Development of a Full Body FE for Reconstructions - IRCOBIolid materi ents; ACL a e knee join pressive loa dled by the not bear c hose repre ar elements riginal mod L/aPCL, post sented

Abstract The goal of this study is to develop a full‐body human Finite Element (FE) model with the

enhanced biofidelity and the numerical robustness necessary for use in dynamic car–pedestrian accident

reconstructions in various impact scenarios. An FE model for the pelvis and lower limbs in a standing position

developed in our past study was combined with FE models for bones and ligaments in the upper body taken

from an occupant model. The biofidelity of the combined model was further enhanced and validated against

experiments. Each body region of the full‐body model was validated in multiple loading conditions, including

neck bending, thoracic impact, isolated pelvis impact, thigh and leg bending, knee bending, knee ligament

tension, and ankle joint rotation. The full‐body model was subjected to lateral and frontal impacts at 70 km/h

against a simplified car model representing a stiff front‐end structure, showing that the model is robust enough

to simulate high speed impacts in different pedestrian orientations.

Keywords biofidelity, human model, pedestrian, numerical robustness

I. INTRODUCTION

Crash simulations using a human Finite Element (FE) model representing a pedestrian in multiple car–

pedestrian accident scenarios are one of the most effective means to investigate car–pedestrian interaction,

quantify dynamic loadings to the pedestrian and predict probability of fatality or injury. This process is essential

to evaluate real‐world relevance of pedestrian safety technologies, including active and passive safety

measures. The currently used subsystem test procedure to evaluate pedestrian passive safety performance was

originally developed by the European Enhanced Vehicle‐safety Committee (EEVC) [1] and essentially assumes

lateral impact of a car to a pedestrian at 40 km/h. Due to the representation of pedestrian accidents with

particular impact scenarios, it is not possible to accurately estimate the real‐world effectiveness of passive

safety measures evaluated by such a test procedure based solely on the test results. It needs to be investigated

against various impact configurations seen in actual car–pedestrian accidents in order to predict reduction in

pedestrian fatalities and casualties expected to be provided by the measure in real‐world accidents. The recent

advancement of active safety technologies would change the distribution of impact speeds as a result of the

activation of systems for crash avoidance. The prediction of the combined effect of both active and passive

safety measures on the reduction of pedestrian fatalities and casualties also requires prediction of car–

pedestrian interactions at various impact configurations. Fredrikson et al. [2] investigated 54 pedestrian

accidents from the German In‐depth Accident Study (GIDAS) database, occurring from 1999 to 2008 and with

AIS3+ head injuries, and estimated isolated and combined effects of specific active and passive safety measures

on head injury mitigation. The study used a particular function between the impact speed and the probability of

AIS3+ head injury to estimate the effect of the measures, which assumes that the injury probability is solely

dependent on the impact speed. This assumption may not apply if the impact speed distribution is significantly

different between urban and rural areas where the distribution of car–pedestrian impact configurations would

be significantly different. These issues can be taken into consideration when all the different impact

configurations in the real‐world accidents are dynamically reproduced by car–pedestrian impact simulations

using a human model.

Accurate predictions of the probability of injuries in multiple accident scenarios by means of impact

simulations using a human FE model would require the biofidelity of the human model. The Global Human Body

Models Consortium has developed a detailed and extensively validated human FE model for a seated occupant

Y. Takahashi is Chief Engineer (tel: +81‐28‐677‐3311, fax: +81‐28‐677‐7500, e‐mail: [email protected]), H. Asanuma is Engineer and T. Yanaoka is Assistant Chief Engineer at Honda R&D Co., Ltd. Automobile R&D Center, Japan.

Development of a Full‐Body Human FE Model for Pedestrian Crash Reconstructions

Yukou Takahashi, Hiroyuki Asanuma, Toshiyuki Yanaoka

IRC-15-63 IRCOBI Conference 2015

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by representing whole‐body skeletal and soft tissues. However, a detailed pedestrian model in a standing

position is still needed to be developed in a future stage [3]. Watanabe et al. [4] developed a full‐body human

FE model representing a pedestrian with a primary focus on brain and internal organ injuries. However, the

result of 4‐point knee bending validation is the only information provided for the lower limb model validation.

In a previous study, the authors developed a human FE model with the upper body represented by articulated

rigid bodies and the pelvis and lower limbs modeled by finite elements to investigate pelvis and lower limb

interactions with a car front‐end structure in car–pedestrian impacts [5‐6]. Since the upper body is modeled by

articulated rigid bodies representing only the external surface, it was not possible to predict injuries to the

upper body such as rib fractures or abdominal injuries. The ankle joint is represented by one single mechanical

joint element, which would require biofidelity improvement when ankle joint injuries are to be predicted. The

knee model was validated in pure (4‐point) lateral bending. Although the primary loading to the knee in

pedestrian crashes would be lateral bending, the knee is also subjected to shear loading and thus additional

validation in combined loading would be needed. The bending response of the lower limb model was validated

for individual long bones (femur, tibia and fibula) and assembly (thigh and leg). This ensured biofidelic overall

bending response of the leg assembly, however the distribution of the loads transmitted by the tibia and the

fibula has not been clarified. In addition to the biofidelity improvement of the model, it is also necessary to

make sure that the model is robust enough to survive severe impacts from a car that need to be simulated to

estimate real‐world relevance of safety technologies by means of computer simulations of car–pedestrian

crashes.

The goal of the study is to develop a human FE model in a standing position that will ultimately be capable of

predicting probability of major injuries to a pedestrian whole body in various impact configurations. This study

developed a full‐body human FE model in a standing position by combining the FE upper body model taken

from Dokko et al. [7] with the pelvis and lower limb FE model developed by the authors [5‐6]. Some

modifications were made to the knee, leg and ankle for improved biofidelity. The neck, thorax, pelvis, thigh,

knee, leg and ankle were validated against experimental data to validate the components of the modified

full‐body model. The model validation included 3‐point bending of the knee and the bending response of the

ankle, which have not been conducted in former studies [5‐6]. The model was subjected to lateral and frontal

impacts from a simplified car model representing a stiff front‐end structure at 70 km/h, in order to evaluate the

numerical robustness of the model.

II. METHODS

The combined FE whole body model was further modified from two different viewpoints: one to enhance

the biofidelity of the model; and the other to improve the numerical robustness of the model in high‐speed

collisions. The components of the modified model were re‐validated against experimental data due to the

modifications made. The 3‐point bending of the knee and the ankle rotation were added to the validation for

enhanced biofidelity. The numerical robustness of the model was evaluated in lateral and frontal collisions. The

lateral impact was used to represent the impact configuration most frequently seen in car–pedestrian accidents.

The frontal impact was employed to investigate the effect of the musculature around the knee on the

resistance to hyperextension of the knee.

Full‐body Pedestrian Model

The upper part of the whole body human FE model in a seated position developed by Dokko et al. [7],

representing the shape of the rib cage for 35 years old, was combined with the pelvis and lower limb human FE

model in a standing position developed by the authors [5‐6] at the T12‐L1 joint. The spine curvature was not

changed. Both models have been developed in PAM‐CRASH. As opposed to the size of the pelvis model scaled

to 50th percentile American male [5‐6], the rib cage model was based on CT scans from a particular individual

[8]. Although the bell‐shaped rib cage of the baseline model can be viewed as part of individual variations, the

width of the bottom of the rib cage was much larger than the width of the pelvis scaled to 50th percentile male.

Since the protrusion of the bottom of the rib cage may predict unrealistic rib fracture mechanism, it was

decided to morph the shape of the rib cage to match the contours of the rib cage in frontal and lateral views


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TABLE I

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of the skin an

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T12 T12‐L1 L1‐L2

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si

te

ompared for

exion, the li

cceleration a

cceleration

nterior‐poste

g. 7. Lateral

Thorax: th

ecause of th

ntervertebral

or the fronta

onditions, w

he impact sp

o the lack of

istories of th

quations:

where is le

eight of the

caled thorac

me (s),

ensity and

isplacement

orce and the

were determi

was compare

mpact, the im

mpact speed

mulations. F

ests.

r both flexion

near accele

about the la

in the late

erior axis we

bending and

he impact re

e change in

l discs. The t

al impact an

hich are sum

peed, the ave

the scaled d

he force and

ength scale f

subject test

ic displacem

is time f

the elastic

was plotted

thoracic def

ined from th

d with the f

mpact force

s of 4.4 m/s,

Figure 8 illus

n and lateral

ration in the

ateral axis w

ral and the

ere compared

d flexion of h

esponse of

the contour

horacic impa

nd the obliq

mmarized in

erage values

ata, each of

the thoraci

λ

factor, ted, i

ment (m),

from the tes

modulus eq

d, and ellipse

flection were

he envelop c

orce‐deflect

time history

, 6.5 m/s an

strates the c

bending aga

e anterior‐p

were compar

e superior‐in

d.

head‐neck co

the thorax

of the rib ca

act test resu

que impact,

Table II, wer

s of the test

the test resu

c deflection

∙

∙

∙

is stan

s scaled forc

is tho

t (s). This sc

qual 1. For

es with the le

e drawn. The

curves. The

ion corridor

y corridors s

d 9.5 m/s w

computer sim

ainst the cor

osterior and

red with the

nferior direc

omplex.

was also re

age and the

lts conducte

respectively

re used. Due

conditions w

ults presente

were scaled

ndard height

ce (N),

oracic deflect

caling proced

each scale

engths of th

en the upper

force‐deflec

for each of

caled to 50t

ere used to

mulation mo

rridors prese

d the superi

e test result

ctions and

e‐validated i

modification

ed by Kroell e

y. For the fr

e to the varia

were applied

ed in the pap

d based on E

t (175.6 cm

is force f

tion from th

dure assume

ed time, the

e axes repre

r and lower b

ction respon

the impact

th percentile

compare the

odels represe

ented by Thu

or‐inferior d

ts. For the la

the angular

n frontal an

ns to the ben

et al. [22‐23]

rontal impac

ations in the

d to the hum

pers [22‐23]

Eppinger et a

for 50th per

from the sub

he subject te

es that the s

e average f

esenting one

bounds of th

se from the

conditions i

e male prese

e impact for

enting the fr

unnissen et a

directions an

ateral bend

r acceleratio

nd oblique i

nding charac

] and Viano [

ct, three dif

mass of the

man model si

was digitized

al. [24] using

rcentile mal

bject tested

ested (m),

scaling facto

force versus

e standard de

e force‐defle

e human mo

n Table II. F

ented in Vian

rce from the

rontal and o

al. [21]. For t

nd the angu

ing, the line

on about t

mpactor tes

cteristics at t

[24] were us

fferent impa

e impactor a

imulation. D

d and the tim

g the followi

(

(

(

(

e ),

(N),

is scal

rs of the ma

s the avera

eviation of t

ection corrid

del simulatio

or the obliq

no [24] for t

human mod

oblique impa

he

lar

ear

he

sts

he

ed

act

nd

ue

me

ng

(1)

(2)

(3)

(4)

is

is

ed

ass

ge

he

dor

on

ue

he

del

act


- 535 -

Th

in

by

be

va

th

th

ra

de

co

al

th

bo

an

de

ap

to

w

am

th

de

th

de

de

st

Th

hi

th

an

Leg: the le

he leg mode

n Fig. 9, and

y Ivarsson e

etween thes

alues from th

he inertial sp

he lack of inf

atio was dete

efined and t

ontribution o

. [27]: 6.0%;

Knee Ligam

he structure

one‐ligamen

nd 1600 mm

evelop force

pplied due to

o a large var

was used. Fir

mong the su

he force from

eflection wa

he ratio up

eflection lev

etermined d

tandard devi

he standard

istory used b

he dynamic t

nd non‐linea

Cond

Test ID M (

42FM

45FM

53FM

60FM

Ave.

Val.M: Mass ofAve.: Avera

FTONTAL

eg model wa

l was subjec

the momen

et al. [26]. T

se two bone

he literature

pace and the

formation fr

ermined, thr

he ratio was

of the fibula

Takebe et a

ment: the mo

of the mod

nt‐bone com

m/s (dynami

e‐deflection

o the lack of

riability of th

rst, the aver

ubjects, and

m the test w

s calculated

to the larg

vel. The aver

directly from

iation in forc

deviation i

by the tests [

tension of aA

ar rate‐depen

dition 1

kg) V (m/s) Test

22.9 4.9 48FM

23.0 5.1 50FM

23.0 5.2 51FM

23.0 4.3 52FM

56FM

58FM

62FM

22.98 4.88 Ave.

23.0 4.9 Val.f impactor, V: Impage, Val.: Validatio

TABLE II

L IMPACT COND

s re‐validate

cted to dynam

t‐deflection

The effect o

s was verifie

e in the axial

e foot define

om the liter

ree different

s calculated

in axial com

al. [28]: 6.4%

odels for the

del for thes

plex extracte

c). The data

corridors an

f the anthrop

he data betw

rage force‐d

the standar

with the sma

and curve‐fi

est ultimate

age values a

the test res

ce was appli

n the deflec

[30‐31]. Figu

ACL. Materia

ndent constit

Condition 2

ID M (kg) V (m/

M 10.4

M 10.4

M 10.4

M 10.4

M 10.4

M 10.4

M 10.0

10.34 6

10.3 act speedon Condition

DITIONS

ed in 3‐point

mic mid‐shaf

response w

of the additi

ed by compa

compressio

ed as a rigid

ature on the

t cross‐sectio

in all of the

pression obt

; Shuler et a

e ACL, PCL an

se ligaments

ed from the

a from Bose

nd the avera

pometric info

ween subjec

eflection cu

d deviation

allest ultima

tted as a fun

e deflection

and the stand

sults to dete

ied to deter

ction was no

re 10 presen

l Type 202 w

tutive charac

Conditio

/s) Test ID M (kg)

7.1 12FF 22

7.3 13FM 22

6.7 14FF 22

7.2 15FM 23

6.9 18FM 23

6.8 20FM 23

6.9 21FF 23

22FM 23

63FM 23

64FM 23

6.99 Ave. 23.

7.0 Val. 23

t bending du

ft 3‐point be

as compared

ion of the i

aring the co

on of the leg

body was fo

e exact cross

ons (mid‐sha

three cross‐

tained from

l. [29]: 6–17%

nd LCL were

s. The valida

knee mode

et al. [30]

age and vari

ormation on

cts, a specific

rve was det

in force was

ate deflectio

nction of def

to estimate

dard deflect

rmine the av

rmine the up

ot taken int

nts the exam

was used for

cteristics we

on 3

V (m/s)

2.9 7.2

2.9 7.4

2.9 7.3

3.6 6.9

3.6 6.7

3.6 6.7

3.6 6.8

3.6 6.7

3.0 6.9

3.0 6.9

27 6.95

3.3 7.0

Fig. 8. Fro

ue to the ad

ending at 1.5

d with that f

nterosseous

ntribution o

. The proxim

orced to disp

s‐section of t

aft, distal thi

‐sections at

the literatur

%).

individually

ation was co

l at the load

and van Do

ability of th

some of the

c procedure

termined up

s calculated

n to that fro

flection. The

e the avera

ion of the fo

verage failur

pper and low

o account d

mple of the b

the bar elem

ere specified.

ontal and ob

dition of the

5 m/s of the

from the res

s membrane

f the load b

mal end of th

lace vertical

the leg at wh

rd and proxi

1.0 kN comp

re ranged fro

re‐validated

onducted in

ding rates of

mmelen et

e failure po

e subjects us

to develop

p to the sma

for each de

om the test

n the functio

ge force‐def

orce and the

re point and

wer bounds o

due to the p

one‐ligamen

ments repres

.

lique thorac

e interosseo

loading rate

sponse corrid

e on the loa

orn by the f

he tibia was

lly at 250 mm

hich the load

imal third) o

pression. The

om 6.0% to 1

d because of

dynamic te

f 0.016 mm/

al. [31] wer

ints. Data sc

sed in the ex

force‐deflec

allest deflec

eflection. The

with the la

on was used

flection curv

e deflection a

d its variation

of the respo

prescribed d

nt‐bone com

senting the l

ic impact.

us membran

e, as present

dor develop

ad distributio

fibula with t

rigidly fixed

m/min. Due

d transmissi

of the leg we

e values of t

17.0% (Funk

f the change

ension using

s (quasi‐stat

re analyzed

caling was n

periment. D

ction corrido

ction at failu

en the ratio

rgest ultima

to extrapola

ve up to th

at failure we

n. The avera

onse corrido

eflection tim

plex model f

igament fibe

ne.

ed

ed

on

he

to

to

on

ere

he

et

in

g a

tic)

to

not

ue

ors

ure

of

ate

ate

hat

ere

ge

rs.

me

for

ers


- 536 -

th

th

lo

re

pr

et

th

(F

fo

O

w

Co

sl

in

en

co

et

F

N

Th

of

tu

se

pe

co

co

an

bu

31

Th

Knee Joint

he ligament

he knee, the

oads, which

esponse in va

resented in

t al. [32]. Fig

heory, the ra

Fig. 11). In or

orce, the tes

ne of the te

was conducte

onsidering t

ack strain of

Ankle: due

n dorsiflexion

nd of the tib

orresponding

t al. [33].

Fig. 11. 3‐po

Numerical Ro

he numerica

f the measur

urned off to

evere condit

edestrian pr

ombination

omponents.

ny pedestria

umper, bum

1 old Japane

he stiffness

Fig. 9. 3‐poi

t: the knee jo

and the kne

model was

normally tak

algus bendin

Ivarsson et a

gure 11 pre

atio of the va

rder to valida

sts with a sm

sts (test ID:

ed under the

he use of th

f 0.2 was app

e to the mod

n/plantarflex

bia was rigid

g joints at 50

int knee valg

obustness Eva

l robustness

res to enhan

provide lar

tions. A simp

rotection reg

of multiple

Unpublished

an protectio

mper, hood e

ese cars in th

characterist

nt bending o

oint model w

ee joint caps

also validate

kes place in

ng of the kne

al. [26]. For

sents the 3‐

algus bendin

ate the knee

maller mome

Comb 7 in B

e moment‐to

he specimen

plied to the l

ification of t

xion and inve

dly fixed to t

00 degrees/s

gus bending

aluation

s of the mod

ce the nume

rgest‐possibl

plified car m

gulations an

rigid surface

d legform, u

n measures

dge and win

e sedan cate

tics of these

of leg

was re‐valida

ule models.

ed in 3‐point

actual car–p

e was comp

3‐point bend

‐point bendi

g moment a

e model unde

ent‐to‐shear

Bose et al. [3

o‐shear force

s in the test

igament and

the ankle in t

ersion/eversi

the inertial s

s. The mome

set‐up used

el was evalu

erical robust

e internal lo

model represe

d NCAPs wa

es connected

pper legform

were analyz

ndshield. The

egory prior to

component

ated in 4‐po

Since 4‐poi

t bending tha

pedestrian c

ared with th

ding, the she

ing test set‐

applied at th

er the condit

force ratio

32]) was omi

e ratio of 2.6

ts without p

d knee joint c

the modified

ion were com

space and th

ent‐angle re

by Bose et a

uated in an im

ness. Failure

oads and ev

enting a stif

as used for

d to linear s

m and headfo

zed to obta

e shape of th

o the compli

ts were dete

Fig. 10. In

oint bending

nt bending e

at applied a

collisions. Fo

e response c

ear force tim

‐up used by

e knee to th

tion with rela

(test ID: Com

tted as an o

69, determin

reconditioni

capsule mod

d model, the

mpared with

he rigid‐body

sponses wer

al. [32].

mpact from a

e representat

aluate the n

ff car front‐e

the impact

spring eleme

orm test res

in the avera

he car mode

iance with th

ermined fro

ndividual kne

due to the c

essentially p

combination

or 4‐point be

corridor scale

me history w

Bose et al.

e shear forc

atively large

mb 4‐8 in Bo

utlier. Three

ed from the

ng the ligam

els.

moment‐an

that from th

y foot was f

re compared

a car at 70 k

tion options

numerical ro

end structure

simulations.

ents represe

ults against

age stiffness

l was determ

he regulation

m the avera

ee ligament t

change in th

provides a pu

n of the bend

ending, the m

ed to 50th p

was compared

[32]. Based

ce is determi

r contributio

ose et al. [32

e‐point bend

e average of

mentous tiss

ngle response

he literature

forced to rot

d with those

km/h to evalu

of the huma

obustness un

e prior to th

. The mode

enting the st

old Japanes

s of the low

mined from t

n for pedestr

age + 1SD. T

tension

he structure

ure bending

ding and she

moment‐ang

percentile ma

d against Bo

on the bea

ned by

on of the she

2]) were use

ding simulatio

the four tes

ues, the init

es of the ank

e. The proxim

tate about t

from Crand

uate the effe

an model we

nder the mo

he mandate

l consists of

tiffness of t

e cars witho

er part of t

the average

rian protectio

The car mod

of

of

ear

gle

ale

ose

am

ear

ed.

on

ts.

tial

kle

mal

he

all

ect

ere

ost

of

f a

he

out

he

of

on.

del


- 537 -

re

ch

lo

th

be

an

hu

hi

ef

sh

ch

F

F

M

la

an

th

an

af

re

th

de

co

im

be

hu

m

ra

epresents th

haracteristics

ongitudinal d

he car surfac

etween the c

nd mid‐gait

uman mode

it by a car b

ffect of the a

hows the exa

haracteristics

Fig. 12. Car‐p

Fig. 13. Force

Model Valida

Neck: Fig.

teral bendin

nd angular a

he flexion of

ngular accele

ffixed to the

Thorax: Fig

esponse corr

hree differen

eflection wa

ompares the

mpact speeds

Leg: Fig. A

ending of th

uman mode

model simulat

anged from 7

0

20000

40000

60000

0.00 0

Force (N)

D

he stiffness

s, as presen

direction don

ce area was

car center pa

cycle. Three

l. The uprigh

oth laterally

addition of t

ample of the

s of the car m

pedestrian im

e‐deflection

ation

A1 in the A

ng of the nec

acceleration

the neck for

eration abou

inertial spac

g. A3 in the

ridor determ

nt combinat

as determine

e impact for

s of 4.4 m/s,

A5 in the App

e mid‐shaft

l. Figure 14 s

tion for the f

7.3% to 16.1%

0.05 0.10 0.15

Deflection (m)

s of the c

ted by Asan

ne by the ear

s also divide

art and the r

impact sim

ht model was

y and frontal

the musculat

e car–pedest

model for dif

mpact simula

characterist

Appendix com

ck for the he

about the fr

r the head C

ut the latera

ce.

e Appendix

ined in the c

tions of the

ed as the d

rce time hist

6.5 m/s and

pendix show

leg between

shows the fo

four differen

% at 1 kN of

Lower Bumper

Bumper

Hood Edge

components

uma et al. [

rlier study to

ed in the lat

right and left

ulations wer

s hit by the c

ly at 70 km/

ture around

trian impact

fferent regio

ation.

ics of simplif

mpares the

ad CG trajec

rontal (X) ax

G trajectory,

l (Y) axis. All

compares t

current study

impactor m

isplacement

tory with th

d 9.5 m/s.

ws the compa

n the respon

orce time his

nt cross‐sect

the total for

0

2000

4000

6000

0.00 0

Force (N)

D

r

using one

34]. In addit

o represent t

teral directio

t parts. Two

re conducted

car model la

/h. The front

the knee on

simulation f

ons along the

fied car mod

III. RESULTS

human mod

ctory, lateral

xis. Figure A2

, frontal (X) a

the accelera

the force‐de

y from the te

mass and th

t of the imp

he scaled for

arison of the

nse corridor

story in axial

ions. The co

rce, which ar

0.05 0.10 0.1

Deflection (m)

e‐dimensiona

tion to the d

he distributi

on to repres

pedestrian lo

d using the s

terally at 70

tal impact si

n the knee k

for lateral im

e centerline o

del.

S

del simulatio

(Y) and vert

2 in the App

and vertical

ations are re

eflection res

est results p

e impact sp

pactor relati

rce time hist

e moment‐d

presented in

compressio

ntribution fr

re within the

15

Hood Front

Hood Center

Hood Rear

al springs w

division of th

on of differe

sent differen

ower limb st

simplified ca

km/h. The m

mulation wa

kinematics du

mpact. Figure

of the car.

on results w

tical (Z) linea

endix prese

(Z) linear acc

elative to the

ponse of th

resented in K

peed shown

ve to T9. Fi

tory corrido

eflection res

n Ivarsson et

on of the leg

om the fibul

e range from

r

0

4000

8000

12000

16000

0.00 0

Force (N)

D

with non‐li

he car front

ent stiffness

nt stiffness

tances were

ar model and

mid‐gait stan

as conducted

uring a collis

e 13 present

with the test

ar acceleratio

nts similar c

celeration of

e global coor

he thorax w

Kroell et al.

in Table II

igure A4 in

or from Vian

sponse in dy

t al. [26] and

from the m

la to the ove

the literatur

0.05 0.10 0.1

Deflection (m)

near stiffne

surface in t

characterist

characterist

used – uprig

d the modifi

nce model w

d to clarify t

sion. Figure

ts the stiffne

results in t

on of the he

comparisons

f the head a

rdinate syste

ith the scal

[22‐23] for t

. The thorac

the Append

o [24] for t

ynamic 3‐poi

d the modifi

odified hum

erall axial for

re.

15

Glass Front

Glass Center

Glass Rear

ess

he

ics,

ics

ght

ed

was

he

12

ess

he

ad

in

nd

em

ed

he

cic

dix

he

int

ed

an

rce


- 538 -

F

kn

po

th

sc

th

50

F

4

in

F

N

A

ki

co

Fig. 14. Load

Knee Ligam

nee ligament

oint variatio

he failure po

Knee Joint

caled respon

he shear forc

0th percenti

Fig. 15. Valid

4‐point valgu

Ankle: Fig

nversion/eve

Fig. 17. Valid

Numerical Ro

ll of the thr

nematics of

onfigurations

distribution

ment: Fig. A6

ts at the loa

ns determin

int variation

: Fig. 15 com

nse corridor f

ce time histo

le male and

dation results

us bending.

. 17 shows

rsion betwe

dation results

obustness Eva

ree impact s

f the modifi

s.

n between tib

6 in the Appe

ding rates o

ned in this st

were not de

mpares the m

from Ivarsso

ory in 3‐point

the simulatio

s of knee ben

the compa

en the exper

s of ankle joi

aluation

simulations

ed pedestria

bia and fibul

endix shows

f 0.016 mm/

tudy and the

etermined fo

moment‐angl

n et al. [26]

t bending of

on result usi

nding mome

rison of the

riment from

int rotation.

were succes

an model is

a in leg com

the compar

/s and 1600

e results of t

or pPCL at 0.0

le response i

and the mod

the knee be

ng the modi

ent in Fig. 1

3‐po

e moment‐a

Crandall et a

ssfully run u

s illustrated

pression.

rison of the f

mm/s betwe

the modified

016 mm/s be

in 4‐point va

dified model

etween the t

fied model.

16. Validatio

oint valgus be

angle respon

al. [33] and t

up to the he

in Fig. A7 i

force‐deflect

een the resp

d model sim

ecause only o

algus bending

simulation r

est data from

n results of k

ending.

nse in dorsif

the modified

ead contact

n the Appen

tion respons

ponse corrid

ulations. Th

one test was

g of the kne

result. Figure

m Bose et al.

knee shear f

flexion/plan

d model simu

with the ca

ndix for the

se of individu

ors and failu

e corridor a

s conducted.

e between t

e 16 compar

. [32] scaled

force in

tarflexion a

ulation.

ar model. T

e three impa

ual

ure

nd

.

he

res

to

nd

he

act


- 539 -

IV. DISCUSSION

The results of the biofidelity validation generally showed that the impact response of the modified human FE

model is more biofidelic at the component level, except the upper limb, for which validation was not conducted.

The upper limb joint models were simplified using joint elements to ensure the numerical robustness at

high‐speed impacts. The elimination of the contact definitions between the rigid‐body cervical and thoracic

vertebrae yielded the head response without unfavorable high frequency oscillations, as shown in Fig. A1 and

Fig. A2 in the Appendix. The modifications to the thorax and the ankle models, validated against experiment,

allow the estimation of impact responses to these body regions. The additional validation of the knee model in

3‐point bending ensured that the model is capable of predicting responses to combined bending and shear

loads. The change in the ligament model structure to the bar models representing the fibers and the solid

models representing the matrix eliminated unrealistic compressive load due to the use of the tension‐only bar

elements, while maintaining a similar level of biofidelity compared to the baseline model. Although the material

property of the interosseous membrane requires validation, the validation of the load distribution between the

tibia and the fibula within the leg is important when local failure of a particular bone needs to be predicted. The

validation of the leg model in 3‐point bending showed that the model prediction almost traced the lower bound

of the experimental corridor from Ivarsson et al. [26]. However, the stiffness of the later phase of loading,

where the bone resistance dominates, is similar to the experimental data. This ‘shift’ of the curve in deflection

can be attributed to the decreased thickness of the leg flesh between the bone and the ram relative to the

thickness in a standing position due to the amputated muscles hanging down the horizontally placed bone. The

impact simulations at 70 km/h succeeded in calculating the kinematics of a pedestrian up to the head contact

with the car without any errors terminating the computation, suggesting that the model can be used to

simulate almost the entire range of the impact scenarios seen in real‐world accidents in terms of the impact

speed.

The knee model validations in 3‐point and 4‐point bending assumed the slack strain of 0.2 due to the lack of

preconditioning of the ligaments in the experiment, as opposed to the tensile tests of individual ligaments

where the ligaments were preconditioned prior to the tests to failure [30‐31]. The assumed initial slack strain

needs further clarifications when experimental data showing the influence of the preconditioning is available.

In order to ensure the numerical robustness of the model up to 70 km/h impact, it was necessary to simplify

the joint models in the upper limb due to the load concentration at the joints. In contrast, the Pedestrian Crash

Data Study database [35] shows only 10 joint injuries out of 858 injuries to the upper limb. This suggests that

the concentrated load to the upper limb joints may not be realistic possibly due to the lack of the muscle model

in the upper limb and its activation upon impact. The lack of bone fracture representation may have resulted in

unrealistic loads to the joints. Further studies are needed to clarify the mechanism of injuries to the upper limb.

As the car front‐end shape and structure would greatly influence car–pedestrian interactions, further

investigation will be needed to ensure numerical robustness of the model against different types of cars.

The addition of the musculature around the back of the knee decreased the peak hyperextension angle of

the knee from 39.5 degrees to 29.8 degrees, showing that the contribution of the passive musculature is

significant in the hyperextension of the knee in a frontal collision. However, the knee response in

hyperextension has not been validated and needs further validations. In addition, the neck model validation

used rigid constraint of the bottom of the neck musculature and skin, which may influence the neck stiffness

and requires further clarifications in a future study. The rigid foot model may need to be further improved in

order for the model to be applicable to various orientation and angle of the ankle upon impact.

V. CONCLUSIONS

A full‐body FE model for a pedestrian was developed by combining the upper body and the pelvis and lower

limb models developed in previous studies. The modifications were made to the neck, upper limb, thorax, knee,

leg and ankle in order to enhance the biofidelity and the numerical robustness of the model. The validation of

the biofidelity of the model for the modified components showed good match with the experimental data,

suggesting that the model is capable of predicting impact responses to the entire body regions, except the

upper limb, where the joint models were simplified to improve the numerical robustness. The impact

simulations against a simplified car model representing a stiff front‐end structure at 70 km/h showed that the


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model is capable of calculating pedestrian kinematics up to head contact to the car over the entire impact

speed range of real‐world accidents.

VI. REFERENCES

[1] European Enhanced Vehicle‐safety Committee. EEVC Working Group 17 Report – Improved test methods to evaluate pedestrian protection afforded by passenger cars, EEVC, 1998.

[2] Fredriksson, R., Rosén, E. Integrated Pedestrian Countermeasures ‐ Potential of Head Injury Reduction Combining Passive and Active Countermeasures. Proceedings of IRCOBI Conference, 2010, Hanover (Germany).

[3] Wang, J. T. Phase II Plan & Status of the Global Human Body Models Consortium. SAE Government & Industry Meeting, 2014, Washington, DC (USA).

[4] Watanabe, R., Katsuhara, T., Miyazaki, H., Kitagawa, Y., Yasuki, T. Research of the Relationship of Pedestrian Injury to Collision Speed,Car‐type, Impact Location and Pedestrian Sizes using Human FE model (THUMS Version 4). Stapp Car Crash Journal, 2012, 56: 269–321.

[5] Takahashi, Y., Suzuki, S., Ikeda, M., Gunji, Y. Investigation on Pedestrian Pelvis Loading Mechanisms Using Finite Element Simulations. Proceedings of IRCOBI Conference, 2010, Hanover (Germany).

[6] Ikeda, M., Suzuki, S., Gunji, Y., Takahashi, Y., Motozawa, Y., Hitosugi, M. Development of an Advanced Finite Element Model for a Pedestrian Pelvis. Proceedings of 22nd ESV Conference, 2011, Washington D.C. (USA).

[7] Dokko, Y., Yanaoka, T., Ohashi, K. Validation of Age‐specific Human FE Models for Lateral Impact. Proceedings of SAE World Congress, 2013, Detroit (USA).

[8] Ito, O., Dokko, Y., Ohashi, K. Development of Adult and Elderly FE Thorax Skeletal Models. Proceedings of SAE World Congress, 2009, Detroit (USA).

[9] Bartleby.com. “Henry Gray (1825–1861). Anatomy of the Human Body. 1918.” Internet: http://www.bartleby.com/107/indextn3.html, accessed 24 March 2015.

[10] Takahashi, Y., Kikuchi, Y. Biofidelity of test devices and validity of injury criteria for evaluating knee injuries to pedestrians. Proceedings of 17th ESV Conference, 2001, Amsterdam (The Netherlands).

[11] Shin, J., Lee, S., et al. Development and validation of a finite element model for the Polar‐II upper body. Proceedings of SAE World Congress, 2006, Detroit (USA).

[12] National Highway Traffic Safety Administration. Laboratory Test Procedure for FMVSS 208, Occupant Crash Protection Sled Tests, TP‐208S‐01.

[13] Hadjipanayi, E., Mudera, V., Brown, R. A. Guiding cell migration in 3D: A collagen matrix with graded directional stiffness. Cell Motility and the Cytoskeleton, 2009, 66(3):121–8.

[14] Susilo, M. E., Roeder, B. A., Voytik‐Harbin, S. L., Kokini, K., Nauman, E. A. Development of a three‐dimensional unit cell to model the micromechanical response of a collagen‐based extracellular matrix. Acta Biomaterialia, 2010, 6(4):1471–86.

[15] Lewis, G., Shaw, K. M. Modeling the tensile behavior of human Achilles tendon. Bio‐Medical Materials and Engineering, 2012, 7(4): 231–44.

[16] Bertram, J. E., Polevoy, Y., Cullinane, D. M. Mechanics of avian fibrous periosteum: tensile and adhesion properties during growth. Bone, 1998, 22(6):669–75.

[17] Institute for Traffic Accident Research and Data Analysis. 2013 Traffic Accident Statistics (in Japanese), 2011.

[18] Bizot, P., Meunier, A., Chrisfel, P., Witvoet, J. Experimental Passive Hyperextension Injuries of the Knee; Biomechanical Aspects of Their Consequences (in French). Revue de Chirurgie Orthopedique et Reparatrice de L'appareil Moteur, 1995, 81(3):211–20.

[19] Lloyd, D. G., Buchanan, T. S. Strategies of muscular support of varus andvalgus isometric loads at the human knee. Journal of Biomechanics, 2001, 34:1257–67.

[20] Ewing, C., Thomas, D. Human head and neck response to impact acceleration. NAMRL Monograph, 1973, 21.

[21] Thunnissen, J., Wismans, J., Ewing, C., Thomas, D. Human Volunteer Head‐Neck Response in Frontal Flexion: A New Analysis. SAE Technical Paper, 1995, Paper No. 952721.

[22] Kroell, C., Schneider, D., Nahum, A. Impact Tolerance and Response of the Human Thorax. SAE Technical Paper, 1971, Paper No. 710851.


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[23] Kroell, C., Schneider, D., Nahum, A. Impact Tolerance and Response of the Human Thorax II. SAE Technical Paper, 1974, Paper No. 741187.

[24] Eppinger, R. Prediction of Thoracic Injury Using Measurable Experimental Parameters. Proceedings of the International Conference on Experimental Safety Vehicles, 1976, London (UK).

[25] Viano, D. Biomechanical Responses and Injuries in Blunt Lateral Impact. SAE Technical Paper, 1989, Paper No. 892432.

[26] Ivarsson, J., Lessley, D., et al. Dynamic Response Corridors and Injury Thresholds of the Pedestrian Lower Extremities. Proceedings of IRCOBI Conference, 2004, Graz (Austria).

[27] Funk, R., Rudd, R., Kerrigan, J., Crandall, J. Analysis of Tibial Curvature, Fibular Loading, and the Tibia Index. Proceedings of IRCOBI Conference, 2003, Lisbon (Portugal).

[28] Takebe, K., Nakagawa, A., Minami, H., Kanazawa, H., Hirohata, K. Role of the fibula in weight‐bearing. Clin. Orthop. Relat. Res., 1984, 184:289–92.

[29] Shuler, F., Dietz, M. Tibial and Fibular Shaft Fractures. In Essential Orthopaedics, p. 703, Elsevier, Philadelphia, USA, 2009.

[30] Bose, D., Sanghavi, P., Kerrigan, J., Madeley, N., Bhalla, K., Crandall, J. Material Characterization of Ligaments Using Non‐contact Strain Measurement and Digitization. International Workshop on Human Subjects for Biomechanical Research, 2002, Ponte Vedra Beach (USA).

[31] van Dommelen, J., Ivarsson, B., et al. Characterization of the Rate‐Dependent Mechanical Properties and Failure of Human Knee Ligaments. Proceedings of SAE World Congress, 2005, Detroit (USA).

[32] Bose, D., Bhalla, K., Rooij, L., Millington, S., Studley, A., Crandall, J. Response of the Knee Joint to the Pedestrian Impact Loading Environment. Proceedings of SAE World Congress, 2004, Detroit (USA).

[33] Crandall, J., Portier, L., et al. Biomechanical Response and Physical Properties of the Leg, Foot, and Ankle. SAE Technical Paper, Paper No. 962424, 1996.

[34] Asanuma, H., Takahashi, Y., Ikeda, M., Yanaoka, T. Investigation of a Simplified Vehicle Model that Can Reproduce Car‐Pedestrian Collisions. Proceedings of SAE World Congress, 2014, Detroit (USA).

[35] National Highway Traffic Safety Administration. National Automotive Sampling System, Pedestrian Crash Data Study, 1994–1998.


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F

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Fig. A1. Head

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V

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VII. APPENDI

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flexion.

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F

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ment bundles in tension.


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F

Fig. A7. Kinematics of pedestrian model during immpact from ssimplified car model at 770 km/h.


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Documents

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