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An experimental head restraint concept for primary prevention of head and neck injuries in frontal collisions
Accid. Anal. and Prev., Vol. 30, No. 4, pp. 535–543, 1998 © 1998 Elsevier Science Ltd. All rights reserved Printed in Great Britain 0001-4575/98 $19.00 + 0.00
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AN EXPERIMENTAL HEAD RESTRAINT CONCEPT FOR PRIMARY PREVENTION OF HEAD AND NECK INJURIES IN FRONTAL COLLISIONS P. H1*, H. H12 and I. E1 1Division of Lightweight Structures, Royal Institute of Technology, S-100 44 Stockholm, Sweden and 2Department of Clinical Neuroscience, Neurosurgery, Karolinska Institute, S-171 76, Stockholm, Sweden (Received 30 September 1996; in revised form 20 March 1997)
Abstract—The Experimental Head Restraint Concept ( EHRC ), a ‘safety belt’ for the head, is designed to reduce forces to the head and neck, in frontal car crashes. The EHRC was evaluated experimentally in frontal collision for a crash severity of 11 m/s, and numerically in frontal collision for a crash severity of 11 and 15 m/s. Experimental data obtained from a frontal barrier test (11 m/s) showed a 67% reduction of the HIC value from 411 (without EHRC ) to 136 (with EHRC ). The same level of reduction was also obtained for the higher speed in the numerical simulation. The moment in the neck was shown in experimental configuration to increase a few percent using the EHRC, but as presented in a numerical analysis, the moment was reduced by stiffening the EHRC. The EHRC clearly has a potential role in the search for primary prevention of neurotrauma injuries in frontal related car crashes. However, there is a strong need for more advanced injury criteria for the neck in order to optimize such complex safety systems. © 1998 Elsevier Science Ltd. All rights reserved Keywords—Head restraint, Injury prevention, Safety systems, Frontal collision
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system. As new car models are introduced, awareness of both passive and active prevention can be heightened. Traffic-related injury to the brain, like other causes, is of two categories: focal and diffuse injury. The focal brain injury is a lesion causing local damage that can be seen by the naked eye. The diffuse brain injury is associated with global disruption of brain tissue usually and is invisible. Both categories of injury are related to transfer of energy from an external impact to the head into the brain tissue. The translational and rotational impact causes a head acceleration often resulting in both focal and diffuse injury. By reducing the head acceleration due to impact, both categories of brain injury can theoretically be diminished significantly. Safety belts and airbags have considerably reduced the number of injuries from traffic accidents; however, both have their limitations. The safety belt reduces the motion of the body but leaves the head and neck unprotected. The airbag used today protects a front passenger in a frontal collision, but does not consider multidirectional forces or the rear seat passenger. There is a need to develop safety systems aimed toward protecting the head and neck and preventing neurotrauma during frontal related impacts. Thus, the aim of the present investigation
Neurotrauma, or injury to the central nervous system, is the most serious type of traumatic injury. Traumatic brain injury accounts for more than a third of all injury deaths (Rice et al., 1989; Kraus, 1991; Kraus, 1993; Kraus and McArthur, 1995; Kraus et al., 1995; Viano et al., 1997). After brain damage, many who survive live with impaired brain function. The toll from spinal cord injury is just as distressing, although the rate of occurrence is lower. The cost to society of neurotrauma is tremendous because of the high rate of incidence among the young (Rice et al., 1989, von Holst et al., 1995; Miller et al., 1995; Kraus et al., 1995). About 40% of all neurotrauma injuries occur as a result of traffic accidents ( Kraus, 1991; Gordon et al., 1995, von Holst et al., 1995; Sosin et al., 1995; Waxweiler et al., 1995; Viano et al., 1997). Current trends suggest an increase in the number of survivors of neurotrauma throughout the world (von Holst et al., 1995). High priority should be given to the long-term goals of prevention, treatment, and rehabilitation of people with damage to the central nervous *Corresponding author. Tel: +46 8 790 6448; Fax: +46 8 207865; e-mail: [email protected] 535
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was to evaluate the design of a new head protection concept, with the possibility of determining the retardation behavior of the head during the entire impact course, using both experimental and numerical frontal barrier tests. E XPE RIM EN TAL HE AD RE ST RAIN T C O N CE P T ( EH R C ) It should be stressed from the beginning that the EHRC was designed solely as an experimental tool for investigating ways to reduce trauma to both the head and neck in frontal associated crashes. As shown in Fig. 1, it consists of: (a) a curved aluminum pipe containing (b) a linear elastic spring; (c) a vertical aluminium plate, which replaces the neck rest and is bolted to the chair frame; and (d ) two endgrips that attach the spring to (e) a solid rod that is bolted to the vertical plate. The EHRC is adjustable vertically on the plate as shown in Fig. 2, and could be described
Fig. 1. Geometry of the EHRC (millimetres).
Fig. 2. Vertical adjustment of the EHRC.
as a ‘safety-belt’ for the head. The function of the EHRC is to decrease smoothly and continuously the G-forces that push the head forward and to keep the head from impacting surrounding objects in a collision. In comparison with the conventional airbag, the EHRC acts on the head as soon as the head leaves the neck rest. The spring retraction force keeps the head in position, and the length of time during which the head and neck are protected is therefore increased. The innovation was designed primarily to reduce the HIC (Head Injury Criteria) value, and secondarily to reduce forces to the neck by theoretically keeping the spine and head in a straight line during a frontal crash. In the experimental test, a spring constant of 1.5 kN/m was chosen to evaluate the effectiveness of the EHRC. E XPE RIM EN TAL SLE D T EST Method Two sled tests, one with the EHRC and one without, were carried out on the 50-metre crash court of VTI ( Va¨g och Transportforsknings Institutet, Linko¨ping, Sweden). The Hybrid III dummy developed in 1975 was chosen as the occupant substitute. The selected measured quantities were head resultant acceleration, moment in neck around occipital condyles, axial force in neck, shear force in neck, and resultant acceleration in thorax. The rotational acceleration was not measured as the Hybrid III dummy was without an angular accelerometer. The dummy was constructed and calibrated for frontal collision speeds of between 10 and 15 m/s; a speed of 11 m/s (40 km/h) was chosen for this study. It is our belief that the EHRC has better protection characteristics in the higher range of crash severity, and therefore, a lower speed was chosen. The dummy was placed in a standard Volvo 740 front seat equipped with a safety belt and the EHRC. The seat was bolted to a four-wheeled sled as shown in Fig. 3 and pulled by a cable to the selected crash speed. The retardation pulse was set to a level of about 18 gravitational units over a range of 90 milliseconds for both tests. Results The test results are shown as functions of time in Fig. 4. The peak translation acceleration of the center of the head was reduced by 32% [Fig. 4(a)]. To compare the acceleration of the head to the head injury criteria, the HIC values were calculated (Society of Automotive Engineers, 1984). With the EHRC, the HIC value was reduced by 67% (from 411 to 136). A small increase (13%) in the peak Y-moment of the neck was seen with the EHRC [Fig. 4(b)]. This result was expected because the
An experimental head restraint concept for primary prevention
Fig. 3. Experimental sled test.
EHRC acts at the top of the head, and a resulting moment is unavoidable. The peak axial and shear forces in the neck were reduced by 37 and 43% [Fig. 4(c,d)], respectively, with the EHRC. The peak acceleration of the thorax was reduced by less than 10% with the EHRC [Fig. 4(e)]. The results show that, compared to injury assessment reference values for the Hybrid III dummy (Mertz, 1984), the chosen collision speed results in responses that are below the respective threshold values for a midsize male. A higher crash severity will be analyzed in the Comparison of Results section. FI NIT E E LE ME NT ANA LYSIS Generally, for this type of analysis, implicit rigid body dynamic programs, such as MADYMO, are the most time-efficient tools. However, in order to evaluate an FE-package and an FE-model of the Hybrid III-dummy ( Fredriksson, 1994), we chose LS-DYNA3D as the numerical program. LS-DYNA3D is an explicit three-dimensional finite element code designed to analyze the large deformation dynamic response of elastic or rigid structures (Stillman and Hallquist, 1992b). The FE model consists of five components, as shown in Figs 5–7: Hybrid III dummy, three-point safety belt, EHRC, chair and floor, and airbag. The deformable model of the occupant substitute, developed by Fredriksson in the pre-processor LS-INGRID (Stillman and Hallquist, 1992a), consists of 17 body sections, which can be chosen as rigid or deformable. The model fulfils the criteria set by U.S. regulations for crash dummies, and gives
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results that are consistent with rigid body dynamic models such as MADYMO ( Fredriksson, 1994). The safety belt was modelled with shell elements and pin-joined truss elements. Triangular shell elements were used where the belt was assumed to contact the dummy; two truss elements, which can carry only tensile loads, were positioned where no contact was expected. Elastic-plastic material behaviour was assumed, with a Young modulus of 2,400 megapascals, a hardening modulus of 1,600 megapascals, and a yield stress of 69 megapascals ( Halldin, 1996). On the Volvo 740 seat, the upper section of the belt leads down from the retractor through a slipring to a pretensioner. Problems occurred while attempting to model this part of the belt because of the small radius of the slipring. The problem was solved by shortening the length of the belt and decreasing the elastic modulus of the material. LS-DYNA3D does contain an algorithm for defining the safety belt with slipring and pretensioner, but the characteristics of these parts are unknown. This work can be left to future researchers. The EHRC was modelled using a rigid 180-degree torus made up of eight-node brick elements; the torus was fastened to the seat with two undamped linear springs (Fig. 6). The chair and floor were modeled using Belytschko–Lin–Tsay shell elements made up of an elastic material. The airbag model was developed by Hallquist et al. (1990). COM PARISO N OF R ESU LT S Three major results are presented in this work: (a) FEM results for the EHRC and airbag for crash severities of 11 and 15 m/s; (b) the effect of spring stiffness on head and neck forces; and (c) comparison between experimental and numerical results. The measured quantities in the numerical studies were head resultant acceleration, moment in neck around occipital condyles, and resultant acceleration in thorax. FEM results for EHRD and airbag: dn=11 and 15 m/s Four different configurations of safety equipment were investigated with the Hybrid III dummy exposed for the crash severities of 11 and 15 m/s: (a) safety belt (SB) only; (b) SB and EHRC; (c) SB and airbag (AB); and (d ) SB, AB, and EHRC. Safety belts were used in all four configurations; comparison of the configurations thus allows evaluation of the effectiveness of the EHRC and airbag, both separately and in combination. The EHRC was equipped with a spring constant of 1500 N/m in the 11 m/s test and 3000 N/m in the 15 m/s test. FEM results for the four
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Fig. 4. Experimental results: With and without EHRC. Dn=11 m/s. (a) Resultant acceleration of head. (b) Moment in neck around occipital condyles, Y-axes. (c) Axial force in neck. (d ) Transverse shear force in neck. (e) Resultant acceleration in thorax.
configurations are shown for the head, neck and thorax in Figs 8 and 9. As expected, the safety belt alone gave the worst result, and the safety belt plus airbag with EHRC gave the best result. HIC values were calculated from the acceleration plots [Figs 8 and 9(a)] for each configuration and are shown in Fig. 10. The percentage improvement of the HIC values from one configuration to another is summarized in Table 1. Comparison of the safety belt with and without the EHRC shows an improvement of 37%. Comparison of the safety belt with and without the EHRC shows an improvement of 56%. Note that adding the EHRC to the safety belt plus airbag configuration gives a substantial additional improvement in the HIC value (from 37 to 60% and from 51% to 68%) but that adding the airbag to the safety belt plus EHRC configuration gives little or
no additional improvement (from 56 to 60% and none). Comparison of the results for the neck was much more difficult since no advanced criteria for the neck exist. It was impossible to analyze the neck moment curves in Figs 8(b) and 9(b) using the amplitude level criterion of Mertz and Patric (1971). However, it is clear that the characteristics of the moment acting on the neck differ considerably between configurations. The airbag and safety belt complemented by the EHRC appears to give the best response, whereas the EHRC and safety belt combination results in the highest moment forces. This negative effect of the EHRC was also shown in the experimental study. The resultant acceleration of the thorax is shown in Figs 8(c) and 9(c). The variations between the
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Fig. 5. FE model of the sled test: (a) EHRC with spring and brick element. (b) Safety belt with triangular shell and truss element. (c) Hybrid III dummy ( Fredriksson, 1994). (d) Chair and floor with shell element.
Fig. 6. FE model of the EHRC (side and top views).
Fig. 8. FEM results for four configurations of safety equipment. Dn=11 m/s.
To analyze a new concept, such as the EHRC, critically, and to optimize its properties, we will need better and more reliable injury criteria for the neck. For the head and neck, the FE analysis shows that the EHRC does give an improvement over the airbag; however, the combination of the two (with safety belt) clearly gives the best result. Fig. 7. FE model of the airbag.
different configurations are within 10%, and are explained by the low crash severity. A higher crash speed reveals the benefits of the airbag in reducing acceleration of the thorax [Fig. 9(c)].
Effect of EHRC spring stiffness on acceleration of the head and moment in the neck: Dn=11 m/s The effect of EHRC spring stiffness on acceleration of the head and the moment in the neck was studied numerically for the safety belt plus EHRC configuration using spring constants of 1,500, 3,000,
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P. H et al. Table 1. Percentage improvement of HIC values with different configurations of safety equipments Crash severity Safety equipment SB versus SB+AB SB versus SB+EHRC SB versus SB+AB+EHRC
11 m/s
15 m/s
37% 56% 60%
51% 68% 68%
and 3,750 N/m. The results are shown in Fig. 11. The acceleration of the head shows the same characteristics and approximately the same HIC value for all spring constants [Fig. 11(a)]; this is primarily the result of keeping the head from impacting the chest. The moment in neck around the occipital condyle shows a strong relationship between spring stiffness and moment [Fig. 11(b)]. A stiff spring removes most of the positive part in the curve while increasing the negative part by a few percent. The amplitude level criteria by Mertz and Patric (1971) state only that injury might occur below −57 Newton metres and above +190 Newton metres. Choosing the ‘best’ spring, therefore, is not a trivial matter as a small negative increase can be worse than a large positive decrease in the moment of the neck. Figure 11(b) clearly shows the need for advanced neck injury
Fig. 9. FEM results for four configurations of safety equipment. Dn=15 m/s.
Fig. 10. Comparison of different configurations of safety equipments for two levels of crash severity. HIC values (typed on the columns) are normalized against the configuration with safety belt only. SB, safety belt; AB, airbag.
Fig. 11. Effect of EHRC spring stiffness. Dn=11 m/s.
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criteria. We propose using a stiffness of 3000–3750 N/m for this crash severity (Dn=11 m/s). Comparison of experimental and numerical results For the safety belt plus EHRC configuration, a direct comparison of the experimental and numerical results was made for the videotape and computer animation, the acceleration plots for the head and thorax, and the moment plots for the neck. The comparisons are shown in Figs 12 and 13. The experimental videotape and computer animation (Fig. 12) show a reasonable correlation in the way in which the Hybrid III dummy moves. Some differences are evident, though, particularly toward the end of the test in which the dummy in the numerical simulation
Fig. 13. Graphical comparison of experimental (dotted line) and numerical (solid line with squares) results for safety belt with EHRC. Dn=11 m/s.
Fig. 12. Comparison of experimental videotape and computer animation for Hybrid III dummy equipped with safety belt and EHRC. Dn=11 m/s.
rebounded earlier. The reasons for this will be discussed below. In Fig. 13(a), a comparison of the results of the acceleration of the head shows an excellent correlation. However, significant differences are evident in comparing the results of the neck and thorax [Fig. 13(b,c)]. Both the moment in the neck and the resultant acceleration of the thorax show an earlier increase in the responses in the numerical simulation than in the experimental. The slope of the numerical curve in Fig. 13(b), however, shows the correct characteristics; therefore, the neck response can be used in comparison analysis. A comparison of results with and without the EHRC (for both experimental and numerical simulations), shows a decrease in thoracic acceleration of less than 10%, which emphasizes the fact that the responses of the thorax are of secondary
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interest in this study. Head and neck responses are of primary interest, and we believe that the present model performs well enough to be used for comparison analysis in this study, but not for a quantitative analyze aiming at predicting injury or no injury. A number of assumptions and simplifications may explain some of these differences. The FE model of the neck was calibrated to the standard Hybrid III pendulum test ( Fredriksson, 1994) and not to the sled test. The thorax was assumed to be rigid, which is one reason for the earlier increase in slope in Fig. 13(c). The experimental sled test and the FE simulation differed in three primary areas: (a) the safety belt, (b) the seat, and (c) the neck rest. The retractor, pretensioner, and slipring of the safety belt were not modeled. The belt dimensions, material properties, and contact definition were all correct. The seat, consisting of foam, a spring system, and bars, was modeled with an elastic plate. These modeling simplifications may explain the mismatch between experimental and numerical results for the acceleration of the thorax [Fig. 13(c)]. The neck rest, to which the EHRC was attached, was joined to the upper bar of the seat back. This bar was deformed in the experimental crash, and resulted in a forward bending of the neck rest ( Fig. 12). It was not possible to reproduce this bending in the FE model, which made the model stiffer in comparison to the test [Fig. 13(b)]. To obtain a better correlation, we suggest that simplifications be made to the experimental test configuration, to allow more accurate modeling of the chair and safety belt construction. D ISCUS SION This experimental and numerical study shows clearly that the Experimental Head Restraint Concept can be an alternative or a complement to conventional safety systems in frontal barrier tests. Two different crash severities were studied, 11 and 15 m/s, and show about the same improvement in HIC value (60 and 68%), using the EHRC plus safety belt compared to the safety belt alone. The stiffness of the EHRC used in the experimental tests, 1500 N/m, resulted in a small increase in the moment of the neck around the occipital condyles. Finite element analyses showed that the moment in the neck could be reduced by using a stiffer spring. The EHRC was designed to be used as an experimental tool. However, we must ask ourselves: could innovation be used as a complement to the conventional safety systems in a car of tomorrow, and what are the potential injuries that may result from the experimental concept? A future design of the EHRC can be either constantly active (as is the
present EHRC ) or activated at the time of impact (as an airbag). Both systems have their advantages and disadvantages. The constantly active system has the benefit of acting during the whole time of impact if it is positioned tightly next to the head. It can be designed to work in a wide range of collision angles, with zero retraction force in the normal range of motion (much like a conventional safety belt), but has a major disadvantage regarding user-friendliness. Probably, only a few will use a constantly active system voluntarily unless the car is equipped with an intelligent safety control system, which implies that the speed is reduced if the EHRC is not in use. A design that is activated at the time of impact is more attractive, with regard to user-friendliness, but a large amount of effort has to be focused on the injury potentials related to out-of-position occupants as the equipment is activated. The major risk attached to the EHRC is neck injury, which may be caused by the added neck moment around the occipital condyles. There is also risk of WAD injuries using the present design with the linear elastic spring producing a rebounding effect on the head and neck system. This risk, however, could be reduced or eliminated by using a non-linear elasto-plastic spring. It was not possible to test the hypothesis of maintaining a straight line between head and spine to prevent neck injuries, but this remains a goal for us to verify. Further investigations may include: a study of the spring material characteristics; analyses of multidirectional impacts as off-set collisions; and analysis of head rotational acceleration, either as a single physical quantity or calculated in more advanced head injury criteria. In addition, more advanced injury criteria are needed for the neck. A number of investigators (e.g. Yoganandan et al., 1989’ Bilston et al., 1993; Eppinger et al., 1994; Pintar et al., 1995) are currently studying both fracture and neurotrauma criteria for the neck, but much work remains to be done. Improved criteria will lead to improved interpretation of experimental and numerical results. Acknowledgements—We want to thank Lars Fredriksson at the University of Linko¨ping, and Niclas Bra¨nnberg at Engineering Research, for their support with the FE model. Thanks to Ingemar Gustavsson at Swedish Road and Transport Research Institute for his help with the experimental part. We also want to acknowledge Mrs Deborah K. Kerwin-Peck for editing the manuscript.
R EF ER EN C ES Bilston, L. E., Meaney, D. F. and Thibault, L. E. (1993) The Development of a Physical Model to Measure Strain in a Surrogate Spinal Cord During Hyperflexion and Hyperextension. Proceedings. IRCOBI, pp. 255–265. Eppinger, R. K., Kleinberger, M., Morgan, R., Khaew-
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pong, N., Bandak, F. and Haffner, M. (1994) Advanced Injury Criteria and Crash Evaluation Techniques. National Highway Traffic Safety Administration. Paper No. 94-S1-O-11. Fredriksson, L. A. (1994) Finite Element Modeling of Occupant Substitutes. Department of Mechanical Engineering. Report No. LIU-TEK-LIC-1994:07. University of Linko¨ping, Linko¨ping, Sweden. Gordon, E., von Holst, H. and Rudehill, A. (1995) Outcome of head injury in 2298 patients treated in a single clinic during a 21-year period. Journal of Neurosurgical Anesthesiology 7, 4, 235–247. Halldin, P. (1996) Evaluation of material parameters for safety belt. Report No 96-08. Department of Aeronautics. Royal Institute of Technology, Stockholm, Sweden. Hallquist, J. O., Stillman, D. W., Hughes, T. J. R. and Tarver, C. (1990) Modeling of air bags using MVMA/ DYNA3D. Livermore Software Technology Corporation, Livermore, CA. von Holst, H., Kraus, J. and Zitney, G. (1995) Prevention, Critical Care and Rehabilitation of Neurotrauma. Perspectives and Future Strategies. A Report to the World Health Organisation, ed. H. von Holst. Society of Automotive Engineers (1984) Human tolerance to impact conditions as related to motor vehicle design. Society of Automotive Engineers, Human Injury Criteria Task Force. SAE Handbook Supplement J885-84, Society of Automotive Engineers, Warrendale, PA. Kraus, J. (1991) Epidemiology features of injuries to the central nervous system. In Neuroepidemiology: Attributed to Bruce Schoenberg, ed. D. W. Andersson, pp. 333–357. CRC Press, Boca Raton, FL. Kraus, J. (1993) Epidemiology of head injury. In Head Injury, 3rd edition, ed. P. R. Cooper, pp. 1–25. Williams & Wilkins. Baltimore, MD. Kraus, J. and McArthur, D. (1995) Epidemiology of brain injury. In Neurology and Trauma, ed. R. Evans, W. B. Saunders, pp 3–17, Philadelphia, PA. Kraus, J., McArthur, D., Silverman, T.A., Jayarman, M. (1995) Epidemiology of Brain injury. In Neuro Trauma, eds R.K. Narayan, J.E. Wilberger, J.T. Povlishock and H. McGraw, pp. 13–30.
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Mertz, H. J. and Patric, L. M. (1971) Strength and response of human neck. In Proceedings of the 15th Stapp Car Crash Conference. SAE 710855. Mertz, H. J. (1984) Injury assessment values used to evaluate Hybrid III response measurements. NHTSA Docket 74-14, Notice 32, Enclosure 2 of Attachment I of part III of General Motors Submission USG 2284, 22 March. Miller, T. R., Pindus, N. M., Douglas, J. B. and Rossman, S. B. (1995) Databook on Nonfatal Injury: Incidence, Cost and Consequence. The Urban Institute Press, Washington, DC. Pintar, F. A., Yoganandan, N., Voo, L., Cusick, J. F., Maiman, D. J. and Sances, A. (1995) Dynamic characteristics of the human cervical spine. In Proceedings of 39th Stapp Car Crash Conference, San Diego, pp. 195–202, SAE/P-95/299. Rice, D. P., MacKenzie, E. and Associates (1989) Cost of Injury in the United States. Report to Congress. Stillman, D. W. and Hallquist, J. O. (1992a) LS-INGRID Users Manual, Version 3:1. LSTC Report No. 1019. Livermore. Stillman, D. W. and Hallquist, J. O. (1992b) LS-DYNA3D Users Manual, Version 920. LSTC Report No. 1007, rev 2. Livermore. Sosin, D. M., Sneizek, J. E. and Waxweiler, R. J. (1995) Trends in death associated with traumatic brain injury, 1979 through 1992. Journal of the American Medical Association 273, 22, 1778–1780. Viano, D., von Holst, H. and Gordon, E. (1997) Traumatic brain injury from traffic related causes: Priorities for primary prevention. Accident Analysis and Prevention 29, 6, 811–816. Waxweiler, R. J., Thurman, D., Sniezek, J. E., et al. (1995) Monitoring the Impact of traumatic brain injury: a review and update. Journal of Neurotrauma 12, 4, 509–526. Yoganandan, N., Ray, G., Pintar, F. A., Myklebust, J. B. and Sances, S. (1989) Stiffness and strain energy criteria to evaluate the threshold of injury to an intervertebral joint. Journal of Biomechanics 22, 2, 135–142.
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AN EXPERIMENTAL HEAD RESTRAINT CONCEPT FOR PRIMARY PREVENTION OF HEAD AND NECK INJURIES IN FRONTAL COLLISIONS P. H1*, H. H12 and I. E1 1Division of Lightweight Structures, Royal Institute of Technology, S-100 44 Stockholm, Sweden and 2Department of Clinical Neuroscience, Neurosurgery, Karolinska Institute, S-171 76, Stockholm, Sweden (Received 30 September 1996; in revised form 20 March 1997)
Abstract—The Experimental Head Restraint Concept ( EHRC ), a ‘safety belt’ for the head, is designed to reduce forces to the head and neck, in frontal car crashes. The EHRC was evaluated experimentally in frontal collision for a crash severity of 11 m/s, and numerically in frontal collision for a crash severity of 11 and 15 m/s. Experimental data obtained from a frontal barrier test (11 m/s) showed a 67% reduction of the HIC value from 411 (without EHRC ) to 136 (with EHRC ). The same level of reduction was also obtained for the higher speed in the numerical simulation. The moment in the neck was shown in experimental configuration to increase a few percent using the EHRC, but as presented in a numerical analysis, the moment was reduced by stiffening the EHRC. The EHRC clearly has a potential role in the search for primary prevention of neurotrauma injuries in frontal related car crashes. However, there is a strong need for more advanced injury criteria for the neck in order to optimize such complex safety systems. © 1998 Elsevier Science Ltd. All rights reserved Keywords—Head restraint, Injury prevention, Safety systems, Frontal collision
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system. As new car models are introduced, awareness of both passive and active prevention can be heightened. Traffic-related injury to the brain, like other causes, is of two categories: focal and diffuse injury. The focal brain injury is a lesion causing local damage that can be seen by the naked eye. The diffuse brain injury is associated with global disruption of brain tissue usually and is invisible. Both categories of injury are related to transfer of energy from an external impact to the head into the brain tissue. The translational and rotational impact causes a head acceleration often resulting in both focal and diffuse injury. By reducing the head acceleration due to impact, both categories of brain injury can theoretically be diminished significantly. Safety belts and airbags have considerably reduced the number of injuries from traffic accidents; however, both have their limitations. The safety belt reduces the motion of the body but leaves the head and neck unprotected. The airbag used today protects a front passenger in a frontal collision, but does not consider multidirectional forces or the rear seat passenger. There is a need to develop safety systems aimed toward protecting the head and neck and preventing neurotrauma during frontal related impacts. Thus, the aim of the present investigation
Neurotrauma, or injury to the central nervous system, is the most serious type of traumatic injury. Traumatic brain injury accounts for more than a third of all injury deaths (Rice et al., 1989; Kraus, 1991; Kraus, 1993; Kraus and McArthur, 1995; Kraus et al., 1995; Viano et al., 1997). After brain damage, many who survive live with impaired brain function. The toll from spinal cord injury is just as distressing, although the rate of occurrence is lower. The cost to society of neurotrauma is tremendous because of the high rate of incidence among the young (Rice et al., 1989, von Holst et al., 1995; Miller et al., 1995; Kraus et al., 1995). About 40% of all neurotrauma injuries occur as a result of traffic accidents ( Kraus, 1991; Gordon et al., 1995, von Holst et al., 1995; Sosin et al., 1995; Waxweiler et al., 1995; Viano et al., 1997). Current trends suggest an increase in the number of survivors of neurotrauma throughout the world (von Holst et al., 1995). High priority should be given to the long-term goals of prevention, treatment, and rehabilitation of people with damage to the central nervous *Corresponding author. Tel: +46 8 790 6448; Fax: +46 8 207865; e-mail: [email protected] 535
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was to evaluate the design of a new head protection concept, with the possibility of determining the retardation behavior of the head during the entire impact course, using both experimental and numerical frontal barrier tests. E XPE RIM EN TAL HE AD RE ST RAIN T C O N CE P T ( EH R C ) It should be stressed from the beginning that the EHRC was designed solely as an experimental tool for investigating ways to reduce trauma to both the head and neck in frontal associated crashes. As shown in Fig. 1, it consists of: (a) a curved aluminum pipe containing (b) a linear elastic spring; (c) a vertical aluminium plate, which replaces the neck rest and is bolted to the chair frame; and (d ) two endgrips that attach the spring to (e) a solid rod that is bolted to the vertical plate. The EHRC is adjustable vertically on the plate as shown in Fig. 2, and could be described
Fig. 1. Geometry of the EHRC (millimetres).
Fig. 2. Vertical adjustment of the EHRC.
as a ‘safety-belt’ for the head. The function of the EHRC is to decrease smoothly and continuously the G-forces that push the head forward and to keep the head from impacting surrounding objects in a collision. In comparison with the conventional airbag, the EHRC acts on the head as soon as the head leaves the neck rest. The spring retraction force keeps the head in position, and the length of time during which the head and neck are protected is therefore increased. The innovation was designed primarily to reduce the HIC (Head Injury Criteria) value, and secondarily to reduce forces to the neck by theoretically keeping the spine and head in a straight line during a frontal crash. In the experimental test, a spring constant of 1.5 kN/m was chosen to evaluate the effectiveness of the EHRC. E XPE RIM EN TAL SLE D T EST Method Two sled tests, one with the EHRC and one without, were carried out on the 50-metre crash court of VTI ( Va¨g och Transportforsknings Institutet, Linko¨ping, Sweden). The Hybrid III dummy developed in 1975 was chosen as the occupant substitute. The selected measured quantities were head resultant acceleration, moment in neck around occipital condyles, axial force in neck, shear force in neck, and resultant acceleration in thorax. The rotational acceleration was not measured as the Hybrid III dummy was without an angular accelerometer. The dummy was constructed and calibrated for frontal collision speeds of between 10 and 15 m/s; a speed of 11 m/s (40 km/h) was chosen for this study. It is our belief that the EHRC has better protection characteristics in the higher range of crash severity, and therefore, a lower speed was chosen. The dummy was placed in a standard Volvo 740 front seat equipped with a safety belt and the EHRC. The seat was bolted to a four-wheeled sled as shown in Fig. 3 and pulled by a cable to the selected crash speed. The retardation pulse was set to a level of about 18 gravitational units over a range of 90 milliseconds for both tests. Results The test results are shown as functions of time in Fig. 4. The peak translation acceleration of the center of the head was reduced by 32% [Fig. 4(a)]. To compare the acceleration of the head to the head injury criteria, the HIC values were calculated (Society of Automotive Engineers, 1984). With the EHRC, the HIC value was reduced by 67% (from 411 to 136). A small increase (13%) in the peak Y-moment of the neck was seen with the EHRC [Fig. 4(b)]. This result was expected because the
An experimental head restraint concept for primary prevention
Fig. 3. Experimental sled test.
EHRC acts at the top of the head, and a resulting moment is unavoidable. The peak axial and shear forces in the neck were reduced by 37 and 43% [Fig. 4(c,d)], respectively, with the EHRC. The peak acceleration of the thorax was reduced by less than 10% with the EHRC [Fig. 4(e)]. The results show that, compared to injury assessment reference values for the Hybrid III dummy (Mertz, 1984), the chosen collision speed results in responses that are below the respective threshold values for a midsize male. A higher crash severity will be analyzed in the Comparison of Results section. FI NIT E E LE ME NT ANA LYSIS Generally, for this type of analysis, implicit rigid body dynamic programs, such as MADYMO, are the most time-efficient tools. However, in order to evaluate an FE-package and an FE-model of the Hybrid III-dummy ( Fredriksson, 1994), we chose LS-DYNA3D as the numerical program. LS-DYNA3D is an explicit three-dimensional finite element code designed to analyze the large deformation dynamic response of elastic or rigid structures (Stillman and Hallquist, 1992b). The FE model consists of five components, as shown in Figs 5–7: Hybrid III dummy, three-point safety belt, EHRC, chair and floor, and airbag. The deformable model of the occupant substitute, developed by Fredriksson in the pre-processor LS-INGRID (Stillman and Hallquist, 1992a), consists of 17 body sections, which can be chosen as rigid or deformable. The model fulfils the criteria set by U.S. regulations for crash dummies, and gives
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results that are consistent with rigid body dynamic models such as MADYMO ( Fredriksson, 1994). The safety belt was modelled with shell elements and pin-joined truss elements. Triangular shell elements were used where the belt was assumed to contact the dummy; two truss elements, which can carry only tensile loads, were positioned where no contact was expected. Elastic-plastic material behaviour was assumed, with a Young modulus of 2,400 megapascals, a hardening modulus of 1,600 megapascals, and a yield stress of 69 megapascals ( Halldin, 1996). On the Volvo 740 seat, the upper section of the belt leads down from the retractor through a slipring to a pretensioner. Problems occurred while attempting to model this part of the belt because of the small radius of the slipring. The problem was solved by shortening the length of the belt and decreasing the elastic modulus of the material. LS-DYNA3D does contain an algorithm for defining the safety belt with slipring and pretensioner, but the characteristics of these parts are unknown. This work can be left to future researchers. The EHRC was modelled using a rigid 180-degree torus made up of eight-node brick elements; the torus was fastened to the seat with two undamped linear springs (Fig. 6). The chair and floor were modeled using Belytschko–Lin–Tsay shell elements made up of an elastic material. The airbag model was developed by Hallquist et al. (1990). COM PARISO N OF R ESU LT S Three major results are presented in this work: (a) FEM results for the EHRC and airbag for crash severities of 11 and 15 m/s; (b) the effect of spring stiffness on head and neck forces; and (c) comparison between experimental and numerical results. The measured quantities in the numerical studies were head resultant acceleration, moment in neck around occipital condyles, and resultant acceleration in thorax. FEM results for EHRD and airbag: dn=11 and 15 m/s Four different configurations of safety equipment were investigated with the Hybrid III dummy exposed for the crash severities of 11 and 15 m/s: (a) safety belt (SB) only; (b) SB and EHRC; (c) SB and airbag (AB); and (d ) SB, AB, and EHRC. Safety belts were used in all four configurations; comparison of the configurations thus allows evaluation of the effectiveness of the EHRC and airbag, both separately and in combination. The EHRC was equipped with a spring constant of 1500 N/m in the 11 m/s test and 3000 N/m in the 15 m/s test. FEM results for the four
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Fig. 4. Experimental results: With and without EHRC. Dn=11 m/s. (a) Resultant acceleration of head. (b) Moment in neck around occipital condyles, Y-axes. (c) Axial force in neck. (d ) Transverse shear force in neck. (e) Resultant acceleration in thorax.
configurations are shown for the head, neck and thorax in Figs 8 and 9. As expected, the safety belt alone gave the worst result, and the safety belt plus airbag with EHRC gave the best result. HIC values were calculated from the acceleration plots [Figs 8 and 9(a)] for each configuration and are shown in Fig. 10. The percentage improvement of the HIC values from one configuration to another is summarized in Table 1. Comparison of the safety belt with and without the EHRC shows an improvement of 37%. Comparison of the safety belt with and without the EHRC shows an improvement of 56%. Note that adding the EHRC to the safety belt plus airbag configuration gives a substantial additional improvement in the HIC value (from 37 to 60% and from 51% to 68%) but that adding the airbag to the safety belt plus EHRC configuration gives little or
no additional improvement (from 56 to 60% and none). Comparison of the results for the neck was much more difficult since no advanced criteria for the neck exist. It was impossible to analyze the neck moment curves in Figs 8(b) and 9(b) using the amplitude level criterion of Mertz and Patric (1971). However, it is clear that the characteristics of the moment acting on the neck differ considerably between configurations. The airbag and safety belt complemented by the EHRC appears to give the best response, whereas the EHRC and safety belt combination results in the highest moment forces. This negative effect of the EHRC was also shown in the experimental study. The resultant acceleration of the thorax is shown in Figs 8(c) and 9(c). The variations between the
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Fig. 5. FE model of the sled test: (a) EHRC with spring and brick element. (b) Safety belt with triangular shell and truss element. (c) Hybrid III dummy ( Fredriksson, 1994). (d) Chair and floor with shell element.
Fig. 6. FE model of the EHRC (side and top views).
Fig. 8. FEM results for four configurations of safety equipment. Dn=11 m/s.
To analyze a new concept, such as the EHRC, critically, and to optimize its properties, we will need better and more reliable injury criteria for the neck. For the head and neck, the FE analysis shows that the EHRC does give an improvement over the airbag; however, the combination of the two (with safety belt) clearly gives the best result. Fig. 7. FE model of the airbag.
different configurations are within 10%, and are explained by the low crash severity. A higher crash speed reveals the benefits of the airbag in reducing acceleration of the thorax [Fig. 9(c)].
Effect of EHRC spring stiffness on acceleration of the head and moment in the neck: Dn=11 m/s The effect of EHRC spring stiffness on acceleration of the head and the moment in the neck was studied numerically for the safety belt plus EHRC configuration using spring constants of 1,500, 3,000,
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P. H et al. Table 1. Percentage improvement of HIC values with different configurations of safety equipments Crash severity Safety equipment SB versus SB+AB SB versus SB+EHRC SB versus SB+AB+EHRC
11 m/s
15 m/s
37% 56% 60%
51% 68% 68%
and 3,750 N/m. The results are shown in Fig. 11. The acceleration of the head shows the same characteristics and approximately the same HIC value for all spring constants [Fig. 11(a)]; this is primarily the result of keeping the head from impacting the chest. The moment in neck around the occipital condyle shows a strong relationship between spring stiffness and moment [Fig. 11(b)]. A stiff spring removes most of the positive part in the curve while increasing the negative part by a few percent. The amplitude level criteria by Mertz and Patric (1971) state only that injury might occur below −57 Newton metres and above +190 Newton metres. Choosing the ‘best’ spring, therefore, is not a trivial matter as a small negative increase can be worse than a large positive decrease in the moment of the neck. Figure 11(b) clearly shows the need for advanced neck injury
Fig. 9. FEM results for four configurations of safety equipment. Dn=15 m/s.
Fig. 10. Comparison of different configurations of safety equipments for two levels of crash severity. HIC values (typed on the columns) are normalized against the configuration with safety belt only. SB, safety belt; AB, airbag.
Fig. 11. Effect of EHRC spring stiffness. Dn=11 m/s.
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criteria. We propose using a stiffness of 3000–3750 N/m for this crash severity (Dn=11 m/s). Comparison of experimental and numerical results For the safety belt plus EHRC configuration, a direct comparison of the experimental and numerical results was made for the videotape and computer animation, the acceleration plots for the head and thorax, and the moment plots for the neck. The comparisons are shown in Figs 12 and 13. The experimental videotape and computer animation (Fig. 12) show a reasonable correlation in the way in which the Hybrid III dummy moves. Some differences are evident, though, particularly toward the end of the test in which the dummy in the numerical simulation
Fig. 13. Graphical comparison of experimental (dotted line) and numerical (solid line with squares) results for safety belt with EHRC. Dn=11 m/s.
Fig. 12. Comparison of experimental videotape and computer animation for Hybrid III dummy equipped with safety belt and EHRC. Dn=11 m/s.
rebounded earlier. The reasons for this will be discussed below. In Fig. 13(a), a comparison of the results of the acceleration of the head shows an excellent correlation. However, significant differences are evident in comparing the results of the neck and thorax [Fig. 13(b,c)]. Both the moment in the neck and the resultant acceleration of the thorax show an earlier increase in the responses in the numerical simulation than in the experimental. The slope of the numerical curve in Fig. 13(b), however, shows the correct characteristics; therefore, the neck response can be used in comparison analysis. A comparison of results with and without the EHRC (for both experimental and numerical simulations), shows a decrease in thoracic acceleration of less than 10%, which emphasizes the fact that the responses of the thorax are of secondary
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interest in this study. Head and neck responses are of primary interest, and we believe that the present model performs well enough to be used for comparison analysis in this study, but not for a quantitative analyze aiming at predicting injury or no injury. A number of assumptions and simplifications may explain some of these differences. The FE model of the neck was calibrated to the standard Hybrid III pendulum test ( Fredriksson, 1994) and not to the sled test. The thorax was assumed to be rigid, which is one reason for the earlier increase in slope in Fig. 13(c). The experimental sled test and the FE simulation differed in three primary areas: (a) the safety belt, (b) the seat, and (c) the neck rest. The retractor, pretensioner, and slipring of the safety belt were not modeled. The belt dimensions, material properties, and contact definition were all correct. The seat, consisting of foam, a spring system, and bars, was modeled with an elastic plate. These modeling simplifications may explain the mismatch between experimental and numerical results for the acceleration of the thorax [Fig. 13(c)]. The neck rest, to which the EHRC was attached, was joined to the upper bar of the seat back. This bar was deformed in the experimental crash, and resulted in a forward bending of the neck rest ( Fig. 12). It was not possible to reproduce this bending in the FE model, which made the model stiffer in comparison to the test [Fig. 13(b)]. To obtain a better correlation, we suggest that simplifications be made to the experimental test configuration, to allow more accurate modeling of the chair and safety belt construction. D ISCUS SION This experimental and numerical study shows clearly that the Experimental Head Restraint Concept can be an alternative or a complement to conventional safety systems in frontal barrier tests. Two different crash severities were studied, 11 and 15 m/s, and show about the same improvement in HIC value (60 and 68%), using the EHRC plus safety belt compared to the safety belt alone. The stiffness of the EHRC used in the experimental tests, 1500 N/m, resulted in a small increase in the moment of the neck around the occipital condyles. Finite element analyses showed that the moment in the neck could be reduced by using a stiffer spring. The EHRC was designed to be used as an experimental tool. However, we must ask ourselves: could innovation be used as a complement to the conventional safety systems in a car of tomorrow, and what are the potential injuries that may result from the experimental concept? A future design of the EHRC can be either constantly active (as is the
present EHRC ) or activated at the time of impact (as an airbag). Both systems have their advantages and disadvantages. The constantly active system has the benefit of acting during the whole time of impact if it is positioned tightly next to the head. It can be designed to work in a wide range of collision angles, with zero retraction force in the normal range of motion (much like a conventional safety belt), but has a major disadvantage regarding user-friendliness. Probably, only a few will use a constantly active system voluntarily unless the car is equipped with an intelligent safety control system, which implies that the speed is reduced if the EHRC is not in use. A design that is activated at the time of impact is more attractive, with regard to user-friendliness, but a large amount of effort has to be focused on the injury potentials related to out-of-position occupants as the equipment is activated. The major risk attached to the EHRC is neck injury, which may be caused by the added neck moment around the occipital condyles. There is also risk of WAD injuries using the present design with the linear elastic spring producing a rebounding effect on the head and neck system. This risk, however, could be reduced or eliminated by using a non-linear elasto-plastic spring. It was not possible to test the hypothesis of maintaining a straight line between head and spine to prevent neck injuries, but this remains a goal for us to verify. Further investigations may include: a study of the spring material characteristics; analyses of multidirectional impacts as off-set collisions; and analysis of head rotational acceleration, either as a single physical quantity or calculated in more advanced head injury criteria. In addition, more advanced injury criteria are needed for the neck. A number of investigators (e.g. Yoganandan et al., 1989’ Bilston et al., 1993; Eppinger et al., 1994; Pintar et al., 1995) are currently studying both fracture and neurotrauma criteria for the neck, but much work remains to be done. Improved criteria will lead to improved interpretation of experimental and numerical results. Acknowledgements—We want to thank Lars Fredriksson at the University of Linko¨ping, and Niclas Bra¨nnberg at Engineering Research, for their support with the FE model. Thanks to Ingemar Gustavsson at Swedish Road and Transport Research Institute for his help with the experimental part. We also want to acknowledge Mrs Deborah K. Kerwin-Peck for editing the manuscript.
R EF ER EN C ES Bilston, L. E., Meaney, D. F. and Thibault, L. E. (1993) The Development of a Physical Model to Measure Strain in a Surrogate Spinal Cord During Hyperflexion and Hyperextension. Proceedings. IRCOBI, pp. 255–265. Eppinger, R. K., Kleinberger, M., Morgan, R., Khaew-
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pong, N., Bandak, F. and Haffner, M. (1994) Advanced Injury Criteria and Crash Evaluation Techniques. National Highway Traffic Safety Administration. Paper No. 94-S1-O-11. Fredriksson, L. A. (1994) Finite Element Modeling of Occupant Substitutes. Department of Mechanical Engineering. Report No. LIU-TEK-LIC-1994:07. University of Linko¨ping, Linko¨ping, Sweden. Gordon, E., von Holst, H. and Rudehill, A. (1995) Outcome of head injury in 2298 patients treated in a single clinic during a 21-year period. Journal of Neurosurgical Anesthesiology 7, 4, 235–247. Halldin, P. (1996) Evaluation of material parameters for safety belt. Report No 96-08. Department of Aeronautics. Royal Institute of Technology, Stockholm, Sweden. Hallquist, J. O., Stillman, D. W., Hughes, T. J. R. and Tarver, C. (1990) Modeling of air bags using MVMA/ DYNA3D. Livermore Software Technology Corporation, Livermore, CA. von Holst, H., Kraus, J. and Zitney, G. (1995) Prevention, Critical Care and Rehabilitation of Neurotrauma. Perspectives and Future Strategies. A Report to the World Health Organisation, ed. H. von Holst. Society of Automotive Engineers (1984) Human tolerance to impact conditions as related to motor vehicle design. Society of Automotive Engineers, Human Injury Criteria Task Force. SAE Handbook Supplement J885-84, Society of Automotive Engineers, Warrendale, PA. Kraus, J. (1991) Epidemiology features of injuries to the central nervous system. In Neuroepidemiology: Attributed to Bruce Schoenberg, ed. D. W. Andersson, pp. 333–357. CRC Press, Boca Raton, FL. Kraus, J. (1993) Epidemiology of head injury. In Head Injury, 3rd edition, ed. P. R. Cooper, pp. 1–25. Williams & Wilkins. Baltimore, MD. Kraus, J. and McArthur, D. (1995) Epidemiology of brain injury. In Neurology and Trauma, ed. R. Evans, W. B. Saunders, pp 3–17, Philadelphia, PA. Kraus, J., McArthur, D., Silverman, T.A., Jayarman, M. (1995) Epidemiology of Brain injury. In Neuro Trauma, eds R.K. Narayan, J.E. Wilberger, J.T. Povlishock and H. McGraw, pp. 13–30.
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Mertz, H. J. and Patric, L. M. (1971) Strength and response of human neck. In Proceedings of the 15th Stapp Car Crash Conference. SAE 710855. Mertz, H. J. (1984) Injury assessment values used to evaluate Hybrid III response measurements. NHTSA Docket 74-14, Notice 32, Enclosure 2 of Attachment I of part III of General Motors Submission USG 2284, 22 March. Miller, T. R., Pindus, N. M., Douglas, J. B. and Rossman, S. B. (1995) Databook on Nonfatal Injury: Incidence, Cost and Consequence. The Urban Institute Press, Washington, DC. Pintar, F. A., Yoganandan, N., Voo, L., Cusick, J. F., Maiman, D. J. and Sances, A. (1995) Dynamic characteristics of the human cervical spine. In Proceedings of 39th Stapp Car Crash Conference, San Diego, pp. 195–202, SAE/P-95/299. Rice, D. P., MacKenzie, E. and Associates (1989) Cost of Injury in the United States. Report to Congress. Stillman, D. W. and Hallquist, J. O. (1992a) LS-INGRID Users Manual, Version 3:1. LSTC Report No. 1019. Livermore. Stillman, D. W. and Hallquist, J. O. (1992b) LS-DYNA3D Users Manual, Version 920. LSTC Report No. 1007, rev 2. Livermore. Sosin, D. M., Sneizek, J. E. and Waxweiler, R. J. (1995) Trends in death associated with traumatic brain injury, 1979 through 1992. Journal of the American Medical Association 273, 22, 1778–1780. Viano, D., von Holst, H. and Gordon, E. (1997) Traumatic brain injury from traffic related causes: Priorities for primary prevention. Accident Analysis and Prevention 29, 6, 811–816. Waxweiler, R. J., Thurman, D., Sniezek, J. E., et al. (1995) Monitoring the Impact of traumatic brain injury: a review and update. Journal of Neurotrauma 12, 4, 509–526. Yoganandan, N., Ray, G., Pintar, F. A., Myklebust, J. B. and Sances, S. (1989) Stiffness and strain energy criteria to evaluate the threshold of injury to an intervertebral joint. Journal of Biomechanics 22, 2, 135–142.