Online supplement: Head Kinematics and Shoulder Biomechanics in Shoulder Impacts similar to Pedestrian Crashes a THUMS study

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1 Online supplement: Head Kinematics and Shoulder Biomechanics in Shoulder Impacts similar to Pedestrian Crashes a THUMS study Ruth Paas a, Johan Davidsson a, Karin Brolin a a Department of Applied Mechanics, Chalmers University of Technology, Gothenburg, Sweden

2 APPENDIX A PARAMETER VARIATION STUDY ON PEDESTRIAN-LIKE SHOULDER IMPACTS Introduction to the parameter variation study Shoulder impact conditions in full-scale pedestrian impacts appear to be different from typical impactor test conditions, based on the available video footage in full-scale pedestrian experiments (Paas et al. 2012, Schroeder et al. 2008, Kerrigan et al. 2005a, 2005b and 2007). Firstly, the impact was not generally perpendicular as in a typical impactor test. Instead, the shoulder tended to roll off the vehicle from the distal to the proximal upper arm. Secondly, the angle between the impacting surface and the midsagittal plane varied depending on whether the bonnet or windshield was impacted (Figure A-1). When the bonnet was impacted, the vertical angle was approximately 0. When the windshield was impacted, the vertical angle was approximately 30. Depending on the rotation of the pedestrian s upper body before shoulder impact, an additional anterior or posterior component was observed. Thirdly, the pedestrian upper body rotated towards the vehicle surface, e.g., the bonnet, while the vehicle moved towards the pedestrian at the same time. Combining the motion of the pedestrian and the vehicle yields the resultant shoulder impact velocity, which was generally from a supero-lateral direction (Figure A-1). Such shoulder impacts and their influence on head kinematics have not been addressed in the literature to date, to the best knowledge of the authors. v res v subj v veh 0 30 v subj v veh v res Pedestrian-to-vehicle Shoulder-tobonnet Shoulder-towindshield Figure A-1: Left: Schematic image of one possible shoulder impact as seen in a vehicle (light grey) vs. pedestrian (black) experiment (Paas et al. 2012). The sum of the vehicle velocity (light grey arrow v veh ) and the shoulder velocity (black arrow v subj ) is the resultant shoulder impact velocity (dark grey arrow v res ). Right: Turning the image around to match a typical impactor experiment, the impact surface is oriented laterally or supero-laterally. The resultant shoulder impact velocity is from a supero-lateral direction. In the experiment, the impacted shoulder was in an elevated posture at the time of impact. 2

3 The main aim of the parameter variation study is therefore to assess the influence of impact angles and shoulder postures, similar to pedestrian shoulder impacts, on head kinematics. A secondary aim is to investigate shoulder girdle biomechanics and shoulder-to-spine coupling in these impacts in order to increase the understanding of head boundary conditions prior to head impact in pedestrian crashes. elevated by 70 mm, anteriorly displaced by 50 mm, and medially displaced by 20 mm. Methods of the parameter variation study In the parameter variation study, simulation of the purely lateral volunteer experiments (Ono et al. 2005) used for THUMS evaluation was used as a reference case. The following parameters were varied (Figure A-2): shoulder initial posture (neutral, elevated or anteriorly-elevated), impactor surface orientation α (0 or 30 ), and impact velocity direction β relative to the impactor surface (90 or 45 ). Prior to each simulation, the impactor was repositioned in order to maintain the distance between the middle of the impactor surface and the mid-point of the humeral head. The choice of angles and shoulder posture in the parameter variation study was based on examination of data and image sequences from Kerrigan et al. (2005a), (2005b), (2007), and Paas et al. (2012). Repositioning of the shoulder was achieved by applying force to the head of the humerus while immobilising the thorax, spine, and head, so that the scapula, clavicle, and the surrounding soft tissue and ligaments deformed accordingly. In the neutral shoulder posture, the original THUMS shoulder was used. In the elevated posture, the head of the humerus was elevated by 70 mm and medially displaced by 20 mm. In the anteriorlyelevated posture, the head of the humerus was Figure A-2: Simulation matrix and naming convention in the parameter variation study. Four impactor configurations (left) were combined with three shoulder postures (top, in lateral and top view). The impactor configurations differed in the impactor surface orientation α and in the impact velocity direction β relative to the impactor surface, where 1: (α = 0, β = 90 ), 2: (α = 0, β = 45 ), 3: (α = 30, β = 90 ) and 4: (α = 30, β = 45 ). The shoulder postures were N = neutral, E = elevated and A = anteriorly-elevated. Data output were: contact forces between the impactor and THUMS, nodal data from the acromia, the head CG, the sternum upper end, and the vertebral body centres of T1. All output was filtered with a CFC 60. 3

4 Results of the parameter variation study This section contains the head linear displacements, the shoulder deflections and the contact forces measured in the parameter variation study. Within each paragraph, the effects of changing the shoulder posture is investigated first; the effects of changing the impactor configuration is presented afterwards. Impactor configuration 4 was a special case and is presented separately from the other configurations in the end of this section. Head y-displacements increased by approximately 50 % when the shoulder posture was changed from the neutral to the elevated and anteriorly-elevated postures, i.e., comparing E1 and A1 to N1, E2 and A2 to N2, and so forth (Figure A-3). In contrast, head y-displacements decreased when the impactor configuration was changed from configuration 1 to the other impactor configurations, i.e., when the impact direction was changed from lateral to superolateral. Generally, the greatest head y-displacements were observed with impactor configuration 1. The simulations N2 and N3 produced similar head y-displacements (Figure A-3, top row). Thus, changing the impact velocity direction only (N2) and changing the impactor surface orientation only (N3) had a similar effect on the head y-displacement in the neutral shoulder posture, compared with the reference case. In the elevated shoulder posture, changing the impactor surface orientation only (E3) affected the head displacements only slightly compared with the lateral impact (E1, Figure A-3, middle row). For both the elevated and anteriorly-elevated postures, impactor configuration 3 resulted in greater head y-displacements than impactor configuration 2. In general, the impact velocity angle β reduced the head y-displacements more than the impactor surface orientation angle. Head linear x- and z- displacements were consistently short (Figure A-3). Compared with the neutral shoulder posture, the elevated and anteriorly-elevated postures resulted in reduced peak acromion-to-sternum and acromion-to-t1 deflections for the impact configurations 1-3 (Figure A-4). In conjunction with the increased head y-displacements in these postures, this indicates stronger shoulder-to-spine coupling (Appendix D) in the elevated and anteriorly-elevated postures compared with the neutral posture. Different impactor configurations did not considerably influence acromion-tosternum deflections, the only exception being impactor configuration 4. A more substantial influence of the impactor configurations was observed in the acromion-to-t1 deflections. Comparing simulations E2 and E4 to E1, and A2 and A4 to A1, the acromion-to-t1 deflections were reduced for the impactor configurations 2 and 4 (β = 45, Figure A-4), and the peak deflection occurred earlier. In impactor configuration 3, the magnitude of peak acromion-to-t1 deflection was similar to impact configuration 1, but the peak occurred earlier for the elevated and anteriorly-elevated shoulder postures. The measured contact force in the neutral shoulder posture was considerably lower than in the elevated and anteriorly-elevated postures (Figure A-5). In simulations E1 and A1, the peak contact force was approximately 25 % higher than in simulation N1. Since the impactor velocity was prescribed, the increased contact force indicates stronger shoulder-to-spine coupling in the 4

5 elevated and anteriorly-elevated shoulder postures compared with the neutral posture. The contact forces did not change considerably when the impactor configuration was changed from lateral to supero-lateral in configurations 1, 2 and 3, indicating that the results in these simulations could be reasonably compared. Figure A-3: Head linear x- (left), y- (middle column), and z-displacements (right). Top: neutral shoulder posture (N1-N4), middle row: elevated shoulder posture (E1-E4), bottom: anteriorly-elevated shoulder posture (A1-A4). Impactor configuration 4 ( = 30, β = 45 ) was a special case since the contact force was considerably lower than in the other impactor configurations. Compared with configuration 1, the peak contact force in configuration 4 was reduced by 30 % - 56 % (Figure A-5). As the impactor velocity was prescribed and the impactor material was nearly rigid, the measured contact force depended mainly on the effective mass and stiffness of THUMS in each impact. In configuration 4, the shoulder was rather movable. Due to this, the impactor did not encounter the same amount of effective mass as in the other impactor configurations. Hence, only a small 5

6 amount of the impactor energy resulted in lateral motion of the upper body of THUMS. Accordingly, head y-displacements and shoulder deflections were considerably reduced in configuration 4, especially in the neutral shoulder posture. Figure A-4: Shoulder deflections (left: impacted acromion-to-sternum, right: impacted acromionto-t1). Top: neutral shoulder (N1-N4), middle row: elevated shoulder (E1-E4), bottom: anteriorly-elevated shoulder (A1-A4). Figure A-5: Resultant contact forces in different impactor orientations ( ) and impact directions (β). Top: neutral (N1-N4), middle: elevated (E1- E4), bottom: anteriorly-elevated shoulder posture (A1-A4). 6

7 Discussion of the parameter variation study The main aim of the parameter variation study was to assess how impact angles and shoulder postures similar to pedestrian shoulder impact influence head kinematics. A secondary aim was to investigate shoulder girdle biomechanics and shoulder-to-spine coupling in these impacts. As described in the discussion section in the main paper, the impact velocity in Ono et al. (2005) was rather low compared with pedestrian experiments. However, these experiments were used as a reference case for the parameter variation study since the head linear displacements of THUMS were shown to be biofidelic in this impact condition. The angles and shoulder postures used in the parameter study were based on video footage and data from previous full-scale pedestrian experiments (Kerrigan et al. 2005a, 2005b and 2007, Schroeder et al and Paas et al. 2012). The authors of those studies have not specifically addressed shoulder impact conditions. However, the geometries of each vehicle s centreline were known. In combination with video footage depicting the subjects at the time of shoulder impact, the vehicle geometries were used to estimate the orientation of the impactor surface α. The vehicle velocity was reported in each pedestrian experiment. The velocity (speed and direction) of the shoulder, immediately prior to shoulder impact, was known from the scapula kinematics in Paas et al. (2012) and estimated from T1 kinematics in Kerrigan et al. 2005a, 2005b and 2007) and Schroeder et al. (2008). From this analysis, a reasonable range was determined for the direction of the shoulder impact velocity β. The shoulder posture after elbow impact in Paas et al. (2012) could be determined using the T2 and scapula kinematics. However, in the other pedestrian experiments, the shoulder posture prior to shoulder impact could only be estimated. The authors of the present study therefore decided to include a reasonable range of shoulder postures based on the available footage from the pedestrian experiments. The scapular motion over the thorax changed when the impact direction was changed to superolateral. The change in scapular motion appeared to influence the shoulder-to-spine coupling. Impacts from a superior direction pushed the lateral edge of the scapula downwards. In the neutral shoulder posture, the amount of coupling was slightly reduced with impactor configurations 2 and 3, and considerably with configuration 4, which was the most extreme angulation. Compared with the purely lateral impact (configuration 1), angulation of the impact speed only (configuration 2) or of the impactor surface only (configuration 3) had a similar effect on the scapular motion. In both cases, the amount of scapular medial sliding was similar to the lateral impact, indicating a similar amount of coupling, and the head lateral displacements were generally similar. Elevation of the shoulder reduced the scapular ability to slide over the thorax in all simulated impact directions. Visual inspection of the simulation results indicated that the scapular orientation and geometry reduced its ability to slide over the thorax. In the elevated posture, the scapula was adducted (medially displaced) and rotated upward. This brought the supero-medial edge close to the spine in the C6-T1 region. During shoulder impact, scapular motion towards T1 was therefore limited by the soft tissue close to 7

8 the spine. In the anteriorly-elevated posture, the scapula was rotated upwards and protracted. The impact then forced the scapula to move more laterally towards the rib cage, limiting posterior sliding over the thorax. Thus, in both the elevated and the anteriorly-elevated posture, shoulder-tospine coupling was increased due to limited medial sliding of the scapula, although through different mechanisms. In these shoulder postures, the acromion-to-t1 and acromion-to-sternum deflections were reduced whereas the contact forces and head lateral displacements were considerably greater than in the neutral posture. Conclusions from the parameter variation study Raised shoulder postures limited the scapular motion over the thorax, increasing the shoulderto-spine coupling and head linear displacements. Head linear displacements were decreased when the impactor orientation, the impact velocity direction or both were changed from lateral to supero-lateral. It is recommended to conduct shoulder impact experiments with human subjects which mimic pedestrian shoulder impacts, to validate the results from the present study and further assess head twist. 8

9 APPENDIX B VOLUNTEER AND PMHS DATA The volunteers in Ono et al. (2005) were, on average, lighter and smaller than THUMS 4.0 (Table B-1). The PMHSs in Bolte et al. (2003) were closer to THUMS in weight than the volunteers in Ono et al. (2005). The height of the PMHSs was not reported (Table B-1). Table B-1: Volunteer and PMHS data (average standard deviation) compared with THUMS 4.0 Age (years) Weight (kg) Height (cm) n Volunteers (Ono et al. 2005) PMHSs (Bolte et al. 2003) Not available 3 THUMS (TMC 2011)

10 APPENDIX C PREVIOUS THUMS 4.0 EVALUATION AND DIFFERENCES BETWEEN PEDESTRIAN AND OCCUPANT VERSION In Appendix C1, the main updates of THUMS 4.0 compared with earlier versions are described in short, and a table with previous evaluation studies of THUMS 4.0 is presented. In Appendix C2, the differences between the THUMS 4.0 pedestrian and occupant models are listed. Appendix C1: Previous Studies on THUMS 4.0 Evaluation THUMS 4.0 was modelled combining several techniques (Shigeta et al. 2009): The head model was taken from THUMS 3.0, the torso (including internal organs and pelvis) model was based on new CT scans (i.e., CT scans which were not used in previous model versions), and the extremities were based on ViewPointTM geometrical data. Since version 4.0 displays major differences regarding model geometry and modelling techniques compared with the previous versions, Table C-1 presents only those evaluation studies which were carried out with THUMS version 4.0. The criteria for inclusion of an evaluation study in this table was that THUMS 4.0 had to be compared with biomechanical data, i.e., kinematics and/or injuries from PMHSs, volunteers or real accidents in rather well-defined loading conditions. Studies in which THUMS was changed or used as a validated model for an application, e.g., for accident reconstructions on the basis of the pedestrian s injuries, were excluded. Table C-1: Previous evaluation of THUMS 4.0 Test Lateral head drop Frontal head impactor Head rotation (1) Head rotation (2) Thorax frontal impactor Thorax lateral impact Thorax dynamic belt compression Evaluated parameters Force over time Contact force, acceleration, and pressure in different brain regions Brain kinematics Brain injury Thoracic force-deflection, bone fractures, and organ strain and pressure distribution Force-deflection and number of rib fractures Reference to experimental data Yoganandan et al. (2004) Nahum et al. (1977) Hardy et al. (2001) and Kleiven et al. (2002) Gennarelli et al. (1982) Kroell et al. (1974) Viano (1989) Chest deflection Césari et al. (1990) Presented in Watanabe et al. (2011) and (2012) Watanabe et al. (2012) Watanabe et al. (2012) Watanabe et al. (2012) Shigeta et al. (2009) Watanabe et al. (2011) and (2012) Shigeta et al. (2009) 10

11 Abdomen frontal impactor Abdominal belt compression Humerus static 3- point bending Humerus dynamic compression Femur static 3-point bending 4-point knee bending Lateral knee impact Pedestrian impact with sedan Pedestrian impact with SUV Pedestrian impact with minivan Abdominal force-displacement and abdominal organ strain distribution Abdominal force-displacement and abdominal organ strain distribution Force-displacement Force-displacement Force-displacement Cavanaugh et al. (1986) Foster et al. (2006) Kemper et al. (2005) Kemper et al. (2005) Yamada et al. (1970) Bending moment over bending angle Bose et al. (2004) Upper tibia acceleration over time 2D kinematics of the head, T1, S1, impacted knee, and impacted heel as well as impact locations of these body parts on the vehicle, also rib, pelvis, and knee injuries 2D kinematics of the head, T1, L5, impacted knee, and impacted ankle (as well as impact locations of these body parts on the vehicle in the 2012 publication only), also rib fractures and knee ligament ruptures 2D kinematics of head, T1, L5, impacted knee, and impacted ankle as well as impact locations of these body parts on the vehicle, also rib and pelvis fractures and knee ligament ruptures and lower extremity fractures Kajzer et al. (1997) and Kajzer et al. (1999) Subit et al. (2008) Schroeder et al. (2008) Schroeder et al. (2008) Shigeta et al. (2009) Shigeta et al. (2009) Shigeta et al. (2009) Shigeta et al. (2009) Shigeta et al. (2009) Watanabe et al. (2011) and (2012) Shigeta et al. (2009) Watanabe et al. (2012) Watanabe et al. (2011) and (2012) Watanabe et al. (2012) Appendix C2: Differences between THUMS 4.0 Pedestrian and Occupant Version In this section, the differences between the THUMS 4.0 pedestrian and occupant version are presented, divided into differences in number of elements in each body part, materials, contact definitions and other differences. Differences in Number of Elements in different Body Parts In several body parts, different numbers of elements were found when comparing the pedestrian and occupant version, indicating that these parts were re-meshed. The authors hypothesise that this re-mesh was done to avoid skewed elements when changing the posture of the model. Regarding the mesh, no changes were made to adjust the occupant to the pedestrian version. 11

12 Regarding the shoulder, the soft tissue / flesh part surrounding each shoulder as well as the shoulder capsules were slightly re-meshed to account for a driver arm posture. However, the influence of this re-mesh on the results was considered to be small since the flesh thickness remained the same and the element size remained similar. In the upper extremities, only the elbow flesh, skin and one of the elbow ligaments were re-meshed. In the thoracic region, several spinal ligaments were slightly re-meshed. In the abdomen and pelvis region, it was only a deep flesh layer in the abdomen, the buttock flesh and skin and the pelvis joint capsules that were slightly re-meshed. The internal organs displayed different numbers of elements in nearly all parts; these were remeshed according to the posture of the occupant. In the lower extremities, the knee flesh and skin were re-meshed for the same reason. No differences in element numbers were found in the head and neck. Differences in Materials The following material differences in the original models were adjusted in the occupant version to match the pedestrian version. The Young s modulus of the skin all over the body was double in the occupant compared with the pedestrian. The Young s modulus of the intervertebral disk annuli was larger by a factor of 17 in the occupant. Strain rate effects were taken into account in a number of materials, using the Cowper and Symonds model which is based on scaling the yield stress according to the strain rate. Differences in the scale factor coefficients were adjusted to the pedestrian in the following cortical bones: vertebrae, sacrum, ribs, sternum, scapulae, clavicles, humerus, radius and ulna. Differences were also found in the spongy bones of the sternum and ribs, where the effective plastic strain at which the material fails was adjusted to the pedestrian. The following additional material differences were found, but not adjusted to the pedestrian version, as their influence on the results was deemed small. The cerebrospinal fluid had a different viscosity coefficient. The Cowper and Symonds coefficients for strain rate effects were different in the cortical bones of the head, but not adjusted to the pedestrian version. In the knee ligaments and the quadriceps femoris tendon, erosion was added in the pedestrian, but not in the occupant version. In addition, the quadriceps femoris tendon was modelled with a different material model and had a slightly different density. The knee ligaments had a different linear bulk modulus, which was larger by a factor of 3 in the occupant. Three shell parts were modelled with an elastic material in the pedestrian and a null material in the occupant: the lower deep-layer flesh, a tissue layer around the anterior and lateral lumbar and thoracic spine, and an additional skin layer on the buttocks. These parts also had half the density in the occupant compared with the pedestrian. Differences in Contact Definitions Several differences were found in the contact definitions in pedestrian and occupant version. Although it was not expected that these differences would lead to considerable changes in the results, all contact definitions in the occupant version were adjusted to the pedestrian version prior to evaluation against Bolte et al. (2003), for which the occupant version was used. 12

13 Specifically, in several automatic-singlesurface contacts with SOFT 2 contact definition, the parameter DEPTH was adjusted from originally 5 in the occupant (checking surface penetrations and edge-to-edge penetrations) to 13 as in the pedestrian (checking surface penetrations at nodes and at the edge with improved energy conservation). This update was implemented in the following occupant contact definitions: body, left leg, right leg, organs, head, skin. In addition, the automatic-single-surface contacts of the left and right arm of the occupant did not have the optional card A included, which enables more detailed control of contacts between materials with a wide variety in elastic bulk moduli. In the related pedestrian contacts, card A was included and specified a soft constraint formulation of the contact. The occupant contact definitions were updated accordingly. A number of tied-shell-edge-to-surface-beamoffset contacts were defined in the brain. In the occupant version, card A specified a soft constraint formulation for the contact, whereas in the pedestrian, card A specified a penalty formulation. The occupant definitions were adjusted to the pedestrian definitions. Other Differences Several hourglass definitions were present in the pedestrian organs but not in the occupant version; these were updated in the occupant to match the pedestrian. No differences were found regarding constrained (nodal) rigid bodies. 13

14 APPENDIX D SHOULDER-TO-SPINE COUPLING The amount of shoulder-to-spine coupling in shoulder impacts can be described, in a simplified way, as the product of forces that inhibit scapular motion over the thorax and the distance which the scapula has moved over the thorax. These parameters were not available from the experimental data. Instead, two main indicators for the amount of coupling were identified from the load path described in the main paper: the magnitude of acromion-to-t1 deflection and the amount of T1 y-displacement. It has to be kept in mind that the impactor velocity was prescribed and the contact force was generally similar in all evaluation simulations, i.e., the inserted energy was similar. The acromion-to-t1 deflection was a measure of the amount of medial sliding of the scapula, although scapular rotation contributed to the deflection to some extent. In addition, the scapular antero-posterior motion contributed to the amount of force transferred to the spinal column. The T1 y-displacement, one measure of the final link in the load path between shoulder and spine, was therefore considered as another indication for the amount of shoulder-to-spine coupling. A larger T1 y-displacement, considering a similar amount of momentum transferred to THUMS, generally suggested greater shoulder-tospine coupling. In the simulations, scapular kinematics could be observed, whereas, in the experiments, the scapular kinematics and the coupling could be estimated from acromion, T1 and head displacements. In the anterior impact, the scapula moved posteriorly in the experiments and the simulation, reducing the coupling between the subscapularis muscle and rib cage compared with the lateral impact. In the experiments, the weaker coupling was indicated by reduced T1 and head y- displacements. This reduction of displacements was also observed in THUMS, but to a lesser extent. Notably, in the anterior impact, the amount of medial sliding over the thorax for the relaxed volunteers and THUMS was comparable with the lateral impact, but the shoulder-to-spine coupling was reduced. The authors hypothesise that the reduced coupling was mainly due to the scapula being pushed posteriorly, lowering the friction force between the scapula, subscapularis and rib cage. In addition, part of the impact energy was used to push the scapula posteriorly, reducing the amount of lateral load. In the posterior impact, the lateral edge of the scapula was pressed into the subscapularis, leading to less scapular medial sliding than in the lateral impact. The reduced sliding, combined with slightly increased friction, appeared to slightly reduce T1 y-displacements in the posterior impact. The tensed volunteers exhibited similar head y-displacements, whereas THUMS displayed less head y-displacements in the posterior compared with the lateral impact. 14

15 APPENDIX E ADDITIONAL RESULTS FROM THE THUMS EVALUATION: HEAD AND SPINE KINEMATICS, SHOULDER DEFLECTIONS AND FORCE-DEFLECTION CURVES Appendix E1: Head rotation Head lateral rotation of THUMS increased slightly earlier than the average relaxed and tensed volunteers and remained within or close to the upper boundary of the volunteer corridors until 120 ms (Figure E-1, left column). The THUMS head returned to the neutral lateral orientation earlier and at a higher rate, compared with relaxed as well as with the tensed volunteers. Head y-rotations were consistently below 3 for all simulations and volunteers (Figure E-1, middle column). However, head twist was considerably greater in THUMS than in the volunteers (Figure E-1, right column). Figure E-1: Head angular x-, y- and z-displacements of THUMS (continuous lines) superimposed on relaxed (dashed lines) and tensed (dotted lines) volunteer corridors. Lateral (left), 15 anterior (middle), and 15 posterior impacts (right). 15

16 THUMS should therefore not be used to study head twist, which was considerably more pronounced than in the response of any volunteer. Various reasons were investigated. The THUMS head centre of gravity was in a biofidelic location with respect to the atlanto-axial joint, which is known to be one of the main contributors to head twist. Shifting the THUMS head centre of gravity backwards or changing the head initial posture (neck extension, neck retraction) within a reasonable range prior to the shoulder impact did not result in any considerable reduction of the head twist. Passive muscle contribution in THUMS was judged to be small and increased later than the head rotations; it was not considered responsible for the substantial twist. In order to balance the head, the volunteers activated their neck muscles and thus possibly displayed a different head rotation compared with the PMHSs. Appendix E2: Volunteer impact head and spine kinematics Comparison of THUMS 4.0 with relaxed and tensed volunteer responses: linear displacements of the head, C3, T1/T2 (the exact vertebra was not always known in the volunteer experiment), T4/T5, T8/T9, T12/L1, and L3. Lateral impact: 16

17 Figure E-2: Lateral impact: linear x (left), y (middle) and z (right) displacements of head and spine. THUMS (continuous lines) with relaxed (dashed lines) and tensed (dotted lines) volunteer corridors. 17

18 15 anterior impact: 18

19 Figure E-3: 15 anterior impact: linear x (left), y (middle) and z (right) displacements of head and spine. THUMS (continuous lines) with relaxed (dashed lines) and tensed (dotted lines) volunteer corridors. 15 posterior impact: 19

20 Figure E-4: 15 posterior impact: linear x (left), y (middle) and z (right) displacements of head and spine. THUMS (continuous lines) with relaxed (dashed lines) and tensed (dotted lines) volunteer corridors. 20

21 Appendix E3: Volunteer and THUMS forcetime and force-deflection curves The volunteers had a mass of 64 ± 12 kg (average ± standard deviation), whereas THUMS 4.0 has a total mass of 77 kg (Appendix B). It can thus be expected that the volunteers had a lower effective mass during shoulder impact than THUMS. The impactor velocity was prescribed in this setup. Thus, according to Newton s second law, the higher effective mass of THUMS has likely contributed to the higher contact force measured in the simulations compared with the experiments (Figure E-5). In addition, the shoulder-to-spine coupling and the THUMS stiffness were possible contributors to the amount of contact force. Figure E-5: Contact force; lateral (left), 15 anterior (middle column), and 15 posterior impact (right), THUMS (continuous lines) with relaxed (dashed lines) and tensed (dotted lines) volunteer corridors. 15 posterior Figure E-6: Individual contact forces over shoulder deflections: acromion-to-sternum (top), acromion-to- T1 (bottom). Lateral impact (left), 15 anterior impact (middle), 15 posterior impact (right). 21

22 Appendix E4: PMHS lateral impact, sternum and T1 linear y-displacements and acromion-toacromion shoulder deflection Figure E-7: Sternum (left) and T1 (right) linear y-displacements of THUMS (continuous lines) and PMHS corridors (dashed lines). APPENDIX F FILTER CLASS The filter class used for all data was CFC 60. The experimental PMHS data had a sampling rate of 10 khz and all nodal output from the simulations was sampled at 5 khz. For these data, a CFC 180 filter would have been sufficient. However, the experimental volunteer data as well as the contact force and head rotations in the simulations had a sampling rate of only 500 Hz, where CFC 60 is more appropriate. All results were analysed with both CFC 60 and CFC 180. Virtually no difference was found, apart from the contact force which appeared to be noisier than the other results. For consistency, all data in this study have been filtered with CFC

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