Accident Analysis and Prevention 80 (2015) 97–105

Influence of stiffness and shape of contact surface on skull fractures andbiomechanical metrics of the human head of different populationunderlateral impacts

Debasis Shaoo a,*, Caroline Deck a,*, Narayan Yoganandan b, Rémy Willinger a

aUniversité de Strasbourg ICube, UNISTRA-CNRS, 2 rue Boussingault, 67000 Strasbourg, FrancebDepartment of Neurosurgery, Medical College of Wisconsin, 9200 West Wisconsin Avenue, Milwaukee, WI 53226, USA

A R T I C L E I N F O

Article history:Received 7 November 2014Received in revised form 27 March 2015Accepted 5 April 2015Available online xxx

Keywords:AnthropometrySkull fracture criterionFinite element head modelLateral impact

A B S T R A C T

The objective of this study was to determine the responses of 5th-percentile female, and 50th- and 95th-percentile male human heads during lateral impacts at different velocities and determine the role of thestiffness and shape of the impacting surface on peak forces and derived skull fracture metrics. A state-of-the-art validated finite element (FE) head model was used to study the influence of different populationhuman heads on skull fracture for lateral impacts. The mass of the FE head model was altered to matchthe adult size dummies. Numerical simulations of lateral head impacts for 45 cases (15experiments � 3 different population human heads) were performed at velocities ranging from 2.4 to6.5 m/s and three impacting conditions (flat and cylindrical 90D; and flat 40D padding). The entire force-time signals from simulations were compared with experimental mean and upper/lower corridors ateach velocity, stiffness (40 and 90 durometer) and shapes (flat and cylindrical) of the impacting surfaces.Average deviation of peak force from the 50th male to 95th male and 5th female were 6.4% and 10.6%considering impacts on the three impactors. These results indicate hierarchy of variables which can beused in injury mitigation efforts.

ã 2015 Elsevier Ltd. All rights reserved.

Contents lists available at ScienceDirect

Accident Analysis and Prevention

journal homepage: www.else vie r .com/ locate /aa p

1. Introduction

Nowadays with rapid growth of technology and protectivesystems, the number of accidents decreased, but still the numbersare high (ONISR Report, 2012). About 1.24 million people die eachyear as a result of road traffic crashes (WHO Report, 2013). Recentfield studies published after the year 2006 have shown that headinjuries continue to occur despite improvements in technology(Faul et al., 2010) and more importantly, skull fractures with andwithout brain injuries are attributed to contact loading in bothfrontal and side impact (Pintar et al., 2007; Zhang et al., 2006).Although more studies are available for frontal impacts than sideimpacts, there is almost a paucity of head injury studies focused onthe role of head anthropometry (Allsop et al., 1991; Yoganandan

* Corresponding authors. Tel.: +33 368852940; fax: +33 368852936.E-mail addresses: deb4uinl@gmail.com (D. Shaoo), deck.caroline@gmail.com

(C. Deck), yoga@mcw.edu (N. Yoganandan), remy.willinger@unistra.fr (R. Willinger).

http://dx.doi.org/10.1016/j.aap.2015.04.0040001-4575/ã 2015 Elsevier Ltd. All rights reserved.

et al., 2004, 2009a; Raymond et al., 2009). The role of head mass isnot investigated in detail.

Finite Element head models (FEHM) are proven to be verypromising tools to study head trauma biomechanics. However,studies using postmortem human surrogates (PMHS) are needed toobtain appropriate data for FEHM validation. Metrics associatedwith skull fracture such as peak force can be used during thevalidation process.Studies have been conducted to determinethese types of variables. Yoganandan et al. (1995) subjected12 intact PMHS heads to failure using a material testing device atloading rates up to 8.0 m/s. Vertex, parietal, temporal, frontal, andoccipital regions were the loading sites. The number of specimenssubjected to lateral impact loading washowever, limited. Yoga-nandan and Pintar (2004) later extended these studies bysubjecting PMHS heads to drop impacts, and a review is given.The authors underscored differences in regional biomechanicaltolerances of the human skull. Variations in regional skull fractureresponses stem from anatomical differences such as the bonethickness and local curvature of the skull in the lateral/parietal-lateral area compared to the frontal and occipital regions

98 D. Shaoo et al. / Accident Analysis and Prevention 80 (2015) 97–105

(Bass and Yoganandan, 2015). Although these studies provideimportant data such as peak forces for skull fracture, because of theassociated cost, ability to obtain experimental specimens andother resources, FEHM is an efficient alternative. Computationalmodels have the ease to describe the complex geometry, to assessthe injury sustained by different parts of the head in various impactconditions and to conduct parametric analysis. Parametric studiesare essential when developing new injury criteria, including onefor skull fracturebased on computational models (Ruan and Prasad2001). Frontal impact simulations were conducted by Ruan andPrasad (2001) to evaluate the effect of skull thickness on skullresponse. Four head models with different skull thickness wereimpacted by a cylinder with a mass of 5.23 kg and an initial velocityof 6.33 m/s. It was found that as skull thickness increased, skulldeformation decreased but this relationship cannot be linearlyinterpolated to the other parameters such as head acceleration. Aparametric study focusing on skull fracture for lateral/side impactsusing FEHM is needed and this is an aim of the present study.

There are limited studies which investigate the influence ofanthropometry and oblique side impacts to different impactingsurfaces. Different populations correspond to different physicalproperties for the head which does not only affect the biomechan-ics, but also influence the load transmitted to the head duringimpact in real-world accidents (Pintar et al., 2005; Yoganandanet al., 2009a,b). Hence the objectives of this study weretodetermine responses of 5th-percentile female and 50th- and95th-percentile male human heads during lateral impacts atdifferent velocities and determine the role of the stiffness andshape of the impacting surface on peak forces and derived skullfracture metrics.

2. Methods

2.1. Experimental data

Results from 17 unembalmed specimens isolated at the level ofthe occipital condyles from PMHS were used to conduct 86 droptests. The specimens included the head with intracranial contents(for some specimens sylgard gel was used as brain substitute). Theinstrumentation for biomechanical data acquisition consisted oftri-axial accelerometers at the vertex, anterior and posteriorregions of cranium, and a-nine accelerometer package (pyramid-shaped PNAP) was attached to the skull at the contra-lateral site ofimpact using screws as shown in Fig. 1a (Yoganandan et al., 2006).Repeated drop tests were conducted on the same specimen withsuccessively increasing input energies (increasing drop heights) tothe specimen until fracture. Three impacting boundary conditions,

Fig. 1. (a) Specimen preparation; (b) test setup

also termed as targets, were used: flat 40- and 90-durometerpadding (50 mm thickness), and cylindrical 90-durometer (50 mmdiameter) padding. The mid-sagittal plane of the specimen wasaligned at an angle of approximately 10 degrees with respect to thehorizontal plane such that the impact was focused on the lefttemporo-parietal regionas shown in Fig. 1b. Acceleration- andforce-time signals were collected using a digital data acquisitionsystem (TDAS Pro DTS Technologies, Seal Beach, CA) according toSAE J 211 specifications and processed using SAE Class 1000 filter.Peak resultant forces and center of gravity linear and angularaccelerations were obtained for each test. More information aboutthe experiments and the validation of skull model in terms oflateral impact and temporo-parietal skull fracture are reported inSahoo et al. (2013b).

2.2. FE head model and boundary conditions for simulation

A state-of-the-art validated FEHM, developed in StrasbourgUniversity (Deck and Willinger, 2008a,b) was used to study theinfluence of stiffness and shape of contact surface on skull fracturesand biomechanical metrics of the human head of differentpopulation under lateral impacts. The advanced model wasenhanced in terms of new constitutive material laws for brainand skull (Sahoo et al., 2013a,b). The previous FEHM wasequivalent to 50th percentile adult human head. The mainanatomical features included the scalp, brain, brainstem andcerebrospinal fluid (CSF), represented by brick elements and theskull, face and two membranes (the falx and the tentorium)modeled with shell elements (Kang et al., 1997). The FEHM wasvalidated against intracranial pressure data from Nahum et al.(1977),Trosseille et al. (1992), Deck et al., (2004), Willinger andBaumgartner, 2003 and Deck and Willinger (2008a,b). The brainmodel was enhanced by implementing anisotropy and fiber data(fractional anisotropy and fiber orientation) from medical imaging(diffuse tensor imaging) into new constitutive law (Chatelin et al.,2013) and recently validated against local brain motion data fromHardy et al. (2001, 2007) by Sahoo et al. (2013a). The skull modelwas also improved by using appropriate composite material modelby taking into account the fracture by Sahoo et al. (2013b). The newcomposite skull model was validated for maximum forces andlateral impact against actual force-time curves from PMHS in theentire time domain at different velocities and for differentboundary conditions. A detailed presentation of different partsof the FE head model is shown in Fig. 2.

The FEHM was used to reproduce these lateral impact experi-ments by conducting numerical simulations and investigating theinfluence of anthropometry on the biomechanical responses. The

(Zhang et al., 2009; Sahoo et al., 2013b).

Fig. 2. Detailed human head FE model (Sahoo et al., 2013a,b,b).

D. Shaoo et al. / Accident Analysis and Prevention 80 (2015) 97–105 99

impact surface was modeled as brick element with MAT63CRUSHABLE_FOAM of thickness 50 mm and rested on the topof a rigid platform. The three impactors (flat and cylindrical 90Dand flat 40D padding) are validated for different velocity by Sahooet al. (2013b). The rigid platform was constrained at the bottom tomimicthe boundary condition. The CONTACT_AUTOMATIC_SUR-FACE_TO_SURFACE interface was used between FE head model andimpactor with a static friction coefficient of 0.7. The mid sagittalplane of the head FE model was aligned at an angle of 10 degreeswith respect to horizontal plane as described in the experimentsuch that the impact occurred to the left temporo-parietal region. Aschematic for the simulation configuration in LS-DYNA isillustrated in Fig. 3. The velocity just before the impact inexperiment was applied to the FE head model in the initial velocitycard. During the reconstruction of experimental cases,variationsbetween the FEM predicted and experimental corridors werequantified by calculating the percentage of difference betweensimulation and experimental peak contact forces and alsocalculating the correlation value “r” (also known as samplePearson correlation coefficient). Hence the FEHM was validatednot only for maximum force but also in the entire time domain(Sahoo et al., 2013b). This validated FEHM was used to study effectof stiffness and shape of different contact surfaces on variousbiomechanical metrics for different population human heads. Thecurrent model was based on a 50th % adult human head. To obtainthe 5th and 95th % FE head model, the mass was altered bydecreasingor increasing the density of different parts of the 50th %FEHM proportionally. A uniform scaling of densities of differentparts of the FEHM was performed to match the mass of differentpopulation human heads. It was assumed that the mass of the headhas more influence than the size of the head. The current approachof altering the properties uniformly from the 50th to the 5th and95th heads was also used by Ruan and Prasad (2001) in their finiteelement modeling study. Numerical simulations of lateral head

Fig. 3. Illustration of SUFEHM model orientation and boundary conditions used forsimulation under LS-DYNA (Sahoo et al., 2013b).

impacts for 45 cases (15 experiments � 3 different populationhuman heads) were performed using the LS-DYNA platform atvelocities ranging from 2.4 to 6.5 m/s and three impactingconditions (flat and cylindrical 90D; and flat 40D padding).Resultant contact force-time histories for each simulation werecompared with experimental data. The extracted results fromsimulations were quantified by the percentage of differencebetween simulation and experimental peak contact forces in thetime domain. Strain energy densities were also extracted for eachsimulated case.

3. Results

A total of 45 simulations (15 experiments � 3 different popula-tion human heads) were conducted. The entire force-time signalsfrom the simulations were compared with experimental mean andupper/lower corridors at each velocity, stiffness of the impactingsurfaces (40 and 90 Durometer) and shapes (flat and cylindrical).The experimental corridors shown in Figs. 4–6 represent plus-minus one standard deviation curves from the mean. They arelabeled as UC and LC, upper and lower corridors. They wereobtained by calculating the mean and standard deviations of eachdata point from each specimen along the entire time history. Theoutputs from the model were also filtered at SAE 1000 Hz as per theexperiments. Fig. 4 shows the comparison of contact forcesobtained from different population head impacts with experi-mental corridors for six different velocities for 40D flat impactor.Eighteen simulations were conducted for six different velocities.The velocity ranges from 6.47 m/s to 3.46 m/s are in accordancewith the experimental data. The peak forces at different velocitiesare listed in Table 1. The average deviations of the peak force frommean experimental data to the 95th- male, 50th- male and 5th-female were 5%, 4% and 15.8%. Average deviation of peak force fromthe 50th male to 95th male and 5th female were 6% and 13.3%. Theskull strain energies for all the cases were calculated. It is observedthat the skull strain energy increased by an average of 7.6% for 95th% male and decreased by an average of 17.7% for 5th % female from50th % male skull strain energies.

Fig. 5 shows the comparison of contact forces obtained fromdifferent population head impacts with experimental corridors forfive different velocities for 90D flat impactor. Fifteen simulationswere conducted for five different velocities. The velocity rangesfrom 5.46 m/s to 2.44 m/s are in accordance with the experimentaldata. The peak forces at different velocities are also listed in Table 1.The average deviations of the peak force from mean experimentaldata to the 95th- male, 50th- male and 5th- female were 5%, 6.4%and 15.5%. Average deviation of peak force from the 50th male to95th male and 5th female were 3% and 10%. The skull strainenergies for all the cases were calculated. It is observed that theskull strain energy increased by an average of 5% for 95th % male

100 D. Shaoo et al. / Accident Analysis and Prevention 80 (2015) 97–105

and decreased by an average of 13.3% for 5th % female from 50th %male skull strain energies.

Fig. 6 shows the comparison of contact forces obtained fromdifferent population head impacts with experimental corridors forfour different velocities for 90D cylindrical impactor. A total oftwelve simulations were conducted for four different velocities.The velocity ranges from 4.89 m/s to 2.44 m/s are in accordancewith the experimental data. The peak forces at different velocitiesare listed in Table 1. The average deviations of the peak force frommean experimental data to the 95th- male, 50th- male and 5th-

-2 0 2 4 6 8 10 12 14

0

2000

4000

6000

8000

10000

95th % Male50th % Male5th % FemaleUCLCMean

Forc

e (N

)

Time (ms)

V=6.4 7m/s

-2 0 2 4 6 8 10 12 14

0

2000

4000

6000

8000

10000

95th % Male50th % Male5th % FemaleUCLCMean

Forc

e (N

)

Time (ms)

V=5.4 6m/s

-2 0 2 4 6 8 10 12 14

0

2000

4000

6000

8000

10000

95th % Male50th % Male5th % FemaleUCLCMean

Forc

e (N

)

Time (ms)

V=4.2 4m/s

Fig. 4. Comparison of contact forces obtained from different population head impa

female were 2%, 7.5% and 13.6%. Average deviation of peak forcefrom the 50th male to 95th male and 5th female were 10.3% and8.5%. The skull strain energies for all the cases were calculated. It isobserved that the skull strain energy increased by an average of10.2% for 95th % male and decreased by an average of 6.7% for 5th %female from 50th % male skull strain energies.

The skull fracture patterns for 45 simulations with differentpopulation FEHMs were obtained by marking the fractureinitiation and accounting for all the element failures until theend of the simulation. The fracture pattern for 95th male, 50th

-2 0 2 4 6 8 10 12 14

0

2000

4000

6000

8000

10000

95th % Male50th % Male5th % FemaleUCLCMean

Forc

e (N

)

Time (ms)

V=5.99m/s

-2 0 2 4 6 8 10 12 14

0

2000

4000

6000

8000

10000

95th % Male50th % Male5th % FemaleUCLCMean

Forc

e (N

)

Time (ms)

V=4.89 m/s

-2 0 2 4 6 8 10 12 14

0

2000

4000

6000

8000

10000

95th % Male50th % Male5th % FemaleUCLCMean

Forc

e (N

)

Time (ms)

V=3.46m/s

cts with experimental corridors for 6 different velocities for 40D flat impactor.

-2 0 2 4 6 8 10 12

0

2000

4000

6000

8000

10000

12000

14000

95th % Male50th % Male5th % FemaleUCLCMean

Forc

e (N

)

Time (ms)

V=5.4 6m/s

-2 0 2 4 6 8 10 12

0

2000

4000

6000

8000

10000

12000

95th % Male50th % Male5th % FemaleUCLCMean

Forc

e (N

)

Time (ms)

V=4.89 m/s

-2 0 2 4 6 8 10 12

0

2000

4000

6000

8000

10000

95th % Male50th % Male5th % FemaleUCLCMean

Forc

e (N

)

Time (ms)

V=4.2 4m/s

0 2 4 6 8 10 12

0

2000

4000

6000

8000

10000

95th % Male50th % Male5th % FemaleUCLCMean

Forc

e (N

)

Time (ms)

V=3.46m/s

-2 0 2 4 6 8 10 12

0

2000

4000

6000

8000

10000

95th % Male50th % Male5th % FemaleUCLCMean

Forc

e (N

)

Time (ms)

V=2.4 4m/s

Fig. 5. Comparison of contact forces obtained from different population head impacts with experimental corridors for 5 different velocities for 90D flat impactor.

D. Shaoo et al. / Accident Analysis and Prevention 80 (2015) 97–105 101

male and 5th female FEHMs for impact with 90 D flat impactorwith impact velocity of 5.46 m/s are illustrated in Fig. 7. Thefracture elements are represented by the black colour. For 95th and50th male similar fractureswere observed, where as in case of 5thfemale FEHM, few elements (accounted for element failure andrepresented in black colour) less than the former were observed.However the strain energies for 95th and 50th male are different,which are 1254 mJ and 1193 mJ respectively.

4. Discussion

The biomechanical response of human skull in lateral impact isdifferent from the frontal impact; the applicability of frontalimpact criterion for side impact of skull is not fully justified. Veryfew parametric studies on lateral impact are reported in literature(Zhang et al., 2009). Ruan and Prasad, 2001 and Kleiven and vonHolst (2002) conducted frontal impact simulation for their

-2 0 2 4 6 8 10 12

0

2000

4000

6000

8000

10000

95th % Male50th % Male5th % FemaleUCLCMean

Forc

e (N

)

Time (ms)

V=4.8 9 m/s

-2 0 2 4 6 8 10 12

0

2000

4000

6000

8000

10000

95th % Male50th % Male5th % FemaleUCLCMean

Forc

e (N

)

Time (ms)

V=4.24m/s

-2 0 2 4 6 8 10 12

0

2000

4000

6000

8000

10000

95th % Male50th % Male5th % FemaleUCLCMean

Forc

e (N

)

Time (ms)

V=3.4 6m/s

-2 0 2 4 6 8 10 12

0

2000

4000

6000

8000

10000

95th % Male50th % Male5th % FemaleUCLCMeanFo

rce

(N)

Time (ms)

V=2.44m/s

Fig. 6. Comparison of contact forces obtained from different population head impacts with experimental corridors for 4 different velocities for 90D cylindrical impactor.

Table 1Peak forces at different impact velocities for the 3 impacting surfaces.

Peak forces (N) for 40D flat impactor

V = 6.47 m/s V = 5.99 m/s V = 5.46 m/s V = 4.89 m/s V = 4.24 m/s V = 3.46 m/s

Experiment 8695 (7420–9970) 8258 (7150–9420) 7635 (6810–8460) 6630 (6010–7250) 5650 (5000–6300) 3985 (3330-4640)Simulation Reference(Sahoo et al., 2013b) 8878 8140 7094 6117 5279 41005th % Female 7669 6974 6155 5276 4591 362250th % Male 8721 8104 7174 6213 5268 412095th % Male 9250 8578 7700 6652 5339 4461

Peak forces (N) for 90D flat impactor

V = 5.46 m/s V = 4.89 m/s V = 4.24 m/s V = 3.46 m/s V = 2.44 m/s

Experiment 9765 (6830–12700) 9215 (7230–11200) 8430 (6760–10100) 6890 (5670–8110) 4545 (3930–5160)Simulation Reference (Sahoo et al., 2013b) 9820 8748 7557 6241 42405th % Female 8703 7832 6825 5541 394150th % Male 9858 8661 7574 6251 428195th % Male 10177 8938 7742 6329 4515

Peak forces (N) for 90D cylindrical impactor

V = 4.89 m/s V = 4.24 m/s V = 3.46 m/s V = 2.44 m/s

Experiment 7280 (5060–9500) 7110 (6530–7690) 6315 (5240–7390) 4050 (3620–4480)Simulation Reference (Sahoo et al., 2013b) 7158 6836 5806 38665th % Female 6286 6502 5146 348250th % Male 7234 6229 5829 366595th % Male 7295 7094 6616 4145

(min–max) = corridors of experiment.

102 D. Shaoo et al. / Accident Analysis and Prevention 80 (2015) 97–105

Fig. 7. Fracture pattern for different population FEHMs for 90D flat impactor at impact velocity of 5.46 m/s. Note the differences in the locations and number of shaded areas inthe three heads.

D. Shaoo et al. / Accident Analysis and Prevention 80 (2015) 97–105 103

parametric studies. Kleiven and von Holst (2002) used FEHMs withvaried size and mesh density for replicating Nahum et al. (1977)frontal impact experiments. The scaling method used based onMerz (1984) may not be adequate for side impacts (Yoganandanet al., 2014). The FEHMs used in the above studies were notvalidated for actual force-time plots coming from lateral impactstest data with PMHS. More over both the studies used singleimpacting surface without considering the stiffness and shapevariation of impactor. In the current study, responses of differentpopulation (95th- male, 50th- male and 5th- female) human headsduring lateral impact at different velocities were investigatedalong with the influence of stiffness and shape of the impactingsurfaces to the peak contact force. It was observed from Figs. 4–6that the 95th % male results highest peak force and 5th % femaleresults least peak forces. However, the response variation issmaller between the 50th and 95th male head anthropometry thanbetween the 50th male and 5th female head anthropometry. This isbecause the mass difference between the two male heads(4.94 � 4.54 = 0.4 kg) is approximately twice lower than the massdifference between the male and female (4.54 � 3.73 = 0.81 kg).The lowest mass of the 5th female head responds with lower forcesfor the same impact level, indicating the response dependence ofbiomechanical metrics to mass variations. This indicates higherhead mass results in greater peak forces than the lower head mass,an expected outcome. Similarly the lower stiffness impactor (lowdurometer) was responsible for decreased force and that was truefor all velocities. However, for the same durometer condition, peakforces were lower with the cylindrical than the flat impactingsurface and this may be attributable to the reduced contact areaassociated with the former boundary condition. The percentagechanges in the peak force from the mean experimental datafollowed similar patterns for all three impactors and threedifferent populations. However when calculating the percentagedeviation of peak force from 50th male response, it was observedthat lower durometer accounted for greater changes than thehigher durometer. The effect of shape of the contacting surface isalso more pronounced when compared with 50% male response(3–10% for 95th % male -5th % female for 90D flat than 10.3–8.5%for 95th % male -5th % female for 90D cylindrical surfaces). The %change of peak force of 5th % female with respect to 50th % maledecreases with increase in velocity. The trend is the opposite

Table 2Comparison of Peak contact forces and skull strain energies for different population humfemale and 95th male head is similar to 50th% and density of different parts is altered; andaltered.

Sa

Mass (Kg) Length (mm) Breadth (mm) Fo

5th % Female 3.73 196 155 8750th % Male 4.54 195 154 9895th % Male 4.94 183 142 10

considering the % change of peak force of 95th % male with respectto 50th % male. However these findings suggest that the peak forcemagnitude is more dependent on velocity and stiffness than theshape of the impacting surface.

One of the shortcomings of this study was not to include factorssuch as change in head size which may be associated with alteredhead mass. The uniform scaling of mass of all head componentsmay be different than the real mass distribution in smaller andlarger heads. As it was assumed that head mass has more influencethan head size, three simulations were conducted by changing thesize of the 50th % male at impact velocity of 5.46 m/s with 90D flatimpactor. The head size of 50th % male FEHM was scaled up anddown to match with the 95th male and 5th female FEHMs. Thepeak contact force and the skull strain energies were recorded forall the population human heads and compared with the previousobtained results as listed in Table 2. It was observed that deviationsin peak contact force and skull strain energy are less than 2%,demonstrating the higher influence of head mass than size.

Also, it should be noted that the physiological process ofosteoporosis affects the cancellous/trabecular bone more than thecortical bone. The effect is seen more in lumbar spine and hip areasfrom where the bone density is measured clinically, using DEXA orQCT techniques. In the human spine, the trabecular bone occupiesmore than 80% of the volume of the vertebral body. The latticestructure of the cancellous bone deteriorates with age leading toosteoporosis in some individuals. The cortical shell does not havesuch consequences due to its different structure. Because thehuman skull has relatively thicker cortical bones, and bears lowload over the life span compared to the lumbar spine (from thetorso), effects of this disease process is low, if any. Thus, age is notconsidered as a dominant variable in skull fracture forcepredictions (Bass and Yoganandan, 2015).

The skull strain energies for 95th% male and 5th% female werecompared with the 50th % male by calculating the % difference.From all the cases it was observed that with increase in mass,velocity and stiffness of the impactor the skull strain energyincreases. For higher mass (95th % male) the discrepancy from 50th% male is lower (7.6 � 5% for 40D and 90D flat for 95th % male) than5th % female head (17.7 � 13.3% for 40D and 90D flat for 5th % male)for similar shape impactor. However the pattern is opposite whencompared with different shape impactor with same stiffness. So

an head for 90D flat impactor at 5.46 m/s. Two configurations: when the size of 5th when the density of 5th female and 95th male head is similar to 50th% and sizes are

me size and density changed Same density but size changed

rce (N) Strain energy (mJ) Force (N) Strain energy (mJ)

03 1016.926 8675 101558 1192.822 9858 1192.822177 1254.04 10217 1293.6

Fig. 8. (A) Skull strain energy for each case (50th % male) along with strain energy variation for 95th male and 5th female (shown by black vertical lines on each column). Thewhite columns represent the cases with no skull fracture and the gray columns represent the cases with fracture. (B) The % risk of skull fracture for all cases. The columnsrepresent the results for 50th % adult male, the upper and the lower limits represent the results for 95th % adult male and 5th % female respectively. The white columns are forcases without fracture and gray columns are for with fracture. The red line represents the 50% risk of skull fracture (544 mJ) (Sahoo et al., 2013b). (For interpretation of thereferences to colour in this figure legend, the reader is referred to the web version of this article.)

104 D. Shaoo et al. / Accident Analysis and Prevention 80 (2015) 97–105

the shape of impacting surface has reasonable influence on skullstrain energy variation. The thickness of skull is another variablewhich may influence the magnitude of this secondary variable.Thevariations of skull strain energies for 95th % male and 5th % femalefrom 50th % male are calculated. These variations are thensuperimposed onto the strain energy data for the 50th % malealong with the 50% threshold limit for skull fracture as shown inFig. 8(A). All the simulations are labeled as Case A, Case B and CaseC as in Sahoo et al. (2013b) and are representing SUFEHM impactsimulations on 40D flat, 90D flat and 90D cylindrical paddingrespectively. The columns represent the strain energies for the50th% male and the upper and lower black lines on each columnrepresent the strain energies for 95th% male and 5th % femalerespectively. The white columns represent the cases with no skull

Fig. 9. S curves obtained after binary logistical regression analysis for the skull strain

simulations. The tabulated data are the Nagelkerke’s R2 values and skull strain energie

fracture and the gray columns represent the cases with fracture. Inall the cases the criterion for 50% risk of skull fracture is notaffected by the variation in population. The % risks of skull fracturefor all cases are calculated and difference of % risk from 50th % maleresults to the other two populations are superimposed in the samefashion as shown in Fig. 8(B). The % risk is more for 95th % malethan the 5th % female when compared for same stiffness impactorbut different shape (90D flat and cylindrical). Hence for strainenergy, shape has reasonable influence than for peak force which ismore influenced by the mass of the different population head andstiffness of different impactors.

Statistical analysis (based on binary logistical regression) for theresults obtained during the simulation with different populationhuman heads was performed. Separate S curves were plotted for

energy data for 5th-percentile female, and 50th- and 95th- percentile male FEHMs for 50% risk of skull fracture.

D. Shaoo et al. / Accident Analysis and Prevention 80 (2015) 97–105 105

different population human heads as shown in Fig. 9. TheNagelkerke’s R2 values for the 5th-percentile female and 50th-and 95th- percentile male human heads are similar (0.68 for 5th %female, 0.696 for 50th % male and 0.685 for 95th % male). From theS curves the 50% tolerance of skull fracture was obtained fordifferent population heads. The 50% risks of skull fracture for 5th %female, 50th and 95th male were 524, 544 and 572 mJ of skullstrain energy respectively. These limits are within +/� 5% ofdeviation from the 50% risk of 50th % head skull fracture.Regardlessof increase/decrease of skull strain energy influenced by popula-tion variation, the average 50% fracture tolerance limit (546 mJ)was not altered much from the referenced limit, which was 544 mJ.This can be used as an initial threshold in the analysis of skullfracture for the conditions used in the present study.

The present parametric study provides information for a betterunderstanding of mechanical responses of the head during lateralimpact by the application of a 3D human head FE model byincorporating impacting surface features and it should beconsidered as a first step in a full analysis of the biomechanicsof head injuries. However, the skull anatomy used in the model wassomewhat simplified. Consequently, in order to more comprehen-sively analyze skull fracture mechanics, it is important to includeall the anatomical features of the human skull and this may includethe suture depending on age. This is considered as a future researchstudy.

5. Conclusions

The responses of 5th-percentile female, and 50th- and 95th-percentile male human heads during lateral impacts at differentvelocities and 3 different stiffness and shape of the impactingsurface were investigated. There is a greater influence of head masson the development of impact force. Impact velocity and padstiffness had a larger influence on impact force than the geometryof the impact surface. These results indicate hierarchy of variableswhich can be used in injury mitigation efforts. These findings offerinsights into contact loading events, identified as a major sourcefor causing head injuries in real-world side impacts in US andEurope, and provide external (force) and other metrics (skull strainenergy etc.) for improved understanding of head injuries in sideimpacts for all sized adult human head populations.

Acknowledgements

The authors acknowledge the ANR-12-EMMA-0026-0(SUFEHM-13) and VA Medical Research for their research supportto this work.

References

Allsop, D.L., Perl, T.R., Warner, C.Y., 1991. Force/deflection and FractureCharacteristics of the Temporoparietal Region of the Human Head. Society ofAutomotive Engineering Paper No. 912907.

Bass, C., Yoganandan, N., 2015. Skull and facial injury biomechanics. In: Yoganandan,N., Nahum, A.M., Melvin, J.W. (Eds.), Accidental Injury: Biomechanics andPrevention. Springer, New York, pp. 203–220.

Chatelin, S., Deck, C., Willinger, R., 2013. An anisotropic viscous hyperelasticconstitutive law for brain material finite element modeling. J. Biorheol. 27 (1–2),26–37.

Deck, C., Nicolle, S., Willinger, R., 2004. Human head FE modelling: improvement ofskull geometry and brain constitutive laws. Proceedings of the IRCOBIConference, Graz, pp. 79–92.

Deck, C., Willinger, R., 2008a. Improved head injury criteria based on head FE model.Int. J. Crashworthiness 13 (6), 667–678.

Deck, C., Willinger, R., 2008b. Head injury prediction tool for predictive systemsoptimization. Proceedings of the 7th European LS-DYNA Conference .

Faul, M., Xu, L., Wald, M.M., Coronado, V.G., 2010. Traumatic Brain Injury in theUnited States: Emergency Department Visits, Hospitalizations and Deaths2002–2006. Centers for Disease Control and Prevention, National Center forInjury Prevention and Control, Atlanta, GA.

Gennarelli, T., Pintar, F., Yoganandan, N., Beuse, N., Morgan, R., 2002. Head injuries tonearside occupants in lateral impacts: epidemiological and full-scale crash testanalyses. Proceedings of the IRCOBI, Munich, Germany.

Hardy, W.N., Foster, C.D., Mason, M.J., Yang, K.H., King, A.I., Tashman, S., 2001.Investigation of head injury mechanisms using neutral density technology andhigh-speed biplanar X-ray. Stapp Car Crash J. 45, 337–368.

Hardy, W.N., Mason, M.J., Foster, C.D., Shah, C.S., Kopacz, J.M., Yang, K.H., King, A.I.,Bishop, J., Bey, M., Anderst, W., Tashman, S., 2007. A study of the response of theHuman cadaver head to impact. Stapp Car Crash J. 51, 17–80.

Kang, H.S., Willinger, R., Diaw, B.M., Chinn, B., 1997. Validation of a 3D human headmodel and replication of head impact in motorcycle accident by finite elementmodeling. Proceedings of 41thStapp Car Crash Conference, Society ofAutomotive Engineers, Lake Buena Vista, USA.

Kleiven, S., von Holst, H., 2002. Consequences of head size following trauma to thehuman head. J. Biomech. 35 (2), 153–160.

Merz, H.J., 1984. A procedure for normalizing impact responsedata, SAE Paper No.840884. Soc. Automot. Eng. 4351–4358.

Nahum, A., Smith, R., Ward, C., 1977. Intracranial pressure dynamics during headimpact. Proceedings of 21st StappCar Crash Conference. SAE Paper No. 770922 .

ONISR Report, France November 2012. Available at: http://www.securite-routiere.gouv.fr/la-securite-routiere/l-observatoire-national-interministeriel-de-la-securite-routiere

Pintar, F.A., Yoganandan, N., Baisden, J., 2005. Characterizing occipital condyle loadsunder high-speed head rotation. Stapp Car Crash J. 49, 33–47.

Pintar, F.A., Maiman, D.J., Yoganandan, N., 2007. Injury patterns in side pole crashes.Ann. Proc. Assoc. Adv. Automot. Med. 51, 419–433.

Raymond, D., Van Ee, C., Crawford, G., Bir, C., 2009. Tolerance of the skull to bluntballistic temporo-parietal impact. J. Biomech. 42 (15), 2479–2485.

Ruan, J., Prasad, P., 2001. The effects of skull thickness variations on human headdynamic impact responses. Stapp Car Crash J. 45, 395–414.

Sahoo, D., Deck, C., Willinger, R., 2013a. Development and validation of an advancedanisotropic visco-hyperelastic human brain FE model. J. Mech. Behav. Biomed.Mater. 33, 24–42. doi:http://dx.doi.org/10.1016/j.jmbbm.2013.08.022.

Sahoo, D., Deck, C., Yoganandan, N., Willinger, R., 2013b. Anisotropic compositehuman skull model and skull fracture validation against temporo-parietal skullfracture. J. Mech. Behav. Biomed. Mater. 28, 340–353.

Trosseille, X., Tarriere, C., Lavaste, F., Guillon, F., Domont, A., 1992. Development of aF.E.M. of the human head according to a specific test protocol. Proceedings of36thStapp Car Crash Conference. SAE Paper No. 922527 .

WHO report, Fact sheet No 358, March 2013. Available at: http://www.who.int/mediacentre/factsheets/fs358/en/

Willinger, R., Baumgartner, D., 2003. Human head tolerance limits to specific injurymechanisms. Int. J. Crashworthiness 8 (6), 605–617.

Yoganandan, N., Pintar, F.A., Sances Jr., A., Walsh, P.R., Ewing, C.L., Thomas, D.J.,Snyder, R.G., 1995. Biomechanics of skull fracture. J. Neurotrauma 12 (4), 659–668.

Yoganandan, N., Pintar, F.A., 2004. Biomechanics of temporo-parietal fracture. Clin.Biomech. 19, 225–239.

Yoganandan, N., Gennarelli, T.A., Zhang, J., Pintar, F.A., Takhounts, E., Ridella, S.A.,2009a. Association of contact loading in diffuse axonal injuries from motorvehicle crashes. J. Trauma 66, 309–315.

Yoganandan, N., Maiman, D.J., Guan, Y., Pintar, F., 2009b. Importance of physicalproperties of the human head on head-neck injury metrics. Traffic Inj. Prev. 10,488–496.

Yoganandan, N., Arun, M.W.J., Humm, J.R., Pintar, F., 2014. Evaluation ofnormalization approaches for developing temporal corridors in oblique sideimpacts. Proceedings of the IRCOBI, Berlin, Germany.

Zhang, J., Yoganandan, N., Pintar, F.A., Gennarelli, T.A., 2006. Brain strains in vehicleimpact tests. Ann. Proc. Assoc. Adv. Automot. Med. 50, 1–12.

Zhang, J., Yoganandan, N., Pintar, F.A., 2009. Dynamic biomechanics of the humanhead in lateral impacts. Ann. Proc. Assoc. Adv. Automot. Med. 53, 249–256.