FOOT AND ANKLE SAFETY EVALUATION IN REAL LIFE CRASH … · 2016-10-04 · FOOT AND ANKLE SAFETY...

FOOT AND ANKLE SAFETY EVALUATION IN REAL LIFE CRASH SITUATIONS

Niklas Höglund1, Per Lövsund1, David Viano1, 2, Stefan Olsén2

1. Chalmers University of Technology

2. Saab Automobile AB

ABSTRACT

The relation between crash severity and lower leg responses are studied using mathematical

models, not only with the 50th percentile Hybrid III dummy, but also using biofidelic feet model and a

female dummy model. A number of countermeasures, including toepan padding and foot airbag, were

evaluated in a broad spectrum of crash situations, occupant sizes and foot positions.

The simulations showed that an airbag in the foot area optimized for the most severe crash

reduced forces higher than 3.5 kN in the right tibia in most crash situations. However, tibial axial

forces were higher in 50% of the simulations with a foot bag optimized for a severe crash than without

foot air bag in various crash conditions. The biofidelic foot and ankle model rotated less than the

Hybrid III foot. The difference in tibia load was relatively small between the dummy model and the

human foot model.

The study shows that an implementation of a countermeasure in cars has to be preceded by an

investigation of dummy responses in several crash situations to avoid sub-optimization of the car

safety performance.

Keywords: Foot and ankle, mathematical modelling, parametric investigation, safety device

FOOT AND ANKLE INJURIES are not fatal but are still common in frontal car collisions

(Kuppa et al., 2001) and often result in disabling injuries (MacKenzie et al., 1993). It is therefore

considered necessary to reduce the frequency and the injury severity. Short drivers and females have

been proposed for higher risk for foot and ankle injuries than large drivers and males in

epidemiological studies by e.g. Dischinger et al. (1994) and Crandall et al. (1996).

The injury mechanism for the most severe injuries is a load in the axial direction of the tibia

(Manning et al., 1998). Critical axial tibial load is 5.2 kN and 3.75 kN for 50th percentile male and

5th percentile female respectively (Kuppa et al., 2001). A hyper dorsiflexion of the ankle joint also

influence injury outcome, especially talus and malleolar fractures (Begeman et al., 1993,

Manning et al., 1998).

Foot and ankle injuries are more likely to occur in crashes with large delta-v (Fildes et al., 1995,

Richter et al., 2001). Toepan intrusion and pedals have been associated with lower limb injuries

(Thomas et al., 1995, Crandall et al., 1998). A more throughout literature review is presented in

Höglund et al. (2003). Previous studies have shown that there is a complex relation between different

compartment deformation characteristics and lower leg responses (Höglund et al., 2003,

Höglund et al., 2002). A foot airbag has shown to be effective reducing tibia index and tibial axial

forces in a series of sled tests by Håland et al. (1998), Kippelt et al. (1998), and Hesse et al. (2001).

The Hybrid III dummy is the most commonly used crash test dummy, but the foot and ankle

joints lack in biofidelity. It is thus a need for a better understanding of the mechanisms causing lower

leg using more detailed models. The objective of this study is to evaluate the safety performance of a

car and countermeasures in situations similar to real life crashes.

METHODOLOGY

An occupant-vehicle model of a mid-size, left hand side drive, production car was developed and

compared to crash test data and sled test data. The toepan and carpet were then replaced by a toepan

from a conceptual model. The model with the conceptual toepan will be hereby being denominated as

IRCOBI Conference – Lisbon (Portugal), September 2003 163

the baseline model. The baseline model was used in parametric studies, divided in two parts;

evaluation of the safety performance of the car model and an evaluation of countermeasures

(Figure 1).

Definitions Study 1 Study 2a, parameters with fixed levels Study 2b, stochastic

levels of the parameters7 Crash severities Baseline padding 3 Paddings Real-life like

crash situationsRandomly selected

paddings Real-life like crash situations

Real-life like crash

situations

Baseline Pedal 2 Occupant sizes3 Foot airbags

Foot model evaluation 3 Foot positions Foot model evaluation 4 Pedals Randomly selected

airbags Hybrid III dummyFoot model

evaluation Evaluation of the safety performance of the car model

Evaluation of countermeasures in real-life crash situation models Human foot model

Fig. 1 - Design of the parametric studies

MODELS

In the baseline vehicle model, the seat, the bracket of the knee bolster and steering wheel and column

were multi-body. Toepan, carpet, frontal airbag, foot airbag and knee panel were finite elements.

Acceleration pulses and intrusion data of the toepan were obtained from full scale finite element

simulations with validated car structural models. The intrusion was applied as pre-scribed motion of

each node of the toepan and carpet. The toepan intrusion from the structural fem models were

converted to carpet intrusion by pre-simulations in which the carpet was substituted with a material

having a low elasticity modulus. Then the substituted carpet was forced to fall on the toepan. In the

end of the pre-simulations, the carpet was tight to the toepan and each node of the carpet was replaced

with the corresponding node number of the toepan. The relative intrusion of these nodes was used as

intrusion of the carpet, with start at the original carpet position. The baseline model was compared to

crash test data. The characteristics of the seat and the crush angle of the knee panel were similar as in

the baseline model in Höglund et al. (2002). The contact definition between the dummy and airbag

and between dummy and pedal were also similar as in the baseline model in Höglund et al. (2002).

The motion of the pedal in the baseline model was estimated from crash test films. The seat belts were

standard MADYMO belt elements with pretensioner.

Fig. 2 - The baseline model

(left) and the toepan (right)

STUDY 1, EVALUATION OF THE MODEL IN REAL-LIFE CRASH SITUATIONS

The evaluation of the protection ability of the car in different crash situation was done by

comparing the responses of different dummies, in different positions, in a number of crash situations

which are listed in Table 1. The impact velocities can not be published, but velocity B is higher than

velocity A and so on. The crash severities were selected to represent a broad spectrum of crash

situations with six different impact velocities, two barrier configurations and three impact directions.

Each crash severity got its unique intrusion of the toepan. The intrusions of a node below the brake

pedal for each crash situation are listed in Table 1. There were large differences in intrusion between

different crash situations. For example, the intrusion distance in crash #4 was 0.08 m and peak

acceleration 1685 m/s2. The peak acceleration in crash #5 was only 833 m/s2 for 0.11 m intrusion.


Table 1: Crash severities in the parametric studies

Impact Toepan intrusion

Number Reference Barrier Direction

Onset time

[ms]

Distance

[m] 1

Peak velocity

[m/s] 1

Peak acceleration

[m/s2] 1

1 Vel. A, FFRB Rigid Full frontal - 0.01 (62) 0.2 (58) 98 (21)

2 Vel. B, FFRB Rigid Full frontal 45 0.05 (59) 4.7 (50) 1005 (46)

3 Vel. B, 30°RB Rigid 30° 40 0.05 (74) 3.8 (58) 472 (55)

4 Vel. C, FFRB Rigid Full frontal 35 0.08 (69) 6.4 (39) 1685 (36)

5 Vel. D, ODB Deformable 40% offset 65 0.11 (101) 5.8 (77) 833 (66)

6 Vel. E, ODB Deformable 40% offset 59 0.17 (100) 10.0 (66) 1616 (61)

7 Vel. F, ODB Deformable 40% offset 49 0.29 (101) 11.8 (63) 1883 (51) 1 First number is peak and the second number is the time for peak in milliseconds

The standard MADYMO hybrid III dummies (TNO, 1999) were used in the simulations,

equipped with standard feet or biofidelic foot models. Two occupant sizes in three foot positions were

investigated (Table 2). The biofidelic lower leg, foot and ankle model was developed by

Hall et al. (1998) and improved by Dubbeldam et al. (1999). The biofidelic foot modelled the major

bones, joints and ligaments in the foot. The tendons were passive with loading functions according to

Hall et al. (1998) and Dubbeldam et al. (1999). Originally, the foot modelled only the right foot of a

50th percentile male without shoes. In this study, the model was mirrored to the left lower leg,

implemented in the Hybrid III and equipped with shoes, similar as in the MADYMO Hybrid350

dummy (TNO, 1999).

Table 2: Occupant size, model complexity and foot position in the parametric studies Parameter Reference Level

HIII, 50th M Hybrid III, 50th percentile male, standard foot model

HIII, 5th F Hybrid III, 5th percentile woman, standard foot model

Human, 50th M Hybrid III, 50th percentile male with biofidelic foot model

Occupant

model

Human, 5th F Hybrid III, 5th percentile woman with biofidelic foot model

Baseline Straight on the accelerator and on the foot rest, as in Euro-NCAP

Angled Both feet moved inward and the feet angled outward to the accelerator and foot rest Foot

position Brake pedal Right foot positioned on the brake pedal

The feet were scaled from 50th percentile male to 5th percentile woman using scale factors. Scale

factors for length in x, y and z direction was presented by Mertz et al. (1989). Mertz et al. (1989)

suggested the force ratio is based on the scale factor for dimension in x and z-direction, which ends up

with a force ratio of 0.68. However, a more recent study by Yoganandan et al. (2000) showed that the

scale factors for load and moment should be 0.63 and 0.5 respectively and these scale factors were

used in this study.

No scale factor for moment of inertia was found in the literature but a simple unit comparison

was conducted as approximation. The unit of moment of inertia is [kgm2] and thus λix, iy, iz=λmass⋅λ2x, y, z.

The shoes of the female dummy were similar as the Hybrid III, 5th percentile female shoes. The scaled

model was not compared with test data. The scale factors are listed in Table 3.

Table 3: Scale factors for the human foot model, 5th percentile female Property Scale factor Property Scale factor

λx 10.85 λmass

10.545

λy 10.85 λix; λiy; λiz 0.39; 0.39; 0.35

λz 10.80 λload

20.63

λmoment 20.5

1. Mertz et al. (1989)

2. Yoganandan et al. (2000).

STUDY 2A, EVALUATION OF COUNTERMEASURES, FIXED LEVELS

The objective of the simulations in this part of the study was to evaluate the model under real-life

safety conditions and investigate countermeasures; different characteristics of toepan padding, foot

airbag and pedal release (Figure 3). The airbag model was just a prototype and thus, all properties

related to the airbag design, such as gas inflation and vent hole are not relevant. However, a foot


airbag causes a motion of the carpet, which is independent of the airbag design. For that reason, the

bag was replaced with a finite element shell with a pre-scribed motion similar to the top of the bag.

That enabled studies of different airbag motions. This simplification reduced the simulation time

greatly. The airbag model was made of steel and deformed plastically in the simulations. The

simulation time was 6.5 hours. Without bag, but with a pre-motioned shape, the simulation time was

reduced to approximately 5 minutes for the Hybrid III, 50th male and 15 minutes for the Hybrid III,

5th female. Four different pedals were investigated as shown in Figure 3. The baseline pedal (pedal0)

did not release. Pedal1 released at 1.7 kN and descended towards the toepan. Pedal2 released and

moved linearly towards the toepan according to a load-displacement translational joint and pedal3

released early in the crash event.

0

0.02

0.04

0.06

0.08

0.1

0 0.02 0.04 0.06

Time [s]

Dis

pla

ce

me

nt

[m]

Footbag1Footbag2Footbag3

0E+0

2E+7

4E+7

6E+7

0 0.5

Strain

Str

ess [

N/m

2]

1

Padding0Padding1Padding2

0

2000

4000

6000

0.00 0.05 0.10

Displacement [m]

Fo

rce

[N

]

Pedal 0Pedal 1Pedal 2Pedal 3

Fig. 3 - Foot airbag (left), toepan padding (centre), pedal joint load characteristics (right)

All combinations of the parameters listed in Table 1 and Table 2 and shown in Figure 3 were

simulated.

STUDY 2B, EVALUATION OF COUNTERMEASURES, STOCHASTIC APPROACH

In order to get an improved understanding of the influence of foot bag characteristics and

padding material properties on the responses, 873 additional simulations were run using the baseline

model with the padding characteristics randomly selected. Another 181 simulations were conducted

with an added airbag and the foot airbag characteristics were stochastically chosen. The variables were

inflation onset time, motion of the top of the bag (simulates inflation volume) and end inflation time.

The crash situation and foot position were selected according to a random function in both

simulation series. All randomly selected variables were evenly distributed. In this part of the study, all

simulations were conducted with the Hybrid III 50th percentile male dummy with standard foot model.

That enable simulations with angled foot and the CPU time was significantly reduced.

RESULTS

The first 5 ms of the response-history curves were removed to leave out initial effects. The

simulations with the human foot initially angled indicated that the model was not realistic in such

situations and these results are not presented in Figure 4 or appendix 2-5 and accordingly, no

conclusions are drawn from these results. Several simulations with the human foot model, especially

for crash #7, were unstable due to numerical errors. In general, the difference in axial tibial load

between simulations with the human-like foot model and the Hybrid III dummy foot were smaller than

the differences in ankle and foot rotation.

STUDY 1, EVALUATION OF THE MODEL IN REAL-LIFE CRASH SITUATIONS

Figure 4 shows that the right axial tibial force in Hybrid III and the human foot model were

almost equally influenced of crash situations for both dummy sizes. The right axial tibial force was in

general lowest for crash #5, an offset crash, and crash #3, an angled impact. The highest forces

occurred for crash #4, frontal crash.

The male dummy, both with the standard foot model and the human foot model was least

influenced of the crash situation when the foot was initially positioned on the brake pedal. The male

dummy was least influenced of the foot position in crash #5, both with the human foot model and the

standard Hybrid III foot model.


The female dummy varied more than the male dummy between different crash situations, foot

positions and dummy foot biofidelity. For example, there was a rather large difference in right tibia

load between the female dummy with the foot initially angled on the accelerator than in the baseline

position for the most severe offset crash, crash #7. The foot slipped off the pedal and hit the toepan at

a time with high toepan intrusion velocity. The axial load in the right tibia in the female dummy

increased with impact velocity when the right foot was positioned on the brake pedal.

The left axial tibial forces were in general lower than the right tibial forces (Appendix 2). The

differences in axial force across the crash situations turned out to be relatively small for the Hybrid III,

50th percentile male. The left axial tibial force in the Hybrid III, 5th percentile female with the foot

initially angled, varied considerable between the crash conditions. With the left foot angled, the female

dummy had higher responses for crashes with angled impact direction or offset impact than full frontal

impact, i.e. higher force with a lateral motion than without.

0

1000

2000

3000

4000

Vel. A, FFRB Vel. B, FFRB Vel. B, 30°RB Vel. C, FFRB Vel. D, ODB Vel. E, ODB Vel. F, ODB

Axia

l fo

rce

[N

]

HybridIII, 5th Female, Baseline HybridIII, 5th Female, Angled HybridIII, 5th Female, Brakepedal

HumanIII, 5th Female, Baseline HumanIII, 5th Female, Brakepedal

0

1000

2000

3000

4000

5000

6000

7000


Axia

l fo

rce

[N

]

HybridIII, 50th Male, Baseline HybridIII, 50th Male, Angled HybridIII, 50th Male, Brakepedal

HumanIII, 50th Male, Baseline HumanIII, 50th Male, Brakepedal

Fig. 4 - Axial force in right tibia for different foot positions, dummy foot and crash severity for

5th percentile female dummy (top) and 50th percentile male dummy (bottom)

Almost all dummy models dorsiflexed as most in the right ankle when the foot was positioned on

the brake pedal. The peak dorsiflexion angle was significantly higher in the Hybrid III foot than the

human foot (Appendix 3). The dorsiflexion angle in the human foot model was never more than 37°.

The human foot was also very stiff in the tarsal joint and everted/inverted just a few degrees

(Appendix 4).

STUDY 2, EVALUATION OF COUNTERMEASURES

Three different countermeasures were evaluated; toepan padding, foot airbag and pedal

deformation characteristics. The difference in axial force in right tibia between padding0 (baseline

padding), padding1 and padding2 was small (Appendix 5). However, the simulations with randomly

selected padding properties showed that it was possible to get lower axial force in this crash situation

using other padding materials. The lines in Figure 5 show the interval between maximum and

minimum of the simulations with randomly selected padding characteristics. The lowest force in

crash #4 was 4600 N for the most optimum padding property. Of the investigated padding materials,

the highest force in crash #4 was 6800 N. For the crash situation #6, the difference between highest

force and lowest force was 2600 N and for crash #5, the difference just 500 N.


0

2000

4000

6000

8000

Vel. A,

FFRB

Vel. B,

FFRB

Vel. B,

30°RB

Vel. C,

FFRB

Vel. D,

ODB

Vel. E,

ODB

Vel. F,

ODB

Axia

l fo

rce

[N

]

Fig. 5 - The lines show the interval between

maximum and minimum of the simulations

with randomly selected padding

characteristics (onset inflation time,

distance, end inflation time). The bars

represent right tibia axial load with baseline

pedal, padding and foot position.

The airbag in the foot area was represented by a motion of the carpet. Figure 6 shows the average

axial force in right tibia for the Hybrid III, 50th percentile male and 5th percentile female, in different

foot positions for different foot airbags. Bag #1 and #2 reduced the axial force in the Hybrid III,

50th percentile male in baseline position. When the foot was angled on the accelerator or initially

positioned on the brake pedal, the force was slightly higher with foot airbag. The axial force in the

female dummy was uninfluenced of bag 1 in the baseline foot position, but the force was substantial

higher in the combination of angled position and foot airbag.

0

2000

4000

6000

305,

Baseline

305,

Angled

305,

Brake

pedal

350,

Baseline

350,

Angled

350,

Brake

pedal

Axia

l fo

rce

[N

]

Bag1 Bag2 Bag3 Baseline countermeasure

Fig. 6 - Average axial force in

right tibia for the Hybrid III,

50th percentile male and

5th percentile female in different

foot positions for different foot

airbags. The dotted bars

represent simulations with the

same countermeasures as in the

baseline model.

The deviation in forces among the simulations with randomly selected foot airbag characteristics

was large (Appendix 6). Figure 7 shows the average axial force in right tibia for randomly selected

foot bag properties in the model with baseline pedal, padding and foot position. High axial forces

occurred in the right tibia in crash situation #4. The load in tibia was 2900 N with an optimized foot

bag, which was a reduction with 50% compared to the simulations with baseline padding.

0

2000

4000

6000

8000

10000

12000

Vel. A,

FFRB

Vel. B,

FFRB

Vel. B,

30°RB

Vel. C,

FFRB

Vel. D,

ODB

Vel. E,

ODB

Vel. F,

ODB

Axia

l fo

rce

[N

]

Fig. 7 - The lines show the interval between

maximum and minimum of the simulations

with randomly selected foot bag properties

(onset inflation time, distance, end inflation

time). The bars represent right tibia axial

load with baseline pedal, padding and foot

position.

The difference in tibia load for different pedals was small. However, the average right ankle

rotation and right foot rotation were slightly reduced with pedal 3, a pedal that deforms early in the

crash event (Figure 8). This was especially obvious for right ankle eversion for the 50th percentile male

with the foot angled on the accelerator.


0

10

20

30

40

Pedal0 Pedal1 Pedal2 Pedal3

Do

rsifle

xio

n a

ng

le [

de

g]

HybridIII, 5th F, BaselineHybridIII, 5th F, AngledHybridIII, 50th M, BaselineHybridIII, 50th M, AngledHumanIII, 5th F, BaselineHumanIII, 50th M, Baseline

-40

-30

-20

-10

0

10

20

30

40

50

Pedal0 Pedal1 Pedal2 Pedal3Eve

rsio

n (

-)/I

nve

rsio

n (

+)

[de

g] HybridIII, 5th F, Baseline

HybridIII, 5th F, AngledHybridIII, 50th M, BaselineHybridIII, 50th M, AngledHumanIII, 5th F, BaselineHumanIII, 50th M, Baseline

Fig. 8 - Average right dorsiflexion angle (left) and eversion/inversion (right) for different pedal

characteristics. Pedal0 is the baseline pedal.

DISCUSSION

An evaluation of the safety performance of a car in real-life crash situations has been performed

with mathematical dummy-vehicle models. A number of simulation series, in total almost 3200 runs,

were made using the models to evaluate the toepan design, safety systems and the dummy model in

different crash situations. The vehicle model was designed and compared to a number of crash tests.

Pelvis acceleration, tibial axial forces and bending moments showed acceptable agreement with crash

tests (Appendix 1). The femur forces agreed well in magnitude but the duration of the femur forces in

the model was shorter than the compared crash test dummy responses.

Previous studies by the authors have indicated that the responses depend on a large number of

parameters describing the toepan configuration. Some parameters influenced the responses differently

depending of the levels of other parameters and these interaction effects have to be taken into account

in lower limb protection evaluation (Höglund et al., 2003). This requires a large number of

simulations, and to reduce the CPU-time, the models have to consist of as few finite element parts as

possible. Most parts in the model were multi-body systems. The toepan, knee bolster, carpet, foot

airbag and frontal airbag were modelled as finite elements to get a shape during intrusion as similar as

possible to real life crash situations.

The model did not include a frontal airbag in the parametric studies. A brief study was conducted

to evaluate the influence of the frontal airbag. 28 combinations with human foot were selected

randomly from the test matrix. The models were simulated both with and without airbag. The

percentage difference with and without airbag were calculated for each lower limb response in all

combinations. The average difference between all peak responses was 1.5%. However, for the most

interesting responses, peak axial tibial forces, foot and ankle rotations and moments, the average was

less than 1%. These results correlate well with results by Crandall et al. (1997) who demonstrated that

airbags have not reduced the frequency of lower leg injuries. Since the difference in responses with

and without frontal airbag was small, the parametric studies were conducted without frontal airbag to

reduce the CPU time.

An important advantage with this study was the use of a biofidelic dummy foot model

(Hall et al., 1998, Dubbeldam et al., 1999). The model consisted of the most important bones, joints

and ligaments of the foot. Unfortunately, the knee joint was similar as in the Hybrid III dummy, a

combination of a translational joint and a revolute joint. Therefore, the external and internal rotation of

the lower leg could not be done in the knee-tibia-fibula-complex. These rotations took place in the

tarsal joints and therefore the ligaments pulled the foot back to the original position. The simulations

with the human like foot model angled on the accelerator and foot rest was then not as realistic as the

simulations with the feet in the baseline or brake pedal position. The results from simulations with the

human foot initially angled are therefore excluded.

The ligaments were modelled as belt elements and Maxwell elements. In some crash situations,

numerical errors occurred when any ligaments become zero in length. This problem occurred most

often for the short ligaments between medial and lateral malleolus and talus. The ligaments in the foot


strongly influenced the kinematics of the biofidelic foot model and the difference compared to the

Hybrid III foot model was particularly in foot rotation, ankle rotation and bending moments.

The axial force was measured in the same position, distal tibia, in both the standard Hybrid III

foot model and the human-like foot model, which reduced the influence of the difference in tibia shape

between the models. Despite differences in kinematics, peak axial forces in especially the

50th percentile male correlated fairly well in the brake pedal and baseline position. The axial loads

were in general higher in the human foot model than the Hybrid III model, which was due to the

Achilles tendon. To investigate the influence of the Achilles tendon, all crash situations with baseline

countermeasures and with the human-like foot model were compared with simulations without an

Achilles tendon. These simulations indicated that the Achilles tendon was the main source for the

difference between the human-like models and the Hybrid III models. For situations with the foot on

the pedal during the whole crash event, there were just small differences in axial load between the

Hybrid III dummy foot and the human like foot model. However, simulations with the female dummy

showed that the ankle rotation often were small, which caused small differences in axial load between

the models with and without a passive Achilles tendon.

The largest difference between the male foot models in baseline position was at low impact

velocities, crash #1 and #2, and in brake pedal position, at angled impact. In that situation, the peak

force occurred at impact to the toepan and in the human foot model, this correlated with a rather high

load in the Achilles tendon.

The axial tibial load in the female dummy with Hybrid III standard shoe was higher than the

dummy with the biofidelic foot model in crash #4. The differences in kinematics were small in that

crash situation, but large enough for the Hybrid III foot to impinge more in the padding than the

human-like foot.

Funk et al. (2002) found that malleolar fractures were common in eversion for ankles preloaded

with 2 kN. In tests with similar conditions and the ankles also initially dorsiflexed 30° resulted in talar

fractures in their study. Therefore, even though axial forces match fairly well, correct description of

the rotation of both the feet and the ankle are important in studies of the interaction between occupants

and vehicle for comparing dummy responses with real-life injuries. It would thus be interesting to

improve the human foot model to be stable in most crash conditions and with a biofidelic

external/internal rotation.

The evaluation of the car model in real-life crash situations showed that the responses were

highly dependent of the crash situation, and within a particular crash with a reasonable variation in

parameters defining the safety systems and occupant. The toepan intrusion was different in all crash

situations and there were no distinct correlation between any response and any parameter describing

the intrusion. With the baseline model and baseline position, the right ankle dorsiflexion in the

Hybrid III, 50th percentile male dummy correlated with intrusion acceleration following a equation of

the second degree (R2=0.91), but this was not the case for any other dummies or foot position. In some

simulations, the axial load was lower when the foot was positioned on the brake pedal, which

correlates with findings in a simulations study be Pilkey et al. (1994). In the baseline position,

50th percentile male dummy and baseline countermeasures, there were a correlation between peak

acceleration of the toepan and tibia load; the load were highest in the three crashes with highest toepan

acceleration. This correlates well with studies by Kuppa and Sieveka (1995) and Sakurai (1996)

Three safety systems were evaluated; foot airbags, toepan paddings and deformable pedals. The

examination of the influence of toepan padding indicated that the property of the padding significantly

influenced the axial load at crashes with high peak acceleration. Crashes with low peak acceleration,

as crash #5 were just slightly influenced of different padding properties. No correlation was found

between intrusion magnitude and the influence of different toepan paddings on tibia load.

The baseline motion of the foot airbag was given by the average node motion of a simulation

with a finite element foot bag. The results from the study indicated that a foot airbag have the potential

to reduce tibial forces. A foot airbag with onset time 12.2 ms, end inflation time 19.6 ms and total

motion, or thickness at full inflation, of 0.0795 m could be effective, especially in combination with

pedal #3. In order to evaluate these countermeasures in different crash situations, foot positions and

dummy sizes, a two-level factorial design study was done (Box et al., 1978). The effects showed that

the most important effects were dummy size, toepan configuration and the interaction between these

factors. The interaction in axial force between dummy and toepan configuration for each crash


situation are summarized in Table 4. As shown in the table, a foot airbag reduced the axial force in

right tibia for the 50th percentile male, baseline positioned, in most situations. For the 5th percentile

female, a foot bag increased the force significantly in four crash situations. However, when the right

foot was positioned on the brake pedal, a foot bag reduced the axial force in the female dummy but not

in the larger male dummy. A foot airbag optimized for a severe crash situation as full frontal in rather

large impact velocity got the potential to reduce axial loads in several crash situations. However, the

bag increased the tibial load in 50% of the simulations in the two-level analysis.

Table 4: Average right axial force in the two-level factorial analysis

Position Dummy

Counter-

measure

Vel. A,

FFRB

Vel. B,

FFRB

Vel. B,

30°RB

Vel. C,

FFRB

Vel. D,

ODB

Vel. E,

ODB

Vel. F,

ODB

Baseline HIII, 5th F Baseline 1383 89 1792 3885 469 1122 1844

Baseline HIII, 5th F Foot bag 1130 1051 1195 3055 2041 3349 3339

Baseline HIII, 50th M Baseline 4735 2888 5149 6065 2489 4792 5819

Baseline HIII, 50th M Foot bag 2112 2079 2257 2836 2800 4547 5962

Brake HIII, 5th F Baseline 405 853 485 1447 1533 3277 3342

Brake HIII, 5th F Foot bag 1085 537 1049 1122 2109 1719 2052

Brake HIII, 50th M Baseline 3008 1548 2792 2873 2588 3927 4311

Brake HIII, 50th M Foot bag 2067 2524 2579 3459 2958 4532 5417

The results in this study correlates well with findings in Hesse et al. (2001), who found in a series

of sled tests that a foot airbag reduced the axial force. However, this study demonstrates that the tibial

axial force become higher with bag than without bag in 50% of all simulations with a foot bag in the

two-level factorial analysis. Introducing a foot airbag is though the only method of the

countermeasures investigated in this study to get right tibia axial forces below 4500 N in a full frontal

crash, rigid barrier, as crash situation #4 keeping the similar toepan intrusion and foot positions. The

foot airbag have to be optimized for a number of crash situations and it would be advantageous to be

able to adjust the airbag inflation depending on the crash situation, driver stature and foot position.

CONCLUSIONS

The evaluation of the mathematical vehicle model in a number of crash situations showed that the

dummy responses varied significantly for different crash situations. Therefore, implementing foot and

ankle safety systems in cars have to be preceded by an investigation of its effects on occupant

responses in a broad spectrum of crash situations, occupant sizes and foot positions. The simulations

with a biofidelic dummy foot showed that the dorsiflexion and inversion/eversion of the ankle and foot

differed significantly compared to the Hybrid III foot, but the axial forces matched fairly well.

The pedal was not a major source for high axial tibial forces in this study, but the eversion of the

right foot was lower when the pedal deformed early in the crash event.

A foot airbag reduced axial tibial loads in some crash situations, but the simulations indicated

that the bag had to be optimized for a number of crash situations. This study showed that an airbag

also could increase the responses for some combination of occupant size, crash situation and foot

position. The axial tibial load was higher in fifty percent of all simulations with a foot airbag

compared to similar crash situation without airbag in this study.

None of the studied padding materials could reduce tibial forces below 4500 N in a full frontal

crash with rigid barrier at velocity C.

ACKNOWLEDGEMENTS

This work was funded by Saab Automobile AB and Programrådet för Fordonsteknisk Forskning

(PFF). The authors would like to thank Mrs Rosemary Dubbeldam at Delphi Automotive Systems and

Prof. Jeff R. Crandall at the Center for Applied Biomechanics at University of Virginia who developed

and provided the biofidelic foot model.


REFERENCES

Begeman, Paul; Balakrishnan, Pradeep; Levine, Robert; King, Albert I., ‘Dynamic human ankle

response to inversion and eversion’ 37th Stapp Car Crash Conference, San Antonio, Texas, USA,

1993, pp 83-93

Box, G. E. P.; Hunter, W. G.; Hunter, S. J. ‘Statistics for Experimenters’ 1978 John Wiley & Sons, Inc.

ISBN 0-471-09315-7.

Crandall, Jeff R.; Martin, Peter G.; Bass, Cameron R.; Pilkey, Walter D.; Dischinger, Patricia C.;

Burgess, Andrew R.; O'Quinn, Timothy D.; Schmidhauser, Carl B., ‘Foot and ankle injury: the

roles of driver anthropometry, footwear and pedal controls’ 40th Annual Proceedings of the

Association for the Advancement of Automotive Medicine, Vancouver, British Columbia, Canada,

1996 pp 1-18

Crandall, J. R.; Martin, P. G, ‘Lower limb injuries sustained in crashes and corresponding

biomechanical research’ International Symposium on Real World Crash Injury Research,

Leicestershire, England, 1997 pp 3.5.1-3.5.16

Crandall, Jeff R.; Martin, Peter G.; Sieveka, Edwin M.; Pilkey, Walter D.; Dischinger, Patricia C.;

Burgess, Andrew R.; O'Quinn, Timothy D.; Schmidhauser, Carl B, ‘Lower limb response and

injury in frontal crashes’, Accident Analysis and Prevention v30 n5 Sep 1998 p 667-677 0001-

4575 AAPVB5

Dischinger, P. C.; Kufera, J. A.; Kerns, T. J. ’Lower extremity fractures in motor vehicle collisions:

The role of driver gender and height’, 38th Annual Proceedings Association for the Advancement

of Automotive Medicine, Lyon, France, 1994 pp 601-605

Dubbeldam, R.; Nilson, G.; Pal, B.; Eriksson, N.; Owen, C.; Roberts, A.; Crandall, J.; Hall, G.;

Manning, P.; Wallace, A., ‘A MADYMO model of the foot and leg for local impacts’ 43rd

Annual Stapp Car Crash Conference, San Diego, California, USA, 1999, pp 185-202

Fildes, B.; Lenard, J.; Lane, J.; Seyer, K., Lower limb injury in frontal crashes. International

conference on pelvic and lower extremity injuries, Washington, DC USA, 1995 pp 1-10

Funk, J. R.; Srinivasan, S.C. M.; Crandall, J. R. ; Khaewpong, N; Eppinger, R. H.; Jaffredo, A. S.; ;

Petit, P. Y., ‘The effects of axial preload and dorsiflexion on the tolerance of the ankle/subtalar

joint to dynamic inversion and eversion’, 46th Stapp Car Crash Conference, Ponte Verdra Beach,

Florida, USA 2002, pp 245-265

Hall, G. W.; Crandall, J. R.; Pilkey, W. D.; Thunnissen, J. G., ‘Development of a dynamic multibody

model to analyze human lower extremity impact response and injury’ International IRCOBI

Conference on the Biomechanics of Impact, Göteborg, Sweden, 1998 pp 117-134

Hesse, S.; Bläßer, S.; Hoffmann, J., ‘Parameter study on different factors influencing lower extremity

injuries’ 17th International Technical Conference on the Enhanced Safety of Vehicles Amsterdam,

The Netherlands, 2001 Paper #101

Håland, Y.; Hjerpe, E.; Lövsund, P., ‘An inflatable carpet to reduce the loading of the lower

extremities - evaluation by a new sled test method with toepan intrusion’ 16th International

Technical Conference on the Enhanced Safety of Vehicles, Windsor, Ontario, Canada, 1998 pp

292-301

Höglund, N.; Lövsund, P.; Viano, D.C. ‘Parametric investigation of the toepan region using

mathematical models’ International Journal of Crashworthiness 2002 (Submitted)

Höglund, N.; Lövsund, P.; Viano, D.C. ‘Mathematical modelling of lower leg responses in offset

crashes’ Journal of Crashworthiness 2003 Vol 8 No 3, pp 285-295

Kippelt, Ulrich; Buss, Winfried; Feldhoff, Uwe; Thelen, Martin, ‘Protection devices and development

tools for reducing foot and leg injuries in frontal crashes’ International IRCOBI Conference on

the Biomechanics of Impact, Göteborg, Sweden, 1998 pp 161-172


Kuppa, S. M.; Sieveka, E. M., 'Dynamic motion of the floor pan and axial loading through the feet in

frontal crash tests', International IRCOBI Conference on the Biomechanics of Impact, Brunnen,

Switzerland, 1995, pp 389-402

Kuppa, S.; Wang, J.; Haffner, M.; Eppinger, R., ’Lower extremity injuries and associated injury

criteria’ International Technical Conference on the Enhanced Safety of Vehicles, Amsterdam,

The Netherlands, 2001, Paper #457

MacKenzie EJ, Cushing BM, Jurkovich GJ, Morris JA Jr, Burgess AR, deLateur BJ, McAndrew MP,

Swiontkowski MF. ‘Physical impairment and functional outcomes six months after severe lower

extremity fractures’ J Trauma. 1993 Apr;34(4):528-38

Manning, P.; Owen, C.; Roberts, A.; Oakley, C.; Lowne, R., ‘Dynamic Response And Injury

Mechanism In The Human Foot And Ankle And An Analysis Of Dummy Biofidelity’, 16th

International Technical Conference on the Enhanced Safety of Vehicles, Windsor, Ontario,

Canada, 1998, pp 1960-1998

Mertz, H. J.; Irwin , A. L.; Melvin, J. W.; Stalnaker, R. L.; Beebe, M. S.; ‘Size, weight and

biomechanical impact response requirements for adult size small female and large male

dummies’ SAE International Congress and Exposition, Detroit, Michigan, USA 1989, pp 133-

140

Pilkey, Walter D.; Sieveka, Edwin M.; Crandall, Jeff R.; Klopp, Gregory, ‘The influence of foot

placement and vehicular intrusion on occupant lower limb injury in full-frontal and frontal-offset

crashes’ 14th International Technical Conference on the Enhanced Safety of Vehicles, Munich,

Germany, 1994, pp 734-741

Richter, M.; Thermann, H.; Wippermann, B.; Otte, D.; Schratt, HE.; Tscherne, H., ‘Foot fractures in

restrained front seat car occupants: a long-term study over twenty-three years’, J Orthop Trauma

2001 May;15(4):287-293

Sakurai, M. 'An analysis of injury mechanisms for ankle/foot region in frontal offset collisions',

40th Stapp Car Crash Conference, Albuquerque, New Mexico, USA, 1996

Thomas, P; Charles, J.; Fay, P., ‘Lower limb injuries~The effect of intrusion, crash severity and the

pedals on injury risk and injury type in frontal collisions’ 39th Stapp Car Crash Conference, San

Diego, California, USA, 1995 pp 265-280

TNO (1999) ‘MADYMO User’s Manual 3D’ TNO Road-Vehicles Research Institute, Delft The

Netherlands

Yoganandan, N; Pintar, F. A.; Pintar, F. A.; Kumaresan, S; Kumaresan, S; Gennarelli, T; Gennarelli,

T. A.; Sun, E; Kuppa, S; Maltese, M; Eppinger, R. H., ’Pediatric and small female neck injury

scale factors and tolerance based on human spine biomechanical characteristics’, International

IRCOBI Conference on The Biomechanics of Impact, Montpellier, France 2000


APPENDIX

Gray shaded lines represents crash test 1, 2 and 3. Black line is the baseline MADYMO model

-700

-600

-500

-400

-300

-200

-100

0

100

0 25 50 75 100 125

Time [ms]

Pelv

is x

-acc [

m/s

2]

-500

-400

-300

-200

-100

0

100

200

0 25 50 75 100 125

Time [ms]

Pelv

is z

-acc [

m/s

2]

-2000

-1000

0

1000

2000

3000

4000

0 25 50 75 100 125

Time [ms]

Rig

ht

fem

ur

forc

e [

N]

-2000

-1000

0

1000

2000

3000

4000

5000

0 25 50 75 100 125

Time [ms]

Left

fem

ur

forc

e [

N]

-1000

0

1000

2000

3000

4000

5000

6000

0 25 50 75 100 125Time [ms]

Rig

ht

tib

ia f

orc

e [

N]

-1000

0

1000

2000

3000

4000

5000

6000

0 25 50 75 100 125Time [ms]

Left

tib

ia f

orc

e [

N]

-100

-50

0

50

100

150

0 25 50 75 100 125

Time [ms]

Lw

r ri

gh

t tib

ia M

y

[Nm

]

-80

-60

-40

-20

0

20

40

0 25 50 75 100 125

Time [ms]

Lw

r le

ft t

ibia

My

[N

m]

Appendix 1: Crash test and

MADYMO simulations with the

baseline models in full frontal rigid

barrier crash situation

0

1000

2000

3000

4000

5000

Vel. A, FFRB Vel. B, FFRB Vel. B,

30°RB

Vel. C, FFRB Vel. D, ODB Vel. E, ODB Vel. F, ODB

Axia

l fo

rce

[N

]

HybridIII, 5th F, Baseline HybridIII, 5th F, Angled HybridIII, 50th M, Baseline

HybridIII, 50th M, Angled HumanIII, 5th F, Baseline HumanIII, 50th M, Baseline

Appendix 2: Axial force in left tibia for different dummies and foot positions

0

10

20

30

40

50

60


Dors

ifle

xio

n [

deg]

HybridIII, 5th F, Baseline HybridIII, 5th F, Angled HybridIII, 5th F, BrakepedalHybridIII, 50th M, Baseline HybridIII, 50th M, Angled HybridIII, 50th M, BrakepedalHumanIII, 5th F, Baseline HumanIII, 5th F, Brakepedal HumanIII, 50th M, BaselineHumanIII, 50th M, Brakepedal

Appendix 3: Right ankle rotation, dorsiflexion, for different dummies and foot positions


-60

-40

-20

0

20

40

60

Vel. A, FFRB Vel. B, FFRB Vel. B, 30°RB Vel. C, FFRB Vel. D, ODB Vel. E, ODB Vel. F, ODBEve

rsio

n (

-),

Inve

rsio

n (

+)

[de

g]


Appendix 4: Right foot rotation, inversion/eversion, for different dummies and foot positions

0

1000

2000

3000

4000

5000

6000

7000

Padding0 Padding1 Padding2

Axia

l fo

rce

[N

]


Appendix 5: Average axial force in lower right tibia for different dummies, foot positions and toepan

padding materials.

0

2000

4000

6000

8000

10000

0 0.05 0.1 0.15Airbag displacement [m]

Axia

l fo

rce

[N

]

Appendix 6: Axial force in right tibia vs. airbag

displacement for all investigated airbags.


FOOT AND ANKLE SAFETY EVALUATION IN REAL LIFE CRASH … · 2016-10-04 · FOOT AND ANKLE SAFETY...

Documents

Transcript of FOOT AND ANKLE SAFETY EVALUATION IN REAL LIFE CRASH … · 2016-10-04 · FOOT AND ANKLE SAFETY...