# Development and validation of a bicycle helmet: Assessment of head injury risk under standard impact conditions

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3 Fig. 2. Test apparatus used to perform experimental impact tests: Monorail shock absorption test equipment (overall dimensions including base: mm. 860x800x5500), Magnesium alloy complete headform The resultant linear acceleration versus time curve was calculated according to equation (1) and filtered by a CFC filter with a 1 khz cutoff frequency. where t t t t x (1) t t is the resultant linear acceleration (g), x y z is the linear acceleration in x axes (g), t linear acceleration in y axes (g) and t is the linear acceleration in z axes (g) z is the A total of 15 locations were impacted on both flat anvil (points A to O) and kerbstone anvil (points 1 to 15) at velocities of 5.42 m/s and 4.57 m/s respectively. Impact sites on flat anvil are shown by an X as illustrated in Fig. 3a c. The 15 other points on kerbstone anvil are represented by a line drawn on the surface of the helmet (Fig. 3d f). To ensure an accurate impact localisation, a laser pointer placed inside the anvil holder was used for tracking target points. y Top view Side view (c) Rear view (d) Top view (e) Side view (f) Rear view Fig. 3. Illustration of the impact points on the helmet during the experimental helmet test campaign

4 Before the impacts were performed, each of the 45 helmets were conditioned under one of the three following environmental conditionings: ( 20 ± 2) C in a BINDER 240L freezer for 4h to 6h (+50 ± 2) C in BINDER M400 incubator for 4h to 6h Wet condition at ambient temperature: the helmet is placed under a shower for 4h to 6h To comply with the EN 1078 requirements, all locations around the helmet were impacted under cold, hot and wet conditions so that a total number of 90 normative impact tests were performed experimentally. Computational Modeling The helmet under study was digitalized to get a Computer Aided Design (CAD) of its geometry. The developed helmet model consists of two parts i.e. an external shell and a liner. The CAD was automatically meshed under the PAMCRASH Mesh module D trias elements were created for the shell and D tetras elements for the lining. Hence, the complete helmet FEM consists of elements. The whole model is presented in Fig. 4. Shell elements have a minimum and maximum size of 2.35 mm and 7.5 mm respectively. Concerning brick elements, the minimum element size is 2 mm. Finally, the meshing was generated to ensure a mesh continuity between the external shell and the liner. Fig. 4. General view of the bicycle helmet finite element model The mesh was imported under the LS DYNA explicit crash code and the material properties of the two helmet model s parts were implemented. It was observed that under impact, the compressed EPS foam liner produces large deformations. In order to simulate such deformations, the EPS behavior was modeled by a tabulated "MAT_CRUSHABLE_FOAM" material law. The material properties implemented in the model were based on EPS sample tests and summarized in table 1. The plastic behavior of the EPS was modeled by a unique stress versus strain curve according to environmental conditions and corresponding to a dynamic loading of the material (Fig. 5). It can be observed that at low temperature, the material is stiffer than at high temperature or under wet condition. The Poisson s coefficient of the foam was set close to zero due to its very high compressibility. The slope of the unloading curve was set to 28 MPa accordingly to Tinard et al. [16]. Table 1. Mechanical properties of the EPS foam ρ (kg/m 3 ) 61.6 E unloading (MPa) 28 ν ( )

5 Fig. 5. Tabulated curves implemented in the material law A linear elastic behavior was selected for the polycarbonate (PC) shell and modeled by the MAT_ELASTIC law. The values of the three parameters needed to characterize the PC material were extracted from the literature [17] and are reported in table 2. A measured thickness value of 1.18 mm was attributed to all shell elements. Table 2. Mechanical properties of the PC shell ρ (kg/m 3 ) 1055 E (MPa) ν ( ) 0.42 To reproduce the 90 normative experimental tests, an ISO headform FEM was imported under the LS DYNA environment. This 5.7 kg headform was meshed with 3280 quadrilateral 4 nodes shell elements as shown in Fig. 6. The rigid behavior of its constitutive material was defined by the MAT_RIGID law with Poisson s coefficient and Young s modulus set to 0.2 and 200 GPa respectively. The head helmet contact of the coupled system illustrated in Fig. 6 was defined by taking into account a CONTACT_AUTOMATIC_SURFACE_TO_SURFACE interface with a 0.2 friction coefficient. Fig. 6. Illustration of the headform model coupled to the bicycle helmet model Two anvil models, flat (306 brick elements) and kerbstone (2820 brick elements), were created to perform the numerical impacts on both surfaces according to the experimental approach, as it can be observed in Fig. 7. The geometry and the dimensions of the flat anvil comply with the EN 1078 prescriptions (minimum diameter of (125 ± 3) mm and 24 mm thick), as do the kerbstone anvil (angle of 90 ± 0.5, 50 mm minimum height and 200 mm minimum length). The MAT_ELASTIC material law and the parameter values as mentioned in table 3 are the same for both impacted surfaces. A "CONTACT_AUTOMATIC_NODE_TO_SURFACE" interface with a friction coefficient of 0.1 was defined to compute the shocks on the two anvils. Table 3. Mechanical properties of the flat/kerbstone anvil ρ (kg/m 3 ) 7800 E (MPa) ν ( )

6 Fig. 7. Illustration of the simulation of standard impact tests against two anvils: Flat anvil, Kerbstone anvil In order to assess the risk of head trauma in terms of diffuse axonal injury (DAI) and subdural hematoma (SDH) under normative impacts, a coupling between the bicycle helmet FEM and the Strasbourg University Finite Element Model (SUFEHM) was performed under the LS DYNA software. This model developed by Kang et al. in 1998 [18] and illustrated in Fig. 8 was meshed with brick elements and 2813 shell elements. Based on anatomical atlas, the head model contains the face, scalp, skull and a homogenous brain. Intra cranial membranes and cerebrospinal fluid (CSF) are also represented in this model. The constitutive material laws of the different parts of the model have been defined by the material laws specified in appendix 1 according to Willinger et al. (1998) [19]. This model was validated under dynamic experimental tests reported by Nahum et al. (1977) [20], Trosseille et al. (1992) [21] and Yoganandan et al. (1994) [22]. This head model has been extensively used for the simulation of a number of well documented real world head trauma in order to correlate intra cranial mechanical parameters with the occurence of injury. This approach permitted us to define model based head injury criteria reported by Deck et al. in 2008 [23], as follows: 50% risk of moderate neurological injury (DAI1+): Brain Von Mises shearing stress of 28 kpa 50% risk of severe neurological injury (DAI2+): Brain Von Mises shearing stress of 53 kpa 50% risk of subdural hematoma: CSF strain energy of 4950 mj The interface between the energy absorbing foam and the human head model (Fig. 9a) was computed by an AUTOMATIC_NODE_TO_SURFACE LS DYNA contact with a friction coefficient set to 0.2. An example of positioning of the coupled head helmet system for a lateral impact on a flat anvil is shown in Fig. 9b. Fig. 8. Strasbourg University Finite Element Head Model

7 Fig. 9. Helmet SUFEHM coupling: Illustration of the head helmet interface, Positioning of the coupled system for a normative impact on flat anvil III. RESULTS During the validation process, each of the 30 helmet target points was impacted experimentally and numerically under three environmental conditionings ( 20 C, +50 C and humidity). The numerical resultant linear acceleration was computed at the centre of gravity of the headform in order to superimpose it with the experimental result. In the framework of this in depth helmet validation effort, results in terms of resultant linear acceleration versus time curve were superimposed with experimental results as well as the peak values. An illustration is given for a crown impact (point C) at 5.42 m/s on a flat anvil under low temperature, high temperature and under humidity treatment in Fig. 10. A good agreement is observed in terms of time duration (8 ms) and peak resultant linear acceleration demonstrating that the stress strain curve defined for each condition is well suited to the energy absorbing material in the considered loading case. Numerical peak g is 163 g, 166 g and 174 g under 20 C, +50 C and H20 condition respectively. Related experimental values are 169 g, 161 g and 165 g leading to 3.7%, 3.0% and 5.2% relative errors. (c) Fig. 10. Superposition of experimental and numerical headform acceleration under crown impact at 5.42 m/s against a flat anvil: Cold condition, Hot condition, (c) Wet condition

8 Time history resultant linear accelerations obtained for a frontal impact (point 1) on kerbstone anvil are shown in Fig. 11 under cold, hot and humid conditionings. By analyzing the time durations (12 ms) and peak values between numerical and experimental results, a good agreement is also observed. Numerical simulation of impacts tests led to peak values of 76 g at 20 C, 88 g at +50 C, 83 g under humidity condition. Compared to experimental values, relative errors are respectively 5.3%, 13.6% and 2.4%. (c) Fig. 11. Superposition of experimental and numerical headform acceleration under frontal impact at 4.57 m/s against a kerbstone anvil: Cold condition, Hot condition, (c) Wet condition Results for 38 impact configurations obtained by impacting the helmet on a flat anvil are shown in Fig. 12 for the experimental and numerical approach. The validity of all simulations has been checked by comparing both hourglass and total energy of the system. To do so, the hourglass energy that is a numerical energy reflecting calculation errors must be lower than 10% of the total energy. In terms of peak resultant linear acceleration, a good agreement is particularly seen for points A to F under each of the three conditionings. Nonetheless, the theoretical acceleration is overestimated for points G and H. This artefact can be explained by the extreme sensitivity of the model response to the positioning of the helmet towards the anvil. Another reason which can explain these differences in peak values is that rupture and high shear deformations of the EPS liner appear when impacting these critical points, phenomena which are not taken into account in the present material model. Missing points (I and L) are points for which numerical instabilities have been observed and for similar reasons

9 (c) Fig. 12. Comparison of experimental and computed maximum headform acceleration under impacts against the flat anvil: Cold condition, Hot condition, (C) Wet condition A comparison between experimental and numerical peak acceleration values for impacts on kerbstone anvil at 4.57 m/s velocity and under all conditionings is presented in Fig For the 35 impact configurations, results show a good agreement between experimental and numerical data. Missing points, respectively points 12, 14 and 15 under all environmental conditions, led to numerical instabilities and hourglass energies higher than 10% of the total energies probably due to rupture as pointed out previously. It should be mentioned that for all impacts, the peak resultant linear acceleration never exceeded the EN 1078 acceleration limit fixed at 250 g. Fig. 13. Comparison of experimental and computed maximum headform acceleration under impacts against the kerbstone anvil: Cold condition, Hot condition

10 Fig. 14. Comparison of experimental and computed maximum headform acceleration under impacts against the kerbstone anvil under wet condition When coupled to the human head model, intra cranial parameters have been computed in order to assess the injury risks in terms of diffuse axonal injury and subdural hematoma under normative impacts in accordance with the model based injury criteria presented in the Methods section. As for the numerical simulations with the headform model, the validity of each simulation has been verified by comparing both hourglass and total energy of the system. Still, hourglass energy must be lower than 10% of the total energy involved. The peaks Von Mises intra cerebral stress computed under impact on both flat and kerbstone anvils for all target points and under the three normative conditionings are shown in Fig. 15. Generally speaking, most of the impact points led to significant but acceptable brain shearing. For normative impacts on kerbstone anvil, Fig. 15b also shows that the Von Mises stress calculated for each impact is lower than 28 kpa leading to a percentage of moderate DAI risk less than 50% for all impact configurations. Injury assessment in terms of percentage of risk to sustain a moderate DAI for the helmeted head impacting a flat or a kerbstone anvil are presented respectively in appendix 2 and appendix 3. Results obtained in terms of internal strain energy within the CSF layer for impacts on flat and kerbstone anvils, both under the three environmental conditionings, are presented in Fig. 16. Risks for the helmeted user to suffer from a SDH in terms of percentage of risk after impacting a flat surface under normative impact conditions are mentioned in appendix 4. As demonstrated in Fig. 16b, impacts realized on kerbstone anvil led to injury risk far below 50%, with a maximum risk of SDH of 5% obtained by impacting the point 2 under cold conditioning as it appears in appendix 5. Fig. 15. Peak intra cerebral shearing stress computed with the helmet model coupled to the human head model for all impact points: Flat anvil, Kerbstone anvil

12 V. CONCLUSIONS In this study, a finite element model of a brand new bicycle helmet was developed, implemented and validated under the LS DYNA explicit crash code. The numerical simulation of 90 experimental normative impact tests, under three environmental conditionings and two anvils, was performed by coupling the helmet FEM to a rigid 5.7 kg ISO headform complying with the EN 960 standard. Results in terms of headform acceleration time history and peak g values between experiment and simulation are in good agreement for most of the impact points thus validating the helmet model. However, for impact points located laterally and close to the helmet edge, further improvement of the helmet model is needed. Once validated, this helmet FE model was coupled to the Strasbourg University FE Head Model in order to assess the head injury risks of both DAI and SDH injuries. Results show that normative impacts on flat anvil are more critical than impacts against kerbstone anvil. Results also show that the computed injury risk is acceptable for most of the impact points. This work is therefore a step towards both helmet optimisation against biomechanical head injury criteria and the consideration of model based head injury criteria into future helmet standards. VI. ACKNOWLEDGEMENT This work has been developed within the ANR PREDIT Project BICYTETE led by the French department of Transport. VII. REFERENCES [1] United Nations, Internet: [2] Thompson R.S., Rivara F.P., Thompson D.C., A case control study of the effectiveness of bicycle safety helmets. N. Engl. J. Med., 320: , 1989 [3] Thomas S., Acton C., Nixon J., Battistutta D., Pitt W.R., Clark R., Effectiveness of the bicycle helmets in preventing the head injury in children: case control study. Brit. Med. J., 308: , 1994 [4] McIntosh A.S., Kallieris D., Mattern R., Svensson N.L., Dowdell B. An evaluation of pedal cycle helmets performance requirements. Proc, 39th Stapp Car Crash Conference, San Diego, USA, pp , 1995 [5] Ching R.P., Thompson D.C., Thompson R.S., Thomas D.J., Chillcott W.C., Rivara F.P., Damage to bicycle helmets involved with crashes. Accid. Anal. Prev., 29: , 1997 [6] NIS (Statistics Belgium), Mobiliteit , Brussels, 2000 [7] Collins B.A., Langley J.D., Marshall S.W., Injuries to pedal cyclists resulting in death and hospitalisation. N Z Med J, 106(969): , 1993 [8] Eilert Petersson E., Schelp L., An epidemiological study of bicycle related injuries, Accid. Anal. Prev., 29(3): , 1997 [9] Oström M., Björnstig U., Näslund K., Eriksson A., Pedal cycling fatalities in Northern Sweden. Int J Epidemiol., 22: , 1993 [10] NF EN 1078 standard, Helmets for pedal cyclists and for users of skateboards and roller skates, S , 1997 [11] EN 960 standard, Headforms for use in the testing of protective helmets, 1995 [12] Willinger R., Diaw B.M., Baumgartner D., Development and validation of a bicycle helmet finite element model, Proceedings of IRCOBI Conference, Göteborg, Sweden, pp , 1998 [13] Gilchrist A, Mills N.J., Finite element analysis of bicycle helmet oblique impacts, Int. J. Impact Engng 35, ,

13 [14] Asiminei A.G., Van der Perre G., Verpoest I., Goffin J., A transient finite element study reveals the importance of the bicycle helmet material properties on head protection during an impact, Proceedings of IRCOBI Conference, York, United Kingdom, pp , 2009 [15] ISO 6487 standard, Road vehicles Measurement techniques in impact tests Instrumentation, 2002 [16] Tinard V., Deck C., Willinger R., Modelling and validation of motorcyclist helmet with composite shell, Int J. Crash, DOI: / , 2012 [17] Deck C., Willinger R., Multi directional optimisation against biomechanical criteria of a head helmet coupling, Int J Crash, 11(6):561 72, 2006 [18] Kang H.S., Willinger R., Diaw B.M.,Chinn B., Validation of a 3D human head model and replication of head impact in motorcycle accident by finite element modeling. Proc. 41 th Stapp Car Crash Conference, Lake Buena Vista, USA, pp , 1997 [19] Willinger R., Diaw B.M., Kang H.S., Finite element modeling of the human head Modal and temporal validation, Int Conf on Vib Eng, Dalian, China, 1998 [20] Nahum A.M., Smith R., Warg C.C., Intracranial pressure dynamics during head impact, Proc. 21 th Stapp Car Crash Conference, New Orleans, USA, pp , 1977 [21] Trosseille X., Tarriére C., Lavaste F., Guillon F., Development of a FEM of the human head according to a specific test protocol, Stapp Car Crash Conf, Seattle, USA, 1992 [22] Yoganandan N., Pintar F.A., Sances A., Walsh P.R., Ewing C.L., Thomas D.T., Snyder R.G., Biomechanics of skull fracture, Journal of Neurotrauma, Vol.12, No. 4/ , 1995 [23] Deck C., Willinger R., Improved head injury criteria based on head FE model, Int J Crash, Vol. 13, No. 6, pp , 2008 [24] Otte D., Injury mechanism and crash kinematics of the cyclists in accident, Proc. 33 th Stapp Car Crash Conference, Washington DC, USA, pp. 1 20, 1989 [25] Larsen L.B., Epidemiology of bicyclist s injuries, Proceedings of the IRCOBI Conference, Berlin, Germany, pp ,

14 Segment Falx and Tent VIII. APPENDIX APPENDIX 1. MECHANICAL PROPERTIES OF THE SUFEHM CONSTITUTIVE PARTS Illustration Mechanical behavior Linear elastic Mechanical parameters e = 1 mm ρ = 1140 kg/m 3 E = 31.5 MPa υ = 0.45 Mechanical parameters / Brain Skull interface Linear elastic ρ = 1040 kg/m 3 E = MPa υ = 0.49 / ρ = 1040 kg/m3 Brain and Cerebellum Viscous elastic K = 1125 MPa G 0 = MPa G inf = / β = 145 s 1 Cortical Spongy e = 2 mm e = 3 mm ρ = 1900 kg/m 3 ρ = 1500 kg/m 3 Skull Fragile elasticplastic E = MPa υ = 0.21 E = 4600 MPa υ = 0.05 K = 6200 MPa K = 2300 MPa UTS = 90 MPa UTS = 35 MPa UTC = 145 MPa UTC = 28 MPa e = 10 mm Face Linear elastic ρ = 2500 kg/m 3 E = 5000 MPa / υ = 0.23 ρ = 1000 kg/m 3 Skin Linear elastic E = 16.7 MPa / υ =

15 APPENDIX 2. Risk of moderate DAI for all impact points on flat anvil and under three conditionings APPENDIX 3. Risk of DAI for 12 impact points on kerbstone anvil and under three conditionings APPENDIX 4. Risk of SDH for all impact points on flat anvil and under three conditionings APPENDIX 5. Risk of SDH for 12 impact points on kerbstone anvil and under three conditionings

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