidated their models using the experimental results.
Chiu & Robinovitch [7] proposed a model of damped two
degrees of freedom spring-mass system to calculate theimpact force. They conducted some experiments to evaluate
the impact force during forward fall on outstretched arms andvalidated their model. One limitation of their study was thatthe elbow extension was ignored. Later they developed theirmodel with an extra spring as the ground stiffness. They set
the impact velocity of 3.83 m/sec for a free-fall of the bodyfrom a height of 0.75 m [8]. A similar study without bodystiffness characteristics was conducted in [9].
DeGoede et al. [10] proposed a three degrees of freedommodel with stiffness and damping elements to calculate theimpact velocity of the hand. They simulated the forward fallby an experiment of arresting a moving mass by one arm andverified their model. Later, they used computer simulation toinvestigate age-related effects on forward fall arresting using
a 2-D five-segment model [11]. Usually, the impact force
profile has two peaks and the fracture occurs by the first one,which is higher in hand-ground contact [12].
Zhou et al. [13] measured the tripping force (e.g.
obstacle–foot impact force), and applied the force to swing
foot of a 12 degree-of-freedom model. They did not consider
separate arms for their model. Lo et al. [14] used a 7 degreeof-freedom model to simulate forward fall motion bycontrolling body segments. They reported wrist impact
velocity in a range of 2.86 to 3.10 m/s. Xu et al. [15] also
used a 5 degree-of-freedom model (e.g. the model withoutarms) and simulated the fall motion using some motion
constraints. Wrist injuries during forward and backward falls
were also investigated in snowboarding using a computer
simulation [16].
The models and experiments clarified that the impact
force is strongly correlated with the impact velocity of thehand and the elbow extension. The previous designed
experiments were not coincident with conditions of areal trip-induced fall. To make an appropriate evaluation of
the forward fall injuries, it is important to investigate thehuman natural reaction during a forward fall to obtain elbow
extension and impact velocity during the impact phase.
In the current study, an experiment was designed to obtaina more realistic pattern of the trip-induced forward fall. We
conducted a tripping experiment using an obstacle colliding
with one leg; while preventing recovery step. Subjects’reactions were investigated during the falling motion withspecial attention to the upper extremity. The experimental
data provided a natural fall process pattern. The subjects wereconnected to a safety harness, which does not affect thehuman reaction, but it can influence the impact velocity.
Thus, to calculate the real forward fall impact velocity, a1485

multibody 12 DOF model was used to simulate the fallmotion. Finally, the impact force was calculated using a 2dimensional model with damping and stiffness elements.
II.

METHODS

A. Subjects
Six healthy young male volunteers participated in thisstudy. The age of the subjects was between 22 and 24 years
old and their average heights and weights were 172.4±3.2
cm and 60.0±8.9 kg respectively. None of the participants
had received fall arrest training such as Jodo or martial arts
techniques, and they did not have any records of illnesses
such as neurological disease, balance disorders or falls. The
Institutional Review Board of Nagoya University approved
all experiment protocols, and all participants provided awritten consent form.
B. Experimental Protocol
Motion capture system (MAC3D System, Motion
Analysis Corp.) with 10 cameras and 25 reflective markers
were used to measure the kinematics of body segments during
the fall motion. For safety reasons, a safety harness was
attached to the shoulders; a knee cover and ankle joint
supporter were provided to the participants. Two load cells(RSCC-200kg, Unipulse Corp.) measured the applied force
on the safety harness and the perturbation on the recovery leg.
An actuated linear slider was used to set the rope length
regarding the obstacle and tripping location. The obstacle wasprovided with three axis force sensors (USL-08-H6, Tech
Gihan Corp.) to determine the trip perturbation timing. The
data were recorded by a frequency of 100 Hz. The experiment
setup is shown in Fig. 1.
The participants were dressed in tight sportswear andmarkers were attached to their clothes. The subjects were
instructed to walk regarding the specific rhythm produced byan electric metronome and a visual guide. The visual guide
was moved by a three-phase induction motor (TO-K,
HITACHI Corp.). The walking speed of 1.5 m/s was set byboth guides. The subjects were instructed to walk
continuously to adjust their walking speed to the guides.
After the gait adjustment trials, the forward fallexperiments started. Tripping was randomly induced to theswing leg (right or left one) using a rigid flat bar. Elevating
and lowering are the most common recovery strategies [3].
For a successful elevating strategy, the subject should step
forward by placing the swing foot anterior to the body. Due tothe shape and height of the obstacle in this experiment, thecollision occurred between the obstacle and subject’s shank.
Consequently, the subject did not have the opportunity to step
forward by the tripped leg. To achieve a successful lowering
strategy, the recovery leg should move forward; however, toprevent the recovery motion, a rope was connected to thestance leg to constraint the recovery step. Such limitation canbe considered as reduced lower extremity range of motionthat occurs among older adults and is a kind of functional
disability [17]. Thus, subjects could not keep the balanceusing the lowering strategy and the forward fall occurred.

The position of the obstacle for each trip was decided
based on the gait motion of subjects. All participants wereasked to wear half-covered glasses to avoid detecting the
obstacle. A total number of 40 trial experiments wereconducted for each subject, where 30 trials were dummywithout any obstacles and in other 10 experiments, theobstacle caused the forward fall. The subjects were permitted
to place their arms in their desirable configuration. The
significant advantage of such marginal number ofexperiments without delay is the reduction of learning effects.
Thus, the subjects could not get accustomed to theexperiments.

Figure 1. Experimental setup to simulate a trip-induced forward fall. One
leg collides with the obstacle, while the other one is attached to a rope toprevent recovery motion. The experiment was equipped with a motion
capture system, visual and vocal guide, safety tether and load cells tomeasure the supporting force and the perturbation on recovery leg.

C. Data Analysis
The whole process of the forward fall in this experiment isdivided into three phases:
C. 1. Non-supported phase
The first phase starts when the tripping over the obstacle
induces. In this step, the contact between the swing leg andthe obstacle occurs at any time during the swing phase and aforward angular momentum of the body, induced by theforward trip causes imbalance. The common recovery
strategy is to obtain balance with producing enough backward
angular impulse.
The subject makes an effort to balance the body using theother leg, however, the recovery motion is prevented. This
phase terminates just before the safety harness supports the
body. In other words, during this phase, the force cell on thesafety harness does not sense any significant force, whichmeans the body is not supported by the harness. It can beinferred that in non-supported phase, a natural forward fall isin process.
C. 2. Supported phase
The second phase starts when the harness supports the
body for safety reasons and finishes just prior to the groundcontact. The reaction to the induced trip is not notably
affected by the safety harness [18]; however, the harness may
support and decelerate the motion. Thus, in this phase, thehuman reaction during the forward fall is considered to be1486

similar to a real forward fall. Our main focus would be theupper extremity, which is a very important part of theevaluation of impact force and injuries.
The dynamic characteristics of the motion may change
because of extra force from the safety harness. Especially thecontact velocity of the hand striking the ground is importantto calculate the impact force, and we need to regenerate themotion dynamic data.
A 12-DoF model is considered to simulate the second
phase motion (Fig. 2). The dynamic equation of the system isgenerated using Lagrange’s equation, which can be expressedas𝑀(𝑞)𝑞̈ + 𝐶(𝑞, 𝑞̇ ) + 𝐺(𝑞) = 𝜏,(1)where M is the inertia matrix; C represents the Coriolis andcentrifugal matrix, and G is the gravity matrix and 𝜏
represents the joint torques. 𝑞 denotes the generalized
coordinate vectors.
In this step, we regenerate the motion in sagittal plane
regarding a free-fall motion using the initial and final
conditions of experimental data. The orientation and velocityof the body segments at the end of the non-supported phaseare considered as the initial condition; however, the bodyorientation at the end of the supported phase is the final
condition. Some motion constraints were also imposed whichwere discussed in [14, 15]. The most important result of thisstep is the contact velocity, which can be used in the next stepto calculate the impact force.

considered as 3,500 kN/m3, and 6.0 (m/s)-1 respectively [10].
𝑥𝐻𝑎𝑛𝑑 , 𝑥̇ 𝐻𝑎𝑛𝑑 are the relative deflection and the velocity of
the hand into the ground. Initially at the moment of contact,
𝑥𝐻𝑎𝑛𝑑 is zero; however, 𝑥̇ 𝐻𝑎𝑛𝑑 is the one calculated at the endof the supported phase.
To calculate 𝑥𝐻𝑎𝑛𝑑 and 𝑥̇ 𝐻𝑎𝑛𝑑 , a sagittal 3-segment model[10] is considered as shown in Fig. 3. 𝜃1 (arm angle) is theangle between the horizontal axis and the arm, and 𝜃2(forearm angle) is the angle between the forearm and thehorizontal axis. Linear shoulder stiffness and damping indices
are 3.5 kN/m and 0,12 kNs/m respectively and the rotational
elbow stiffness is determined as 2.28 Nm/° [11]. The effective
mass can be calculated as [7]𝑀=

𝐹𝑠𝑡𝑎𝑡𝑖𝑐,𝑔−𝑚(3)where 𝐹𝑠𝑡𝑎𝑡𝑖𝑐 is the force measured by the force plate in staticstate, when all oscillations damped. The gravitational constantis denoted by 𝑔 and 𝑚 is the total mass of hand, arm andforearm. This model was simulated in MSC.visualNastran4D, 2002.

Figure 3. A sagittal 3-segment model of the impact phaseIII. RESULTS

Figure 2. Human model with 12 degrees of freedom to simulate thesupported phase in sagittal plane

C. 3. Impact phase
In this phase, hand-ground contact occurs. The collision
between heel of the hand with the ground generates
a significant instantaneous force. The impact force consists of
two local extremes. In case of the hand striking the ground,usually, the first peak is more important and is the main
reason of injuries. To calculate the impact force, we used a
non-linear model that was proposed in [10]3|(1 − 𝐵𝐻−𝐺 𝑥̇ 𝐻𝑎𝑛𝑑 ),𝐹𝑖𝑚𝑝𝑎𝑐𝑡 = 𝐾𝐻−𝐺 |𝑥𝐻𝑎𝑛𝑑(2)

where 𝐾𝐻−𝐺 and 𝐵𝐻−𝐺 are the stiffness and damping
parameters of the hand-ground interface, which are

A. Supported phase – Body disposition
Fig. 4 shows a typical falling process during the supported
phase, where the obstacle struck the left leg and the recovery
step of the right leg was prevented. The subject landed firston his hands followed by the knees. Mostly, the participants
fell symmetrically without any steps forward or backward;
however, in some cases, asymmetrical fall motion wasobserved. The trips were induced on an average of 32.0 ±
3.7% of gait cycle. The obstacle-side body is marked withcircles on the body segments.
The absolute arm angle during the supported phase isillustrated in Fig. 5.a. The initial angle of the obstacle-side
arm is higher than the other one. The angles of both arms
prior to ground contact are similar and around 75°. At the
beginning of the phase, both arms have ascending trends,
followed by a descending trend. Arms have the lowest angle
at the end of this phase.

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Fig. 5.b. shows the absolute forearm angle. Despite thearm, the forearm angle is always ascending. Similar to thearm angle, at the starting point of the supported phase, therecovery-side arm angle has more variation. At the end of the
supported phase, the angle of the obstacle-side forearm(82.8°) is higher than the recovery-side forearm (75.7°).

mass height in the sagittal plane. As expected, the center ofmass has a descending trend; and the slope of the trend is lessthan the hand graph.

Figure 5. Average, minimum and maximum of (a) arm angle, (b) forearm
angle, and (c) elbow extension for obstacle-side arm, and recovery-side arm

B. Supported phase – motion data
The average initial and final condition of body segments
were extracted from the experiments. The dynamic model isdeveloped to simulate forward fall process. The mostimportant result of this section is the hand velocity at the endof the supported phase.

Figure 4. (a) A sample forward fall motion in supported phase, with an
obstacle facing the left leg, (b) sagittal plane view of the fall

The elbow extension is shown in Fig. 5.c. The average
obstacle-side elbow extension is higher than the other side.
The final elbow extension for the obstacle-side is 157.8°;
however, it is 150.7 ° for the recovery-side. The rate ofchanging elbow extension prior to ground contact is low,which can be beneficial to reduce the peak force [6]. The
subjects had not received any training for such elbow
extension.
The hand height in the sagittal plane is illustrated in Fig.
6.a. The initial average height is around 0.62 m. The obstacleside hand has a lower height. Based on this graph, it can beinferred that in average, the obstacle-side hand strikes theground prior to the other one. Fig. 6.b. shows the center ofFig. 7.a. shows the regenerated height of the hand during
the trip-induced forward fall using the Lagrange’s equation.
The initial average height is considered about 0.62 m and therelated velocity graph is illustrated in Fig. 7.b. The impact
velocity is about 3.89 m/s.
C. Impact phase
To calculate the contact force during the forward fall, themodel in Fig. 3 was used. The arm and forearm angles werealso obtained from the experimental data. The contact force
during the impact phase is shown in Figure 8. The firstmaximum impact force is approximately 2.3 kN. The firstpeak is related to the high-frequency transient occurring
slightly after heel of the hand strike. The second peak occurs

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after a local minimum and is usually smaller in case of handcollision to the ground.

These experiments were enough to obtain a reasonable rangeof results. The elbow extension was 157.8° for the obstacleside and 150.7° for the recovery-side. DeGoede et al. [6]reported an average of 168° from the experiments of fivesubjects. In their study, the subjects were aware of the fallupon the release of the supporting tether. Thus, their priorintention may have influenced the elbow angle. Moreover the
subjects had less time to react the forward fall.
Results show that during the falling process, the obstacleside hand has lower height compared with the recovery-side
hand. Therefore, it is more probable that the obstacle-side
hand strikes the ground prior to the other hand. The reason isrelated to the usual attitude of the recovery motion. The
subjects rotate the body towards the obstacle-side [19] and therotation axis is the inclined stance leg. Consequently, theobstacle-side hand has lower height during the falling
process, however, our experiments show that the subjects
decrease the height difference between hands prior to groundcontact. The other reason can be the different elbow angles.
In average, the obstacle-side elbow has higher extension
compared with the recovery-side elbow as shown in Fig. 5.c.

Figure 6. (a) The hand and (b) the COM (center of mass) height from theground level

Figure 8. The ground reaction force during the hand-ground contact

The higher elbow extension in obstacle-side may occur asa natural reaction to control the balance. The difference
between the elbow extensions may cause diverse impact
forces on left and right sides. To the best of the author’s
knowledge, no previous work discussed the influence of suchdifferent elbow angles. All proposed analytical models to
calculate the impact force were 2-dimensional, and suchdifference in elbow angles has not been considered yet.

Figure 7. (a) The position and (b) velocity of the hand from the groundlevel in vertical direction of the sagittal plane

IV. DISCUSSION

This study was designed to investigate the natural reaction
against the forward fall. We have limited the scope of ourcurrent investigation to upper extremity motion during a tripinduced fall. Due to difficulty of conducting such experiments
and the injuries may occur for volunteers, we limited theexperiments to six young subjects and did not use elderly.

The average elbow extension and the rate of changing theelbow angle, were both appropriate to reduce the peak force
applied to the distal forearm. The subjects chose thisconfiguration of the arm and forearm arbitrarily and without
any training. The elbow angle in this study was lower than thenatural elbow angle in [6]. One reason can be the type of theirexperiments, which was a controlled laboratory experimentwith a lower falling height. The other reason may be thedifferences between the types of falls. In their experiments,
the subjects did not face any obstacles and the fall occurredby releasing a supporting tether (i.e. No external force applied
to the swing leg). To confirm our results, it is necessary to reconduct the experiment using people with different ages andraces. The impact velocity was calculated as 3.89 m/s. In
another study [8], it was calculated using the assumption of
free-fall motion from 0.75 m height. They evaluated theimpact velocity as 3.83 m/s. In this study, the first impact
force was calculated as 2.3 kN. This value can be considered1489

as the natural first impact force of forward falling induced bytripping.
V.

CONCLUSION

The main purpose of this study was to investigate naturalforward fall pattern. Elbow angle and impact velocity play
important roles in intensity of impact force and should betaken into consideration to evaluate fall injuries. Most of thefall studies conducted some experiments to reproduce a partof real fall motion; however, we carried out lifelike fallexperiments to obtain the human reaction against forward fall
without any special instructions. We have constrained thescope of our study to a forward fall arrested with both hands;
which is a quite common strategy. Results may not beapplicable to the other fall arrest strategies or different types
of falls such as sideways fall.
The outcome shows that the arm angle reduces during thefall process; however, the forearm angle increases and theelbow extension has a slight ascending trend. The average
elbow extension at the moment of hand-ground contact was
about 154.25°; however, the obstacle-side elbow extension
was higher (157.8 ° ). The responses to the trip were noteminently influenced by the safety tether. Thus the elbow
angle can be considered as the natural elbow angle against
forward fall without any practice or instruction. This angle
has a significant role in the intensity of the impact force andstrangely, in these experiments, the average elbow angle wasin a range that reduces the impact force.
Furthermore, results show that the obstacle-side hand has
lower height during the falling process. Therefore, theobstacle-side hand may strike the ground prior to the otherside. This may occur due to twisting the body towards the
obstacle-side with the inclined stance leg as the rotation axis.
One more reason can be the higher elbow extension angle in
the obstacle-side. The subjects decreased the height
difference between hands prior to ground contact.
Due to the application of the safety harness, the forward
fall dynamic data could not be used and we regenerated the
motion using Lagrange’s equation. We used the experimental
data of the initial and final conditions of body segments andthe motion constraints to calculate the impact velocity in afree-fall motion. The impact velocity was about 3.89 m/s,
which is not very different from the previous studies [7]. The
related impact force was also estimated regarding the average
elbow extension and the impact velocity, which was about 2.3
kN. The results of this study can be useful in human-robot
collaboration, where a collision between human and robot
may cause a forward fall.
ACKNOWLEDGMENT
This work was supported by JSPS KAKENHI Grant
Number 26750121, and METI “Strategic international
standardization
promotion
project:
International
standardization of human tolerance against fall injuries”.
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