Finite Element Analysis of Passenger Multiple Belt Restraint Configurations

Finite Element Analysis of Passenger Multiple Belt Restraint Conﬁgurations by Mao Yu A thesis submitted to the Graduate Faculty of WAKE FOREST UNIVERSITY SCHOOL OF BIOMEDICAL ENGINEERING & SCIENCES

In Partial Fulﬁllment of the Requirements

for the Degree of

MASTER OF SCIENCE

Biomedical Engineering

August 2010

Winston-Salem, North Carolina

Approved by: Joel D. Stitzel, PhD, Advisor, Chair

Examining Committee: H. Clay Gabler, PhD

Warren N. Hardy, PhD Acknowledgements

Thanks to my advisor Dr. Stitzel and my committee for their patience in me and my lab mates for their assistance.

i Contents

Acknowledgements i

Contents ii

List of Figures v

List of Tables x

Acronyms xi

Physical Constants xii

Symbols xiii

Abstract xiv

1 Introduction and Objectives 1 1.1 Motivation ...... 1 1.2 Human volunteers, PMHS, animals, and ATDs ...... 2 1.3 Computer Simulations ...... 3 1.4 Injury Measures ...... 4 1.5 Injury Risk Functions ...... 5 1.6 Safety Equipment ...... 6 1.7 Study Objective ...... 7

2 Finite Element Simulation of Frontal Crash using THUMS 8 2.1 Introduction ...... 8 2.2 Model Structure and Development ...... 10 2.2.1 Vehicle Interior ...... 10 2.2.2 THUMS model ...... 11 2.2.3 Restraints ...... 12 2.2.4 Boundary Condition ...... 14 2.2.5 Contact Deﬁnition ...... 15 2.3 Model Review ...... 18

ii Contents iii

2.3.1 Static Checks ...... 18 2.3.2 Preliminary Runs ...... 19 2.4 Results ...... 22 2.4.1 Simulation snapshots ...... 22 2.4.2 Overall fringe plots ...... 24 2.4.3 Skeletal fringe plots ...... 26 2.5 Discussion ...... 28 2.5.1 Model development ...... 28 2.5.2 Post Mortem Human Subject (PMHS) validation ...... 30 2.5.3 Simulation results ...... 35 2.6 Conclusion ...... 38

3 Development of Finite Element Injury Measures for THUMS 39 3.1 Introduction ...... 39 3.2 Methods ...... 41 3.2.1 Head Measure ...... 42 3.2.2 Neck Measure ...... 46 3.2.3 Chest Measures ...... 49 3.2.4 Abdominal Measures ...... 56 3.2.5 Vehicle Interior Measures ...... 58 3.2.6 Test scenario ...... 59 3.3 Results ...... 59 3.3.1 Head ...... 60 3.3.2 Neck ...... 72 3.3.3 Chest ...... 86 3.3.4 Abdomen ...... 101 3.3.5 Derived measures ...... 111 3.4 Discussion ...... 112 3.4.1 Results ...... 112 3.4.2 Injury measures ...... 114 3.5 Conclusion ...... 118

4 Injury Risk Functions for THUMS 119 4.1 Introduction ...... 119 4.2 Methods ...... 124 4.3 Results ...... 131 4.4 Discussion ...... 137 4.5 Conclusion ...... 140

5 Comparison of diﬀerent load limits using Total HUman Model for Safety (THUMS) 141 5.1 Introduction ...... 141 5.2 Methods ...... 143 5.3 Results ...... 149 5.3.1 Simulation snapshots ...... 149 Contents iv

5.3.2 Head injury measures ...... 153 5.3.3 Neck injury measures ...... 155 5.3.4 Chest injury measures ...... 157 5.3.5 Pelvis injury measures ...... 162 5.3.6 Risks ...... 168 5.4 Discussion ...... 174 5.5 Conclusion ...... 175

6 Comparison of diﬀerent pretensioner placement using THUMS 176 6.1 Introduction ...... 176 6.2 Methods ...... 178 6.3 Results ...... 182 6.3.1 Simulation snapshots ...... 182 6.3.2 Head injury measures ...... 188 6.3.3 Neck injury measures ...... 191 6.3.4 Chest injury measures ...... 193 6.3.5 Pelvis injury measures ...... 196 6.3.6 Risks ...... 200 6.4 Discussion ...... 206 6.5 Conclusion ...... 208

7 Comparison of four-point and three-point restraint using THUMS 209 7.1 Introduction ...... 209 7.2 Methods ...... 210 7.3 Results ...... 214 7.3.1 Simulation snapshots ...... 214 7.3.2 Head injury measures ...... 221 7.3.3 Neck injury measures ...... 224 7.3.4 Chest injury measures ...... 226 7.3.5 Pelvis injury measures ...... 231 7.3.6 Risks ...... 237 7.4 Discussion ...... 243 7.5 Conclusion ...... 245

Bibliography 246 List of Figures

2.1 Vehicle Interior Components ...... 10 2.2 FE surrogate models ...... 12 2.3 THUMS chest cavity ...... 13 2.4 Three-point restraint ...... 14 2.5 Crash Pulse ...... 15 2.6 Crash Pulse (Velocity) ...... 15 2.7 Contact deﬁnitions ...... 17 2.8 Viscus properties ...... 19 2.9 THUMS simulation snapshot ...... 23 2.10 Hybrid III (HIII) simulation snapshot ...... 23 2.11 THUMS fringe plot ...... 24 2.12 HIII fringe plot ...... 25 2.13 THUMS fringe plot skeleton ...... 26 2.14 HIII fringe plot skeleton ...... 27 2.15 PMHS diagram ...... 31 2.16 PMHS test pulse ...... 31 2.17 PMHS validation ...... 32 2.18 PMHS displacements ...... 32 2.19 PMHS head movement ...... 33 2.20 Validation z - displacement ...... 33 2.21 PMHS passenger side kinematics ...... 34

3.1 THUMS injury measure ...... 41 3.2 THUMS head ...... 44 3.3 HIC node placement ...... 45 3.4 Nij illustration ...... 47 3.5 THUMS neck springs and dampers ...... 49 3.6 Horizontal rib deflection ...... 51 3.7 Parallel rib deflection ...... 52 3.8 CTI illustration ...... 53 3.9 THUMS chest components ...... 54 3.10 Modified Rib Cortical ...... 55 3.11 HIII vs. THUMS rib differences ...... 56 3.12 THUMS abdominal organs ...... 57 3.13 Lap belt force measure ...... 58

v List of Figures vi

3.14 Iliac crest force measure ...... 58 3.15 Test three-point case ...... 60 3.16 THUMS head acceleration ...... 62 3.17 HIII head acceleration ...... 63 3.18 THUMS HIC15 ...... 65 3.19 HIII HIC15 ...... 67 3.20 THUMS HIC36 ...... 69 3.21 HIII HIC36 ...... 71 3.22 THUMS neck normal force ...... 74 3.23 HIII neck normal force ...... 75 3.24 THUMS neck shear force ...... 77 3.25 HIII neck shear force ...... 78 3.26 THUMS neck bending moment ...... 79 3.27 HIII neck bending moment ...... 81 3.28 THUMS neck lateral moment ...... 82 3.29 THUMS Nij ...... 84 3.30 HIII Nij ...... 85 3.31 THUMS chest resultant acceleration ...... 88 3.32 HIII chest resultant acceleration ...... 90 3.33 THUMS chest sternal deflection ...... 91 3.34 HIII chest deflection ...... 92 3.35 Normalized sternal deflection ...... 93 3.36 Horizontal left rib deflections ...... 95 3.37 Parallel left rib deflection ...... 97 3.38 Forces on shoulder ...... 98 3.39 Chest rib fractures ...... 100 3.40 Abdominal acceleration ...... 103 3.41 THUMS pelvic acceleration ...... 105 3.42 THUMS resultant pelvic acceleration ...... 106 3.43 HIII resultant pelvic acceleration ...... 107 3.44 Abdominal lap belt force ...... 109 3.45 Left iliac crest force ...... 110

4.1 Injury risk function for HIC36 ...... 125 4.2 Injury risk function for HIC15 ...... 126 4.3 Injury risk function for Nij ...... 127 4.4 Injury risk function for chest 3ms clip ...... 128 4.5 Injury risk function for chest sternal deﬂection ...... 129 4.6 Injury risk function for Combined Thoracic Index (CTI) ...... 130 4.7 THUMS Head Injury Criteria (HIC)15 risk of injury ...... 131 4.8 HIII HI! (HI!)15 risk of injury ...... 131 4.9 THUMS HIC36 risk of injury ...... 132 4.10 HIII HIC36 risk of injury ...... 132 4.11 THUMS Neck Injury Criterion (Nij) risk of injury ...... 133 List of Figures vii

4.12 HIII Nij risk of injury ...... 133 4.13 THUMS chest 3ms clip risk of injury ...... 134 4.14 HIII chest 3 ms clip risk of injury ...... 134 4.15 THUMS chest deﬂection risk of injury ...... 135 4.16 HIII chest deﬂection risk of injury ...... 135 4.17 THUMS CTI risk of injury ...... 136 4.18 HIII CTI risk of injury ...... 136

5.1 Model setup and belt configurations ...... 144 5.2 Seatbelt properties ...... 145 5.3 Three-point retractor properties ...... 147 5.4 Four-point retractor properties ...... 148 5.5 Crash pulse applied ...... 148 5.6 Three-point load limiter 1 ...... 149 5.7 Three-point load limiter 2 ...... 150 5.8 Three-point no load limiter ...... 150 5.9 Four-point load limiter 1 ...... 151 5.10 Four-point load limiter 2 ...... 152 5.11 Four-point no load limiter ...... 152 5.12 Maximum head acceleration ...... 153 5.13 HIC36 ...... 154 5.14 HIC15 ...... 154 5.15 Neck tension ...... 155 5.16 Nij ...... 156 5.17 Chest 3ms clip ...... 157 5.18 Chest deflection ...... 158 5.19 Chest CTI ...... 159 5.20 Chest rib stress ...... 160 5.21 Chest lumbar force ...... 161 5.22 Abdominal acceleration ...... 162 5.23 Pelvic acceleration ...... 163 5.24 Abdominal maximum Viscous Criterion (VC) ...... 164 5.25 Abdominal minimum VC ...... 165 5.26 Abdominal lap belt force ...... 166 5.27 Abdominal left iliac crest force ...... 167 5.28 HIC36 risks ...... 168 5.29 HIC15 risks ...... 169 5.30 Nij risks ...... 170 5.31 Chest 3ms clip risks ...... 171 5.32 Chest deflection risks ...... 172 5.33 CTI risks ...... 173

6.1 Crash pulse applied ...... 178 6.2 Pretensioner setup ...... 179 List of Figures viii

6.3 Pretensioner pull-in ...... 180 6.4 Both pretensioners load limiter 1 ...... 182 6.5 Buckle pretensioner load limiter 1 ...... 183 6.6 Retractor pretensioner load limiter 1 ...... 183 6.7 Both pretensioners load limiter 2 ...... 184 6.8 Buckle pretensioner load limiter 2 ...... 184 6.9 Retractor pretensioner load limiter 2 ...... 185 6.10 Both pretensioners no load limiter ...... 186 6.11 Buckle pretensioner no load limiter ...... 186 6.12 Retractor pretenseioner no load limiter ...... 187 6.13 Maximum head acceleration ...... 188 6.14 HIC36 ...... 189 6.15 HIC15 ...... 190 6.16 Neck tension ...... 191 6.17 Nij ...... 192 6.18 Chest 3ms clip ...... 193 6.19 Chest deﬂection ...... 194 6.20 Chest CTI ...... 194 6.21 Chest rib stress ...... 195 6.22 Chest lumbar force ...... 195 6.23 Abdominal acceleration ...... 196 6.24 Pelvic acceleration ...... 197 6.25 Abdominal maximum VC ...... 197 6.26 Abdominal minimum VC ...... 198 6.27 Abdominal lap belt force ...... 198 6.28 Abdominal left iliac crest force ...... 199 6.29 HIC36 risks ...... 200 6.30 HIC15 risks ...... 201 6.31 Nij risks ...... 202 6.32 Chest 3ms clip risks ...... 203 6.33 Chest deﬂection risks ...... 204 6.34 CTI risks ...... 205

7.1 Model setup and belt conﬁgurations ...... 211 7.2 Belt force vs strain ...... 211 7.3 Four-point restraint ...... 212 7.4 Crash pulse applied ...... 213 7.5 Three-point both pretensioners load limiter 1 ...... 214 7.6 Three-point buckle pretensioner load limiter 1 ...... 215 7.7 Three-point retractor pretensioner load limiter 1 ...... 215 7.8 Three-point both pretensioners load limiter 2 ...... 216 7.9 Three-point buckle pretensioner load limiter 2 ...... 216 7.10 Three-point retractor pretensioner load limiter 2 ...... 217 7.11 Three-point both pretensioners no load limiter ...... 217 List of Figures ix

7.12 Three-point buckle pretensioner no load limiter ...... 218 7.13 Three-point retractor pretenseioner no load limiter ...... 218 7.14 Four-point retractor pretensioner load limiter 1 ...... 219 7.15 Four-point retractor pretensioner load limiter 2 ...... 219 7.16 Four-point retractor pretensioner no load limiter ...... 220 7.17 Maximum head acceleration ...... 221 7.18 HIC36 ...... 222 7.19 HIC15 ...... 223 7.20 Neck tension ...... 224 7.21 Nij ...... 225 7.22 Chest 3ms clip ...... 226 7.23 Chest deﬂection ...... 227 7.24 Chest CTI ...... 228 7.25 Chest rib stress ...... 229 7.26 Chest lumbar force ...... 230 7.27 Abdominal acceleration ...... 231 7.28 Pelvic acceleration ...... 232 7.29 Abdominal maximum VC ...... 233 7.30 Abdominal minimum VC ...... 234 7.31 Abdominal lap belt force ...... 235 7.32 Abdominal left iliac crest force ...... 236 7.33 HIC36 risks ...... 237 7.34 HIC15 risks ...... 238 7.35 Nij risks ...... 239 7.36 Chest 3ms clip risks ...... 240 7.37 Chest deﬂection risks ...... 241 7.38 CTI risks ...... 242 List of Tables

3.1 Parts of THUMS head ...... 45 3.2 Nij limits ...... 46 3.3 THUMS neck properties ...... 48 3.4 CTI limits ...... 53 3.5 THUMS thoracic materials ...... 56 3.6 THUMS abdominal parts ...... 57 3.7 Derived injury measures ...... 111

4.1 Select Injury Assessment Reference Value (IARV) values ...... 122 4.2 AIS code description ...... 124 4.3 HIC15 injury risk ...... 131 4.4 HIC36 injury risk ...... 132 4.5 Nij injury risk ...... 133 4.6 Chest 3ms clip injury risk ...... 134 4.7 Chest deﬂection injury risk ...... 135 4.8 CTI injury risk ...... 136

5.1 Restraint activation sequence ...... 147 5.2 Case matrix ...... 148

6.1 Restraint activation sequence ...... 181 6.2 Case matrix ...... 181

7.1 Case matrix ...... 213

x Acronyms

AIS Abbreviated Injury Scale ...... 122 ATD Anthropomorphic Test Device...... 2 CDF Cumulative Distribution Function ...... 123 CFL Courant-Frederick-Levy ...... 30 CG Center of Gravity ...... 45 CSDM Cumulative Strain Damage Measure ...... 115 CTI Combined Thoracic Index ...... vi DDM Dilatation Damage Measure ...... 115 DV Delta-V...... 14 FE Finite Element...... 3 FMVSS Federal Motor Vehicle Safety Standard...... 50 HANS Head And Neck Support ...... 244 HIC Head Injury Criteria ...... vi HIII Hybrid III...... v IARV Injury Assessment Reference Value...... x LSTC Livermore Software Technology Corporation ...... 22 MVC Motor Vehicle Crash ...... 1 NCAP New Car Assessment Program ...... 177 NHTSA National Highway Traﬃc Safety Administration ...... 1 Nij Neck Injury Criterion ...... vi PMHS Post Mortem Human Subject ...... iii RMDM Relative Motion Damage Measure ...... 116 SI Severity Index ...... 42 SIMon Simulated Injury Monitor...... 116 SNPRM Supplemental Notice of Proposed Rule Making...... 126 THUMS Total HUman Model for Safety ...... iii TL Tolerance Limit...... 42 VC Viscous Criterion ...... vii

xi Physical Constants

Natural Number e = 2.718 28 ··· m Standard Gravity g = 9.806 65 ··· s2

Pi π = 3.141 59 ···

xii Symbols

m2 a acceleration G or s

a3ms 3 ms clip G b scale (Logistic Distribution) N k spring constant strain m location (Logistic Distribution) t time s m2 A eﬀective acceleration Gs or s D deﬂection in or mm E Young’s modulus P a

Ei initial Young’s modulus P a K bulk modulus P a T duration s

rads α angular acceleration s2 strain (mechanics) µ mean (statistics) σ stress (mechanics) P a σ standard deviation (statistics) rads ω angular velocity s

xiii Abstract

Injury measures have been developed for crash test dummies to associate the data generated to some meaningful measure to the human body. While the crash test dummies oﬀered an important surrogate for the human body in real world crash testing, in ﬁnite element modeling, models of the human body could be an improvement over FE models of crash test dummies.

In this study, a set of injury criterion are developed for the Total Human Model for Safety (THUMS). While the crash test dummies uses load cells and accelerometers to measure force and acceleration data, through THUMS, these forces and accelerations can be measured more directly. Several areas of the THUMS will be measured including head, neck, shoulder, chest, abdomen, and pelvic injury criterion. The simulations will be run in LS-Dyna 971.

To test the validity of the injury measures, simulations will be run using THUMS with a 40 mph delta-v crash pulse. The seat and floor board will be modeled with rigid body components to simplify the simulation. The knee bolster will be modeled as a simple elastic material. No other interior components will be modeled. Two types of belt configuration will be tested, a three point seat belt configuration, and a four point seatbelt configuration. In addition, three load limiter levels will be used: 1500 N load limit, 3000 N load limit, and no load limit. This will determine if the injury measures provides realistic outputs in a simulated motor vehicle crash. “All models are false but some models are useful.”

George Edward Pelham Box Chapter 1

Introduction and Objectives

The ﬁeld of Injury Biomechanics spawns from a need to protect occupants. As automobile usage have become widespread around the world, Motor Vehicle Crash (MVC)s are becoming more problematic.

1.1 Motivation

MVCs were the leading cause of death in 2005 for ages 4-34 and was ranked third in terms of years of life lost in the United States [1]. While safety devices such as seatbelts and frontal air bags have saved many lives [2, 3], more than 37,000 people were killed and around 1,711,000 people were injured in MVCs in 2007 [4].

The National Highway Traffic Safety Administration (NHTSA) has shown that the seatbelt is an effective safety apparatus in saving lives [3], there is still some controversy regarding the specific configuration of the seatbelts [5–9]. Further research into the response of the occupant to varying seatbelt parameters can enhance our understanding of this safety system and fine tune its properties to better protect the occupant. We will simulate the MVC and evaluate the human response to differing safety system parameters.

1 Chapter 1. Introduction and Objectives 2

1.2 Human volunteers, PMHS, animals, and ATDs

A crucial process of MVC simulations is a valid method to measure the human response. Analysis of the human response to crash situations are performed with human volunteers [10–21], animals [22], PMHSs [9, 23], Anthropomorphic Test Device (ATD)s [9, 24–27], and computer models [28–40]. Each method is useful in enhancing the knowledge of human response to MVC scenarios. The use of volunteers in crash research is limited and ethically restrictive. Research volunteers can not be subject to severely harmful events, which limits the amount of experimentation that can be done. Human volunteers offer the most valid response, although most human volunteer data come from earlier experiments not directly applied to MVCs [14–21, 41]. This data is useful but lacking, hence PMHS and animals have been used especially in helping to understand failure criterion for various parts of the body that can not be determined from volunteer data. The drawback of using PMHSs comes from the inability to reproduce physiological responses such as bleeding, inflammation, or muscle activity [42]. Furthermore, due to limited availability and the various conditions of the cadaver, the PMHS may not be representative of the driving population. Animal models can exhibit physiological responses, but they are structurally different from humans.

To create a uniform, mass-producible, and repeatable standard for crash analysis, ATDs were created. ATDs are used as surrogates to human subjects in crash tests and experiments and are equipped with various sensors to measure acceleration, deflection, and forces. These measurements offer insight into the mechanics of the human body, but the relationship between the measurements and injury still must be determined through other means. Although ATDs have gained widespread adoption in research and government rule-making, they still suffer from several draw backs compared to PMHS and human volunteers. First, they do not offer muscle response to the event as a human volunteer would [43]. Second, they do not suffer injury or fail as human or PMHS would. Third, limited space inside the ATD means that instrumentation is limited to taking a finite number of measurements, while human volunteers or PMHS may suffer injury at a site not instrumented. Fourth, multiple ATDs have been developed for different impact scenarios, and there is no single ATD that performs best in all scenarios [24, 25]. Chapter 1. Introduction and Objectives 3

Fifth, the measured values are gross motion measurements and do not reﬂect tissue level behavior.

Of these issues, adding muscle response would require extra hardware to be implemented inside the dummy, which could be technically challenging. Realistic injury and failure is partially addressed by creating failure criterion for measurements and validating the criterion with PMHS, but this does not address behavior of the dummy after the failure criterion has been exceeded. Limited space in the ATD is addressed by instrumenting the most life-threatening injury sites, and the argument that sites where no instrumentation exist are less likely to cause life-threatening injury. The existence of multiple ATDs shows that current ATDs are specialized towards one MVC scenario and still do not capture the behavior of humans at an adequate level to completely replace PMHS and volunteer testing. No human surrogate may ever be adequate enough to replace real humans, but it might be feasible to create a uniﬁed human surrogate that behaves well in most MVC scenarios. Lack of tissue level behavior raises an inherent limitation of ATDs that cannot be addressed with the current technology. Because of these issues, ATDs are supplemental to injury biomechanics research, and are constantly validated with PMHS or human volunteer data.

1.3 Computer Simulations

To address some of the limitations of PMHS, animal testing, and ATDs, computer simulations, especially the Finite Element (FE) method, are an alternative or supplement to ATDs. Although early FE Models were reproductions of ATDs [44, 45], advancements in processing power have allowed more sophisticated human body models to be used. The THUMS and Ford full body ﬁnite element model [46, 47] are Finite Element Models that attempt to reproduce the human body in many aspects. Models such as THUMS can be used in place of ATDs as they are validated against PMHS data [29, 30, 46, 48, 49].

FE models offer several benefits compared to ATDs, PMHS, and animals. First, FE models are versatile, they can be changed and modified easily, which allows running models with many different MVC scenarios feasible. Second, FE models can be more Chapter 1. Introduction and Objectives 4 detailed than ATDs. THUMS contains the full rib cage with cortical and trabecular bones, which makes it possible to distinguish between failure of the cortical and failure of the trabecular. THUMS also has multiple layers of soft tissue and skin that are anatomically realistic. Third, FE models can also simulate muscle force through spring/muscle elements. Fourth, material properties for computational models can be more complex as mathematical models of human tissue becomes more realistic. Fifth, FE models can output more information than traditional ATDs. Data such as stress, strain, force, displacement, and kinematic data can be measured in virtually any location on the model. There are several downsides to computational models, however. First, since computational models are mathematical models, they can sometimes become numerically unstable, which can cause the model to behave unexpectedly. Second, sophisticated computational models require in depth knowledge of many fields as well as extensive testing to create a realistic set of parameters that are mathematically feasible to implement. Third, FE models must be validated against PMHS or human volunteers. Validation can be approached from many aspects, from gross motion to stresses and strains at specific locations of the body. Therefore FE models can supplement ATDs in a variety of situations to enhance understanding of the human body in MVCs.

1.4 Injury Measures

With ATDs and FE models, measured values such as forces, stresses, and displacement needs to be correlated to injury. Since the amount of instrumentation was limited for ATDs, early research focused on areas of life-threatening injuries such as head and chest. NHTSA has established limits for these measured values as the absolute critical that should not be exceeded [50]. The establishment of these limits is based on previous research and that certain measured values can be correlated to a chance of injury [51]. The decision on which values to measure was the subject of some controversy, values such as the spinal acceleration and head accelerations were used initially [52, 53], but later changed. Measurements such as the Gadd severity index served as intermediate injury measures for head injury until the development of HIC, which is still controversial. For example, HIC only considers translation acceleration. Chapter 1. Introduction and Objectives 5

However controversial, NHTSA has established some injury measures that are used in government regulations [54]. While these measures do no necessarily represent an exhaustive list, they are well researched measures. The HIC is the accepted measure for the head, with a limit of 700 for HIC15. The neck criterion used is the Nij, which measures a composite of the forces and moments experienced at the neck. The chest deﬂection and chest acceleration was used as the chest injury measures.

1.5 Injury Risk Functions

Determination of the injury measures was an important first step, but correlating these measures to risks of injury and failure was also necessary. Furthermore, a determination injury measures interaction with other measures are needed. There are two methods to correlate injury measures to injury occurrence, one is to define the absolute threshold that the injury measure could not exceed, the other is to define a probability distribution function that gives the risk of injury. The first method has the advantage of describing a universal limit that can be used as a definitive pass/fail criterion. A problem with this is that the measured limits does not necessarily mean a fail will guarantee injury if the limits are exceeded, nor does it mean injury will not occur if below the limit. This method might offer too much simplification in some instances. The second method has an advantage of giving a probability of injury, which is a more realistic criterion. The probability of injury comes from either assuming one of several data distributions, or using non-parametric techniques [55]. The data distributions commonly used are normal, lognormal, weibull, and logistic. While it is useful mathematically to assume the distribution of the data using parametric techniques, it has been shown that non-parametric methods perform better [56]. Nevertheless, parametric risk functions, especially the logistic distribution, are widely used due to their close form solutions. Although useful, there are sources of inaccuracies from oversimplification, noise, and measurement errors. Chapter 1. Introduction and Objectives 6

1.6 Safety Equipment

The goal of injury measurement and injury risk functions is to improve the outcome of MVCs through the assessment of crash safety devices. There is a division of crash safety into crash prevention and crash worthiness. Crash prevention technologies utilizes safety systems to help decrease the likelihood of getting into a MVC. On the other hand, crash worthiness is the ability of a vehicle to protect occupants in an MVC. This study focuses on the crash worthiness aspect of an MVC, which can be attributed to several components of the vehicle, including the general structure and geometry of the vehicle and the interior occupant safety systems. The crumple zone of a car absorbs a large amount of crash energy and reduces the load on the occupant safety systems. Vehicles with shorter front ends usually endure more loads in the seatbelts and airbags. Occupant safety equipment also improve the outcome of MVCs. Seatbelts are one of the oldest but most eﬀective safety technologies in the vehicle, but there are still improvements that could be made to them.

Three enhancements to the safety system that this study concentrates on are pretensioners, load limiters, and four-point belts. The pretensioner addresses an issue of extra slack in the seatbelt. In regular driving, the seatbelt is often loosely coupled to the occupant’s body, for comfort reasons. In the event of an MVC, it is optimal to pretension the seatbelt to the occupant’s body before the crash to reduce the maximum acceleration of the occupant and decrease the velocity debt that the occupant incurs from being decoupled from the vehicle, or a pretensioner. This is done diﬀerently depending on the implementation and the manufacturer; however, it is similar to the detection system used by airbags. The pretensioner can be placed at either the retractor, which loops through the shoulder D-ring, or the buckle, the eﬀect of placement will be studied.

In addition to pretensioning, the load limiter addresses another issue with seatbelts. Seatbelts are relatively narrow pieces of fabric that puts a concentrated stress across the occupant thorax. Although designed to be positioned across the strongest parts of the thorax, signiﬁcant injury can still be caused by the seatbelts as a result of a severe crash, more so for older occupants, who may have weaker bone structures due to osteoporosis or other factors. Thus seatbelts could be more harmful than helpful in a severe crash. Chapter 1. Introduction and Objectives 7

One solution is to incorporate devices to limit the maximum load on the occupant’s thorax, or load limiters. Load limiter implementation are divided into an extra loop of fabric in the seatbelt that tears apart when a load threshold is reached, or a torsion bar in the retractor that yields at a load threshold. Either way, the behavior of the load limiter is similar.

Besides add-ons to the seatbelt, there are proposed geometrical changes to the seatbelt such as a four-point seatbelt that restrains both shoulders. One of the perceived beneﬁt of this belt system is to reduce the likelihood of chest injury from severe collisions. Preliminary tests have shown both promising results and problems.

1.7 Study Objective

This study uses THUMS, a FE model of the human body, to study the eﬀects of pretensioners, load limiters, and four-point belt systems. Pretensioners will vary by placement between retractor and buckle. Load limiters will vary by load limit. Seatbelt geometry will vary by three-point to four-point systems. Diﬀering combinations of these safety systems will be simulated. Chapter 2

Finite Element Simulation of Frontal Crash using THUMS

2.1 Introduction

Development of occupant crash protection initially started with human volunteers, animals, and PMHS. While each of these methods shed insight into the behavior of the human body, each has drawbacks. Human volunteers can only be subjected to relatively safe levels of load, animals have differing geometry than humans, and PMHS suffer from degradation and lack of muscle response. Computer modeling of the human body through the FE method can overcome some of the drawbacks of these methods. Computer models can be subjected to many load scenarios with less effort and cost compared to using physical surrogates. They can be created to be detailed as seen by THUMS. The THUMS model offer a sophisticated and validated model for use as a human surrogate in a FE simulation [29, 30, 34–36, 46, 48, 49, 57].

FE simulations with full body FE models are often performed with pre-made models such as THUMS which is a general purpose FE crash model. General purpose FE models are put through a thorough development cycle that requires details on most important aspects of the model’s body to match well with that of a human counterpart. This matching is often done with experimental data as well as previous research. The

8 Chapter 2. Finite Element Simulation of Frontal Crash using THUMS 9 advantages of using a well-designed pre-made FE models are a uniform model can be compared across different studies, the model is often more validated if developed separately, and the model is designed to withstand a variety of different scenarios and still behave realistically. Disadvantages of using a pre-made FE models include difficulty in incorporating the FE model into the current design, problems debugging model errors due to lack of understanding of the model, and less flexibility in modification of the model.

Despite the drawbacks of using a pre-made model, developing a full body FE model is also not feasible for most studies as it requires a heavy investment of time and resources, therefore the pre-made THUMS will be used in this study. A FE simulation using THUMS was developed and the experiences and results are presented. The simulation was performed for a vehicle interior with standard three-point belt. Chapter 2. Finite Element Simulation of Frontal Crash using THUMS 10

2.2 Model Structure and Development

The FE simulation development was divided into five phases. First, a minimal vehicle interior was modeled. Second, the THUMS model was positioned in the vehicle interior. Third, restraints were fitted onto THUMS. Fourth, boundary conditions and the crash pulse was applied to the vehicle interior. Finally, contacts were defined between THUMS and the vehicle interior.

2.2.1 Vehicle Interior

The vehicle interior consists of the seat, the knee bolster, and the ﬂoor pan. No steering wheel or airbag was included in this study. Center console or doors were not modeled since this was a frontal impact simulation. The interior components are shown in Figure 2.1.

Head Rest

Seat Back

Knee Bolster Seat Bottom

Floor Pan Pedals

Figure 2.1: Vehicle Interior Components

The THUMS back had a finer element mesh while the THUMS bottom and legs had a coarser element mesh. This could be to allow more accurate information from the chest Chapter 2. Finite Element Simulation of Frontal Crash using THUMS 11 and spine. As a result, the seat back was created with a finer mesh than the seat bottom to provide better interaction with THUMS. During a frontal impact, the occupant is thrown forward due to the pulse and minimal contact forces exist between the occupant and the seat back. The rest of the seat was meshed coarsely as detailed information of the seat was not desired in this study. Furthermore, the rigid material in LS-Dyna was used to model all of the seat and the floor pan. The rigid material is not a reflection of the material property as much as the element processing [31]. It is a cost-effective way of modeling support structures in a simulation where stress and deformation information is not necessary. The rigid material still has properties for contact treatment. The seat in this study has a Young’s Modulus of 52.5 kPa and a Poisson’s ratio of 0.33 (approximately the material properties for foam). This ensures that the seat is not too hard for contact with the THUMS. The knee bolsters are modeled with an elastic material with a Young’s Modulus of 10 MP. All the parts are constrained to the seat back and held in place via global constraints.

2.2.2 THUMS model

The THUMS model was placed in the seat and moved until no initial penetration existed between the THUMS and the seat. THUMS itself was designed with initial penetration between various segments of the limbs so that it could be placed in diﬀerent postures by simply rotating and translating the limbs. This means that there are initial penetrations between the THUMS and some of the limbs, but this is not problematic in this cases. As long as no contacts are deﬁned between THUMS and its own limbs, initial penetrations are acceptable if unrealistic. The THUMS model represents a 50% male with height of 175 cm, weight of 77 kg, and material properties that represents a 30-40 year old [58]. The model itself has been validated in multiple directions of loading with multiple body parts [29, 30, 33–35]. The model is a more detailed representation of the human body compared to ATD based FE models such as the HIII as seen in Figure 2.2. The THUMS uses approximately 91 thousand elements and 66 thousand nodes to simulate the human body. Chapter 2. Finite Element Simulation of Frontal Crash using THUMS 12

Hybrid III Total HUman Rigid Model Model for Safety 50% Male (THUMS) 50% Male

Figure 2.2: FE surrogate models: HIII vs. THUMS

THUMS has a biofidelic skeletal structure and outer geometry. The THUMS body come from anatomically correct geometry data that include bones, skin, and joints. Tendons and ligaments are also closely modeled at each joint. No idealized joints (CON- STRAINED JOINT) keywords are used. Instead, all joints are modeled with contact definitions. The THUMS model uses shell elements to model the cortical bones while using solid elements to model the trabecular. The THUMS skin is also composed of a shell outer layer and a solid inner layer. While THUMS attempts to accurately model the skeletal structure of the human body, some soft tissue information such as internal organs and facial details are not modeled accurately. This lack of biofidelity in THUMS should be noted when analyzing organ level injury measures. The lack of organs in THUMS is compensated for by the use of solid bodies inside the chest cavity to fill the empty volume. This is not an ideal solution, but it reduces the run time while still maintaining some biofidelity for gross behavior.

2.2.3 Restraints

A standard three-point restraint with shoulder retractor and pretensioner was modeled. Belt ﬁtting was performed with LS-PREPOST (Livermore Software Technology Corpo- ration, Livermore, CA). A mixture of 1D seatbelt and 2D shell elements with thickness were used rather than pure 1D seatbelt elements. The 2D shell elements had a thickness of 1 mm and a width of approximately 47 mm. The seatbelt was oﬀset approximately 2mm from the THUMS body and contains triangular elements where it interacts with Chapter 2. Finite Element Simulation of Frontal Crash using THUMS 13

THUMS Chest Cavity

Parts

Figure 2.3: THUMS chest cavity: the lack of organs are shown with lines

the body, and seatbelt elements where it feeds into the retractor and anchor. A pure seatbelt element formulation would have required a node to surface contact with a 1- dimensional element seatbelt, which would is less realistic. The ideal scenario would be to have only shell-type seatbelt elements, but this element is in the experimental phase and was not deemed stable enough for this simulation.

The seatbelt elements use a linear elastic material with a Young’s Modulus of 20 GPa, such that there is minimal elongation in the shell elements. The seatbelt elements do not have an elastic modulus, but are controlled by loading and unloading curves that deﬁne their behavior (force vs. engineering strain). The belt loads at approximately 9.39 × 106 N per change in length. The density of the belt was assumed to be 1080 kg/m3 and the this was converted to mass per unit length by assuming the cross sectional area of the shell elements, computed as thickness times width.

The three-point belt is shown in Figure 2.4. The three-point belt contains a series of 1D belt elements that contains two shell element segments to cover the thorax and lap. The belt can freely slide though the buckle and the shoulder ring. The retractor allows belt pull-out before locking. The three-point belt configuration also allows pyrotechnic pretensioner placement at the retractor or the buckle. Because a pretensioner was modeled at the buckle, the buckle connector was given a stiffness of 800 × 106N/m to Chapter 2. Finite Element Simulation of Frontal Crash using THUMS 14 simulate a rigid bar which could be pulled down. The rest of the buckle shared similar material properties as that of the seat, with rigid body definitions.

Slipring

Shell Elements Seatbelt Elements

Retractor

Buckle

Figure 2.4: Three-point restraint with seat (isometric view)

2.2.4 Boundary Condition

There are some boundary conditions that must be applied to the model. These can be separated into the global constraints and applied kinematics. The constraints applied include chaining the interior structure of the vehicle so that they do not move relative to each other. This was be done by tying each part of the interior structure (seat, knee bolster, ﬂoor pan, retractor, buckle, belt anchor) to the seat back through either rigid body deﬁnitions or nodal constraints (CONSTRAINED RIGID BODIES and CONSTRAINED EXTRA NODES SET).

The applied kinematics include the crash pulse and gravity. The crash pulse applied is a 40 mph Delta-V (DV) pulse that comes from a severe crash. The pulse is shown in Figure 2.5, and the integrated velocity vs. time curve is shown in Figure 2.6. The crash pulse applied drops down dramatically after 200 ms, which means that the simulation does not need to run much longer than 200 ms. The acceleration is applied to the vehicle interior via the BOUNDARY PRESCRIBED MOTION keyword, which Chapter 2. Finite Element Simulation of Frontal Crash using THUMS 15 applies a kinematic curve as a function of time to the part. No initial velocity is applied. In the end, the vehicle interior will have a 40 mph velocity. Standard gravity is applied via a LOAD BODY Z keyword, which loads every part in the simulation by a constant.

Applied Crash Pulse

300 ) 2 s m

( 200

100 acceleration 0

0 0.05 0.1 0.15 0.2 0.25 time (s)

Figure 2.5: Crash pulse: acceleration of the seat vs. time

Applied Crash Pulse

15 ) s m ( 10 y cit elo v 5

0 0 0.05 0.1 0.15 0.2 0.25 time (s)

Figure 2.6: Crash pulse: velocity of the seat vs. time. The maximum velocity is m approximately 17.8 s or 40mph.

2.2.5 Contact Deﬁnition

After the creation of geometry and application of boundary conditions, the interaction between parts, or the contact definitions, in the simulation must be defined. Contact Chapter 2. Finite Element Simulation of Frontal Crash using THUMS 16 between parts, such as colliding, sliding, and coupling, must be defined for the simulation to be realistic, otherwise the parts would fall through each other. For collision type contacts, the contacts used are surface to surface contacts. The surface to surface contact definition uses a penalty based formulation in which the contacting surfaces are given a force normal to the surface when penetration occurs that increases as the amount of penetration increases. There are also the option of using nodes to surface, or part to part contact. The automatic surface to surface contacts are stable contacts that ensure the correct normal vectors. In addition, surface to surface contact definitions are more thorough than node to surface contact definitions. For the default contact between nodes and surface, there is a master surface and slave nodes. The slave node is checked for penetration into the master surface at each time point. In a default surface to surface contact, the nodes of one surface is assigned to be the slave nodes and check against the other, then the nodes of the other surface is assigned to be the slave nodes and checked also. Thus in surface to surface contact definitions, it is irrelevant which surface is the master and which is the slave.

Three considerations when implementing the contact surfaces include the mesh density of the contact surface, the contact stiffness between the contacting surfaces, and bucket sorting. The mesh density of the contact surface directly affects the distribution of force on the contacting nodes. With more nodes on the contacting surface, the contact force is more evenly distributed, and less likely to cause instability in the model. Dissimilar or badly formed meshes can also cause problems with surface to surface contacts as the distinction of which slave node belongs to which master node maybe difficult to determine, and this may cause difficulties in determining initial penetrations between contacting surfaces. Contact between surfaces of similar stiffnesses are well handled by the default contact algorithm. However, for contacts between two dissimilar materials such as the THUMS soft tissue and the floor pan, additional measures may need to be taken to prevent excessive penetration due to overly soft contact stiffnesses. The soft constraint penalty formulation (SOFT = 1 or 2) can be used to handle these types of contact situations. This formulation modifies the contact stiffness so that it is a function of the time step and a scale factor. Another option is to manually scale the default contact stiffness. There are a few differences between SOFT = 1 and Chapter 2. Finite Element Simulation of Frontal Crash using THUMS 17

SOFT = 2 in terms of formulation, but SOFT = 2 produced simulations that were less likely to become unstable for this study. The contact bucket sorting attempts to optimize contact algorithm so that every slave node is not compared against every master segment. Instead, nodes are grouped and only nearest groups are compared. Since bucket sorting is an expensive part of the contact algorithm, increasing the number of cycle between bucket sorts can dramatically increase the simulation speed. However, for simulations where there maybe large motions between time steps, the bucket sort may need to be done every cycle. There are also some additional considerations such as edge to edge contact and warped segment checking, which are useful when using irregularly shaped meshes. Sometimes it is necessary to increase the size of the segments via the MAXP AR parameter for irregular shapes and sharp corners [31, 59].

Contact deﬁnitions were implemented in several discrete areas. The contacts deﬁned are shown in Figure 2.7.

Head to headrest

Body to Belt

Knee to Bolster

THUMS back to Seatback Feet to Floor pan

Figure 2.7: Contact deﬁnitions between THUMS and vehicle interior Chapter 2. Finite Element Simulation of Frontal Crash using THUMS 18

2.3 Model Review

Once the design of the model is complete, a review of the model for computational issues, or model stability, should be addressed. Several issues should be addressed before beginning simulation runs, while other issues do not become apparent until some preliminary runs have been performed. Model stability issues that are not resolved may lead to unrealistic behavior, negative volumes, and shooting node errors.

2.3.1 Static Checks

Initial penetrations between contacting surfaces are an issue where one of the contacting surface begins the simulation with penetration into the other surface. This can cause large initial forces, numerical instability, or breakdown of the contact algorithm. Initial penetrations between non-contacting surfaces are, however, not a problem, but defining a contact might be considered for the sake of realism. Initial penetration can be check with several methods. First, visual inspection of the simulation set-up should be performed first and last to ensure that there are no gross penetrations. Attention must be paid to shell thickness when shell to shell contacts are defined. In this study, shell to shell contact was not used, instead, a separate segment set was defined for each contact surface. The segment set definition is based on defining corner nodes for the segments, so that shell thickness do not come into play. The segment set definition also has the advantage of allowing defining a subset of a part as the contacting surface, which saves simulation time. Second, initial penetration can also be checked via third-party software such as LS-PREPOST and Altair Hypermesh (Troy, MI). Finally, run time checks can also be performed for initial penetrations. LS-DYNA outputs warning messages for initial penetrations and specifies the penetration distance. Large initial forces and contact energies are usually a sign of initial penetrations (this is best checked without applied loads).

Element quality is another important static check that should be performed before any simulations. Element quality check for the vehicle interior was relatively simple since most of the parts were not complex. The varying degree of complexity for the THUMS Chapter 2. Finite Element Simulation of Frontal Crash using THUMS 19

THUM viscus properties

40 ) a P

k 30 (

20 stress

0 0 0.2 0.4 0.6 0.8 1 1.2 strain (unitless)

Figure 2.8: THUMS modiﬁed viscus properties

model itself make element quality checking diﬃcult. Due to the large number of elements for THUMS, most of the parts are coarsely meshed to optimize simulation speed. In addition, high warp angles and sharp corners are sometimes unavoidable in some parts such as near joints. The parts themselves sometimes have diﬀerent sized elements. Since anatomical data can be complex, poor element quality are sometimes necessary. In this study, poor element quality for THUMS are only corrected once the preliminary runs deemed it necessary.

2.3.2 Preliminary Runs

Preliminary runs show that there are a few problem areas in THUMS. One persistent problem area was the THUMS viscus (the lung-like part) that often ended a run in negative volume errors. Negative volume errors are caused by overly high deformation in elements that experience excessive forces. The problem could come from the material property, the contact deﬁnition, element formulation, hourglassing, or the surrounding parts. In this case, multiple steps were taken to remedy the situation.

The viscus was originally designed based on the LOW DENSITY FOAM material model in LS-DYNA for the current version of THUMS. This proved to be mostly unstable, especially with high loads on THUMS. Thus the material properties for the THUMS Chapter 2. Finite Element Simulation of Frontal Crash using THUMS 20 viscus was changed to CELLULAR RUBBER, from version 1.52b of the THUMS version. Figure 2.8 shows the stress vs. strain curve for this material as implemented in LS-DYNA. The material itself represents a hyperelastic material with entrapped air. The hyperelastic formulation behaves similarly to a generalized form of the Mooney- Rivlin rubber model [59] and is much less compressible than the foam formulation due to a much higher poisson’s ratio. The THUMS viscus material still was shown to be problematic even with this modification so that a more robust element formulation was used. The default element formulation for the viscus was a single point formulation, which calculates the volumetric integral by approximating it using the Jacobian at the center of the element. This formulation is designed for a fast solution and can result in unrealistic behavior in extreme deformations. These unrealistic behavior are hourglass modes and can cause early termination. To remedy this, either hourglass control can be implemented or another element formulation can be substituted. There are viscous and stiffness forms for hourglass control and it seems that the stiffness form works better for higher impacts. However, hourglass control by itself was not sufficient to prevent early termination and the element formulation was also modified. Two additional element formulation can be considered instead of the one-point integration. There is the fully integrated 8-node brick, which increases the cost of stress update by a factor of 8, and the selectively reduced fully integrated brick, which assumes constant pressure throughout the element. The fully integrated and the selectively reduced element both offer increase stability under most circumstances. Fully integrated solids are more expensive than the default one-point integration, but are more accurate and do not suffer from hourglassing. One-point integrated solids use the Jacobian at the center of the element to approximate the volumetric integrals. They are faster to compute than fully integrating the solid, but can suffer from hourglassing. Hourglass modes are often controlled by using extra hourglass algorithms. The one-point integrates solids have a default hourglass formulation, but this is not always optimal. In this case, it was deemed more effective to use the fully-integrated solid to avoid hourglassing. It was also noted that the surrounding ribs often failed excessively, the failure criterion for these ribs were corrected to have a more realistic failure mode [60]. The selectively reduced element formulation was used as it offered stable simulations for the scenarios in this study. Chapter 2. Finite Element Simulation of Frontal Crash using THUMS 21

The viscus itself offered little resistance to compression compared to the ribs and was therefore sensitive to negative volumes. The modification to the ribs, the change in hourglass control, and the change in element formulation all increased the stability of the viscus, but this does not guarantee that THUMS can handle any arbitrary load. At most, the threshold for the load THUMS can handle has been increased. On the other hand, it might be that the modifications to THUMS only work for the current set of scenarios and that new modifications will need to be implemented for different scenarios. As the viscus represents important organs such as the heart and lung, modification to the material properties should be considered carefully. The current modification was based on previous version of THUMS that were validated. The modified viscous stress vs. strain curve for the material shown in Figure 2.8 demonstrates a stiffening of the material as the strain increases. This property is often necessary to ensure stable simulation. Realistically, the viscus would fail under extreme strains, but it is often computationally difficult to implement realistic failure criteria and sometimes it is not desirable. Thus the question of what to do when the strain reach extremely high values is an interesting study in itself. For the current study, materials that do not have failure criteria defined are relegated to experiencing increasingly stiffer behavior as the strain increases, which acts to stymie further stretching, so that as not to promote extreme stretching of the tissue and control negative volume behavior.

The THUMS pelvis is frequently in contact with seats and thus often experiences upward forces. The THUMS pelvis bone often causes sharp concentration of forces on the THUMS flesh internally and would result in early termination. The stress concentrations could be resolved by creating a more rounded pelvic bone and using a contact definition that would allow a more distributed force. This was proven difficult as it included dramatic changes to the pelvic model. Instead, the material property for the pelvic flesh was changed to an elastic model and hence the tissue did not use the viscoelastic model usually used by soft tissue with THUMS. The material was changed to elastic and stiffened until negative volume errors were eliminated. The pelvis flesh material was originally a viscoelastic material with a bulk modulus of 2.23 GP a. The bulk modulus was converted to an elastic modulus using linear elastic and incompressible assumptions and then gradually stiffened until the negative volume errors were eliminated. Chapter 2. Finite Element Simulation of Frontal Crash using THUMS 22

Another factor that affected the stability of the model was the coefficient of friction. A higher friction factor between THUMS resulted in a more stable model. Dynamic friction had less of an effect than the static friction.

The model was also run with the crash pulse in the side directions (near side and far side) and with the crash pulse from the rear. It was found that in most of these case, the THUMS failed in the viscus with the current modiﬁcations. Thus further modiﬁcations would be necessary to let THUMS run to completion in the other directions. These runs were a check on the stability of THUMS and are not part of the study.

After stability problems were minimized with the preliminary runs, both the THUMS and the HIII were run with the same crash pulse to compare the overall behavior of the models. The HIII model was acquired from the Livermore Software Technology Corporation (LSTC) FTP server. It has been validated against a physical HIII ATD.

2.4 Results

Three sets of simulation outputs are shown in this section. First, the gross motion are shown in a series of simulation snapshots taken at six points in time throughout the simulation. Second, a fringe plot showing maximum principal stress is shown at the most severe point of the crash pulse to show overall stress levels in the model. Finally, a fringe plot of just the skeletal structure is shown to compare the THUMS with the HIII.

2.4.1 Simulation snapshots

The gross motion of both THUMS (Figure 2.9) and HIII (Figure 2.10) are shown at times: 0 ms, 25 ms, 50 ms, 75 ms, 125 ms, and the end time. The motion of THUMS and HIII both show farther excursion from the seat back in the thorax than the abdomen, so that no submarining occurred. The THUMS head was bent more than the HIII head in the sagittal plane, which is also there was more bending in the THUMS neck and spine. The THUMS shows less forward motion in the abdomen than the HIII. Chapter 2. Finite Element Simulation of Frontal Crash using THUMS 23

Figure 2.9: Snapshot of the simulation (THUMS) at various times

Figure 2.10: Snapshot of the simulation (HIII) at various times Chapter 2. Finite Element Simulation of Frontal Crash using THUMS 24

2.4.2 Overall fringe plots

The fringe plots at time = 0.075 (75 ms) for the THUMS (Figure 2.11) and HIII (Figure 2.12) shows the maximum principal stress at the time with the most severe acceleration. The arms of the models were removed as they obstructed viewing the thorax. THUMS fringe levels were set to a range between 0 and 10 MPa, which is 100 times that of the range of HIII, set between 0 and 100 kPa. THUMS had more stress in the shoulder and neck, especially in the right shoulder where it contacts the seatbelt. HIII had more stress in the abdomen, around the lapbelt. In addition, there is more penetration of the shoulder belt into THUMS while minimal penetration is shown for HIII.

Figure 2.11: Fringe plot (maximum principal stress) of THUMS (without arms) at time after maximum acceleration (t=75 ms) Chapter 2. Finite Element Simulation of Frontal Crash using THUMS 25

Figure 2.12: Fringe plot (maximum principal stress) of HIII (without arms) at time after maximum acceleration (t=75 ms) Chapter 2. Finite Element Simulation of Frontal Crash using THUMS 26

2.4.3 Skeletal fringe plots

The skeletal fringe plots for THUMS (Figure 2.13) and HIII (Figure 2.14) are also taken at the time of maximum acceleration (t=75 ms). The skeletal fringe plots show high stresses in the ribs, which have the highest stress on the side, and the lower extremity for THUMS. The lower extremity also has stress concentrations on the femur and tibia. The HIII model has less stress information than the THUMS due to extensive use of RIGID BODY keywords. As a result, only the ribs show stresses, which are an order of a magnitude higher than THUMS.

Figure 2.13: Fringe plot (maximum principal stress) of THUMS (skeleton) at time after maximum acceleration. Note that head and the spine are rigid bodies, so that stress information was not available.

The total simulation time for THUMS is approximately 20 hours, with approximately 44% of the time used for element processing, and 55% of the time used for contact algorithms. The amount of time spent on THUMS to seat contact was minimal, which was approximately 4% of the total processing time, the rest of the contact processing time was on THUMS. Chapter 2. Finite Element Simulation of Frontal Crash using THUMS 27

Figure 2.14: Fringe plot (maximum principal stress) of HIII (skeleton) at time after maximum acceleration. Note that the rib stresses are relatively symmetrical, so two views are not provided. Chapter 2. Finite Element Simulation of Frontal Crash using THUMS 28

2.5 Discussion

This study focused on the implementation of THUMS in a frontal crash simulation and deals with some speciﬁcs of the FE modeling process with LS-DYNA. The results presented here are to show a successfully working THUMS model and compare it with the standard HIII model. Further development of result processing will be presented in later chapters.

2.5.1 Model development

The vehicle interior development consisted mostly of rigid materials. Since the seat is made from a rigid material, no deformation of the seat was modeled. The use of the rigid material is designated for supporting structure where the stress and strain measurements are not required. Ideally, rigid structures do not deﬂect much. This is a valid assumption for classical mechanics involving materials such as steel, but less valid for an automotive seat. While not accurate, the deformation of the seat was not the focus of this study and would have contributed little to the overall outcome of the simulation in terms of gross motion or overall stresses. The rigid material still oﬀered realistic interactions at the contact site, so that forces are still transmitted from the seat. The only component of the vehicle interior structure that was not rigid was the knee bolster. A linear elastic material was used for the knee bolster in order to see the interaction between the THUMS knee and the knee bolster. In addition, simulations with a rigid knee bolster were attempted where we found that the rigid bolster caused excessive deformation in the THUMS knee because it was nonyielding. This often caused stability issues with THUMS.

Compared to the HIII simulation, THUMS had more stability issues. There were often failures in the soft tissues of THUMS, such as the viscus in the chest cavity, the abdomen, and the buttocks. The viscus was a major source of stability problems. The initial viscus material was a low density foam, but it was changed to cellular rubber to maintain the hyperelastic aspect of the material and ensure the stability. The cause of these failures is also partially the 40 mph DV crash pulse that is more severe than the Chapter 2. Finite Element Simulation of Frontal Crash using THUMS 29 crash pulses THUMS was validated against (35 mph). This was tested by lowering the crash pulse by 75% and the simulations ran without failures. This points out a few issues with the pre-made models. First, pre-made models of the human body often uses soft tissues not specifically designed for biological material. For the THUMS model, much of the hyperelastic materials used where designed for rubbers and foams rather than human soft tissue. While there are a few specific biological materials in the LS-DYNA material database, they are often neglected. The THUMS viscus, instead of using the MAT HEART TISSUE or MAT LUNG TISSUE, uses MAT LOW DENSITY FOAM. The specific reasons for using low density foam is unknown, but switching out the low density foam material with the lung tissue did not result in a more stable model. This suggests FE materials designed specifically for biological tissues are still immature compared to general hyperelastic materials and the latter is still preferred. Second, validation of pre-made human body models do not offer enough coverage to ensure stability of the model in all scenarios. While THUMS have been validated in the frontal direction, it still suffered many stability issues due to the higher than normal crash pulse. Other than the higher crash pulse, other parts of the simulation are within the norm of a frontal crash test. Finally, the inherent complexity in the full body model creates more sources of instability that would be easy to eliminate in a simple model. In the THUMS, there are three parts that constantly have to be adjusted to prevent failure while none were unstable in the HIII model.

Debugging stability errors in the simulation were approached from a few directions. The most time efficient debugging technique was to isolate problematic parts and test its behavior in a controlled environment. This allows processing only the necessary parts of the model and allows quicker turn around time for the debugging cycle. In addition, isolating the problematic part helps with understanding the source of the error. However, this proved to be difficult to accomplish with the viscus as the complex interactions inside THUMS chest cavity was difficult to replicate independently. In addition, the complexity of THUMS coupled with the sheer code size made it difficult to understand the full scope of the interaction of a single part. While it was possible to test the material property of the viscus via a simple model, the geometry aspects also contributed to the instability and this interaction could not be accounted for. Therefore Chapter 2. Finite Element Simulation of Frontal Crash using THUMS 30 a more time consuming approach to debugging the THUMS was used. The debugging cycle for FE simulations consisted of attempting various modifications and running the full model. Although time consuming, this form of debugging provides the most realistic feedback as each run is not run in an isolated environment and all the interactions are always present in each run. A disadvantage of this method is that understanding of the source of the problem suffers due to the complex interactions.

The debugging cycle necessitated reducing the simulation time whenever possible. As the simulation were 20 hours per run, rigid materials were used as much as possible for reducing simulation time while at the same time allowing for modeling of supporting structures. These changes are not as signiﬁcant compared to the processing time devoted to THUMS which is much longer. However, THUMS does oﬀer a balance between complexity and processing speed. Much of THUMS is coarsely meshed, which contributes to the instability, but also decreases processing time. In addition, hyperelastic materials that are softer (lower bulk modulus) have faster processing times due to a higher Courant-Frederick-Levy (CFL) Criteria, or the stable time step. A full body FE model that was oriented more towards geometric realism and stability could easily quadruple the processing time of THUMS.

2.5.2 PMHS validation

In addition to validation against the HIII model, the model was also validated against existing PMHS tests [61]. While THUMS has been validated in many studies, this speciﬁc loading scenario - 40 mph frontal DV with no airbag deployment should also be compared to a PMHS study for validation. The test used is test number 3084 from NHTSA Biomechanics test database [62]. The PMHS test is a frontal test sled with a 58.2 km/h delta V (36 mph) without airbag deployment. Further test information can be viewed in the full report.

The PMHS test conﬁguration is reproduced in Figure 2.15. As can be seen in Fig- ure 2.15, the PMHS test set-up also includes the occupant compartment and steering wheel, rather than just the seat and knee bolster in the simulation. It was mentioned in the report that the occupant struck the B-pillar on rebound, which is not possible in Chapter 2. Finite Element Simulation of Frontal Crash using THUMS 31

Figure 2.15: PMHS diagram with measurements.

the simulation. In addition, the test uses a 2 point shoulder belt and an aluminum plate to anchor the PMHS midsection.

Figure 2.16: PMHS applied sled pulse.

The applied sled pulse is shown in Figure 2.16. Compared to the crash pulse in the simulation (Figure 2.5), the magnitude of the peak acceleration is lower by approximately 5 G’s while the duration of the crash pulse is slightly longer by approximately 30 ms.

Photographic snapshots of the test can be seen in Figure 2.17. It can be seen that the PMHS does not rebound as symmetrically as shown in the simulation, but rather leans towards the right side. Chapter 2. Finite Element Simulation of Frontal Crash using THUMS 32

(a) PMHS pretest. (b) PMHS post-test.

Figure 2.17: PMHS validation photos. 2.17(a) shows the pretest snapshot. 2.17(b) shows the post-test snapshot.

(a) PMHS x - displacement

(b) PMHS z - displacement

Figure 2.18: PMHS validation displacement traces. 2.18(a) shows the x - displacements. 2.18(b) shows the z - displacements.

The maximum head z - displacement as shown in 2.18(b) is -14.38 cm, while the maximum head displacement in the simulation was around -50 cm to -55 cm, almost four times that of the cadaver test.

Using the high speed motion capture displacement measurements (Figure 2.18), the Chapter 2. Finite Element Simulation of Frontal Crash using THUMS 33

Figure 2.19: Diagram of maximum PMHS head movement, the white marker near the windshield shows the approximate maximum position of the head.

approximate position of the head at maximum displacement can be calculate based on the PMHS diagram shown Figure 2.15 by calculating the ratio of centimeters to pixels and converting the maximum head displacements to pixels. The result is shown in Figure 2.19. It can be seen that the cadaver head comes close to that of the windshield and the cadaver most likely would have moved further forward if the steering wheel was not present.

PMHS 0 THUMS

−20

−40 z - displacement (cm)

0 50 100 150 200 time (ms)

Figure 2.20: PMHS vs THUMS head z - displacement

The z - displacement of the PMHS vs THUMS is shown in Figure 2.20. It appears that the first 75 ms of the two are similar, but at approximately 75 ms it appears that Chapter 2. Finite Element Simulation of Frontal Crash using THUMS 34 the PMHS hits the maximum z displacement possibly due to the steering wheel while the THUMS head continue to move downward. Also of interest is the fact that the PMHS head displacement slope before the maximum is steeper than the THUMS head displacement slope. This further shows that it is likely the PMHS head would have dropped steeper if not blocked by the steering wheel or windshield. Given that the steering column stroke was 0, it is unlikely that the amount of force required to stop the PMHS was significant enough for the head displacement to increase by a significant amount.

Figure 2.21: PMHS passenger side kinematics as shown in Forman et al [63]

Since the PMHS test shown above is a driver side test, the steering wheel becomes a major difference between the PMHS test and the simulation, therefore a passenger side test is also shown from the Forman et al study in 2006 [63]. Figure 2.21 shows the lateral view of the passenger in a 48 kph (≈ 30 mph) DV PMHS sled test. It can be seen that the approximate displacement of the head in the z - direction is 15 cm, closer to the NHTSA biomechanics test than the simulation. The effects of the airbag is still an unknown in this test, however. Thus it shows that the simulation head z - displacement would be likely higher than a PMHS test using the same configuration, if the results of the passenger side PMHS test is extrapolated to a 40 mph DV.

The ideal PMHS validation test would include both a 40 mph DV and a passenger side Chapter 2. Finite Element Simulation of Frontal Crash using THUMS 35 interior with no airbag. This would show the kinematic behavior of the PMHS with conditions more similar to that of the simulation.

The results of the PMHS validation suggests that the THUMS head displacement was more than a PMHS would experience, which could point to overly soft material model in THUMS or a less rigid restraint system.

2.5.3 Simulation results

The simulation snapshots show that much of the behavior of THUMS is similar to that of the HIII. The extra extension and bending of the spine and neck could be attributed to the softer tissues of THUMS compared with the HIII. The general motion of THUMS does not seem unrealistic for such a severe crash however. Especially without an airbag and instrument panel, there is nothing to stop the momentum of the head. On the other hand, had an instrument panel existed in the simulation, it would have been a severe impact judging from the ﬁnal position of THUMS head It was expected that the motion be severe given the crash pulse; however, the belt pay out seems to be normal. In the HIII simulation, the shoulder belt restrained the model to a smaller range of motion than the THUMS. The HIII also showed a higher amount of forward shifting in the lap belt, but not enough to cause submarining. The stresses caused by the belt on the model exterior as shown by the fringe plots (Figure 2.11 and Figure 2.12) are similar in the overall distribution of the stress in the thorax. Both the THUMS and HIII chest are experiencing less stress than the shoulder, which indicates that the restraint system is working appropriately by shifting the force experiences to the stronger parts of the body. Both models are also experiencing a concentration of stress shown in red at areas close to the lap belt on the sides; however, THUMS is experiencing less stress than HIII. This could indicate that the mass of HIII is balanced diﬀerently, as the THUMS also did not shift forward as much in the lower body. Although the stress distribution is comparable, the HIII stresses are about 100 times less than that of THUMS for the same color. This could indicate that the stresses are less evenly distributed for the HIII and that stresses tend to concentrate in a few areas such that only by lowering the maximum value of the legend could we see stress variation for all of the body. Chapter 2. Finite Element Simulation of Frontal Crash using THUMS 36

This lack variation in the stress is further demonstrated by the skeletal fringe plots (Figure 2.13 and Figure 2.14). In this instance, we see that the stresses are concentrated in the ribs of the HIII while distributed more evenly across the skeletal structure for the THUMS. The highest points of stress on the ribs are to the sides rather than the front of the models. This means that the restraint systems are not causing excessive concentrated stress on the thorax and that the stresses are caused more by overall compression of the rib cage.

The HIII FE model is based on a physical ATD that is itself suppose to behave in a similar manner to a cadaver in a crash test. It was designed from the beginning to follow the speciﬁcations of the physical HIII ATD. While ideally this means that it should behave similarly to a cadaver, this is an indirect approach. In addition, this also means that the HIII FE model inherits all the disadvantages of the physical model without taking advantage of some of the features of the FE method. On the contrary, THUMS is designed from the geometry of the human body and therefore oﬀers much better geometric accuracy. In terms of material properties, while the THUMS does not have detailed material parameters for every part of the body and much of the soft tissues are similar, it does use more distinct materials than the HIII.

However, THUMS raises a more general issue with geometrically complex and detailed model. The idea of a model such as THUMS is to simulate the human body in a virtual environment that is geometrically close to the actual human body. The material properties may be less accurate as soft tissues cannot be described easily. The inherent complexity, however, can hinder understanding of the behavior of THUMS which is a simpliﬁed version of the actual human body. Since one of the goal of the model is to contribute to our understanding of the actual phenomenon, a balance between complexity and ease of analysis should be carefully established. The material properties of the model is not realistic enough such that it can be treated as a valid model to stand on its own in all scenarios. Therefore it can be diﬃcult to determine the parts of the results that need to be validated if the geometry is overly complex.

THUMS is still a more direct approach than HIII due to its direct correspondence with the human body. The drawbacks of THUMS can be minimized by increasing the amount Chapter 2. Finite Element Simulation of Frontal Crash using THUMS 37 of validation done and also performing more validation with individual body components as applicable. Temporary simpliﬁcation of certain parts such as the viscus could improve the stability until more robust material models are developed that are well understood. THUMS is a good step in the right direction of FE modeling of the human body, but more eﬀort is needed to fully realize the potential of FE modeling. Chapter 2. Finite Element Simulation of Frontal Crash using THUMS 38

2.6 Conclusion

A 40 mph crash simulation was created using THUMS and the experiences documented here. It was found that THUMS was not stable at such a high velocity impact without modification to some of its material properties. Modifications to the chest material properties, careful usage of contacts, and some modification to the THUMS mesh in the chest allowed the simulation to be stable for the duration of the crash pulse. While these modifications might detract from the realism of the simulations, the THUMS model was still more realistic than the crash test dummy models in its geometry. Chapter 3

Development of Finite Element Injury Measures for THUMS

3.1 Introduction

FE models of the human body have been improving in realism over the past couple of years through advancements in computational power. Over time, they could become a suitable alternative to ATDs for mass and repeatable use in analysis of crash tests. FE models such as THUMS offer several benefits over ATDs. First, THUMS offer measurement possibilities that are not possible with traditional ATDs. Outputs such as stress, strain, local forces, and energy are useful measurements that should be also be used to supplement the existing injury measures. While these injury measurements might not be easily validated with experimental tests, they can offer a better understanding of the model behavior and possibly contribute to validating existing injury measures. FE injury measures are different from crash test injury measures because they are not measured with equipment that the crash test dummies have. Instead, FE injury measures can be measured directly as part of the FE output. This is a more robust way of measuring output as the desired measurement is not encumbered by extraneous equipment. Virtually any part of the FE model can be measured. Second, although FE models do not directly exhibit injury behavior, they do offer material properties that can response

39 Chapter 3. Development of Finite Element Injury Measures for THUMS 40 realistically to applied loads. Deﬁning realistic material properties can be enhanced with well-correlated injury measures in addition to proper material property testing procedure. The material properties can also be supplemented with injury criterion that are developed as part of the injury measure, which could give more realistic failure in the model. Third, they do not suﬀer as much from experimental variations and overall are more repeatable. Accurate injury measures can be developed from the FE model (with experimental validation) and be used to better describe injury scenarios.

A caveat of FE models is that they are less validated compared to ATDs. Thus it is desirable to see how injury measures from FE models compare to those from ATDs. FE models and ATDs have some differences that can create difficulties when translating injury measures designed for ATDs to those of FE models. First, FE models for the human body such as the THUMS have complex and accurate outer geometrical detail and also validated physical response in most critical areas [29, 30, 33–35, 46, 58], but the THUMS does not have all organs of the human body, so organ level injury measures are not possible with THUMS. Although ATDs mimics the human body [25], they also have to simplify the body as they cannot present all the complexities in the human body. Second, ATDs have dedicated pieces of equipment to measure both kinematic and kinetic responses. While FE models of the human body can output these responses too, there is a question of whether they can be correlated directly to the ATD outputs. There is often some discrepancy between the location of the measurement, the method of measurement, and the output accuracy of the measurement. Third, there is the issue of an experimental model vs a numerical model. When dealing with physical measuring equipments such as with ATDs, there are additional concerns with the correct filtering, the sensitivity of the equipment, and correctly calibrating the equipment before the test. On the FE side, there are issues with numerical realism, filtering of the output, and dealing with different ways of measuring the output.

A problem with both methods of injury measures is that it is questionable if either can be used as the accurate representation of the human body response. However, since injury measures have been developed using physical measurement devices, the physical measurements should be used as a more validated method of injury measures. When developing injury measures for FE models, it should correlate as closely to the ATDs as Chapter 3. Development of Finite Element Injury Measures for THUMS 41 possible for validation or be validated separately. PMHS would be a good candidate to validate FE models if such data was uniform and well controlled.

3.2 Methods

Correlating injury measures between an FE model and an ATD can be a simple or a complex procedure, and can depend on whether there are variations in model geometry, material property differences, constraints on measurement options, or simplifications assumed in either the ATD or the FE model. For example, since the HIII model is not as anatomically correct as the THUMS, there are difficulties when translating some of the injury measures designed for the HIII to THUMS. While any part of the THUMS body can be directly measured for loads, injury measures from existing ATDs and models should be used initially as these are well validated measures. The injury measures that are often measured in HIII are the HIC, Nij, chest deflections, and chest acceleration. The injury measures are located in Figure 3.1. Some additional measurement points were added to provide more information. These injury measures are not necessarily the best injury measures, but most of these injury measures have been extensively validated and offer a common way of correlating quantitative measures to injury.

HIC Nij Chest 3 ms Clip

Chest Deflection

Pelvic Abdominal Acceleration Acceleration

Figure 3.1: THUMS injury measure locations Chapter 3. Development of Finite Element Injury Measures for THUMS 42

3.2.1 Head Measure

The HIC is an injury measure that is often associated with head injury. HIC, originally called the Tolerance Limit (TL) formula is formulated by Versace [64] with Equa- tion 3.1. Where a is the acceleration, T is the duration that maximizes the function.

(R a dt)2.5 TL = (3.1) T 1.5

The time dependent nature of the function means that high magnitude, short duration pulses will yield the same value as that of low magnitude, long duration pulses. Naturally, a cap on the maximum duration is necessary. The original NHTSA established limit on the HIC is 36 ms, this limit was later reduced to 15 ms [65]. Ideally, this value should correlate to some risk of head injury. There is still some controversy regarding the use of HIC, but it is the most widely used injury measure when it comes to head injury.

The HIC is often written as HIC36 or HIC15 to distinguish between the 36 ms and 15 ms limit on the duration. One interpretation of HIC has been provided by Eppinger as “a measure of the rate of change of speciﬁc kinetic energy . . . modulated by the square root of the average acceleration” over the time duration [66].

Z SI = an dt (3.2)

The HIC function originally came from the Gadd Severity Index (SI), as shown in Equation 3.2 [53]. If the exponent n is removed from the Equation 3.2, it can be seen that the formula becomes the change in the velocity over the duration. This injury measure associates higher accelerations with greater contribution to the index. The formula of the SI comes from Gadd’s reasoning of the behavior of animal tissue. Gadd assumed that animal tissue failed somewhere between brittle and viscous failure mode. In brittle failure mode, after a certain force has been reached, the tissue fails suddenly. This is the idea behind an acceleration limit (since acceleration is proportional to force), Chapter 3. Development of Finite Element Injury Measures for THUMS 43 in which a maximum value is placed on the acceleration. In viscous failure mode, the tissue failures occurs after a certain amount of shear strain, which can be correlated to the duration of loading as well as an increase in load intensity. Thus the SI comes from the idea that soft tissue partially fails in brittle mode, in which higher accelerations contribute more to the index, and it partially fails in viscous mode, in which the increases in acceleration as well as the duration aﬀects the index. A limit of 50 ms was originally suggested by Gadd as the limit to the time duration, and n was approximated to be 2.5 from the Wayne State Tolerance Curves [67]. The value that Gadd suggested to be the tolerance limit was set at 1000. Although there are some questions regarding the formulation of SI, such as whether it can be assumed that the composite materials of the head can be assumed to behave like animal tissue, Versace also criticized Gadd’s interpretation of the Wayne State data. Since Gadd attempted to ﬁt a straight line to the log plot of the Wayne State data using the Equation of the form log A = m log T +log K, where m is −0.4 and k is 15.85, it can be shown that the solution can be in the form of Equation 3.3, Equation 3.4, or Equation 3.5. So that in fact there is nothing special about the exponent 2.5 or the limit of 1000.

A = 15.85T −0.4 (3.3)

TA2.5 = 15.852.5 = 1000 (3.4)

AT 0.4 = 15.85 (3.5)

It should be noted that the eﬀective acceleration A is deﬁned as the magnitude of the acceleration-time trace if the acceleration is relatively constant, or as Equation 3.6 if it is irregular in shape. In other words, it is the average acceleration.

1 Z A = a dt (3.6) T Chapter 3. Development of Finite Element Injury Measures for THUMS 44

Thus Equation 3.4 can be rewritten as Equation 3.7 by substituting for eﬀective acceleration and the formula for HIC follows naturally. It can be seen that only if a is assumed to be constant, then the SI equation (Equation 3.2) is equivalent to Equation 3.7. So in fact the SI does not oﬀer a mathematically correct representation of the Wayne State data unless the pulse is a square wave.

Current NHTSA regulation sets the maximum duration for HIC as 15 ms or 36 ms, usually written as HIC15 and HIC36, respectively. The limit for HIC15 is set at 700 [50] and the limit for HIC36 is set at 1000.

TA2.5 = 1000 1 Z T ( a dt)2.5 = 1000 (3.7) T

It should be noted that the formulation for HIC is based purely on a curve ﬁtting of data and did not start from any physiology-based theory of head injury.

Soft Tissue

Diploë Brain

Figure 3.2: The three layers of THUMS head

THUMS has a simpliﬁed model of the head that has approximately three layers. Select material properties for each part is detailed in Table 3.1. As can be seen, the outer skin is composed of two layers of slightly tougher shell with solid tissue in between. The skin layer is around 1 cm in thickness, which might also account for the eﬀects of hair, but this is unknown. The skull is also composed of three layers, with two layers of cortical sandwiching a layer of cancellous. The brain is a single solid part with a much softer Chapter 3. Development of Finite Element Injury Measures for THUMS 45

Young’s Modulus. Of note is the material properties of the skull and the brain, which are all rigid bodies. This means it will not be possible to acquire stress and strain values from the skull or brain. This might not be problematic for the skull, but strain values for the brain could be useful in previously mentioned injury measures for the brain. If necessary, these parts could be changed to use a linear elastic material.

Table 3.1: THUMS head divided into three layers with material properties.

Layer Part Material Type Material Property Outer Shell Elastic E = 22 MP a Skin Solid Tissue Viscoelastic K = 2.3 MP a Inner Shell Rigid E = 22 MP a Exterior Lamina Rigid E = 8 GP a Skull Diplo¨e Rigid E = 200 MP a Interior Lamina Rigid E = 8 GP a Brain Brain Tissue Rigid E = 102 kP a

The HIC for THUMS is measured at the Center of Gravity (CG) of the head as shown in Figure 3.3. This is the ideal measurement location if eﬀects of rotation of the head are to be minimized. As the HIC only uses translational components of acceleration, any rotational eﬀects should be minimized. This does not mean that translational acceleration are a better indicator of injury, merely that HIC does not utilize rotational kinematics.

Brain

HIC Node

Figure 3.3: Node location for measuring HIC Chapter 3. Development of Finite Element Injury Measures for THUMS 46

3.2.2 Neck Measure

The major injury measure for neck injury in frontal crashes is the Nij neck injury

criterion. The development of Nij has its roots with the experiments of Prasad and Daniel [68, 69] and Mertz et al [70]. Experiments by Prasad and Daniel showed a relationship between the neck moment and axial force of matched piglet and child dummy

tests. This information prompted the creation of the Nij neck injury criterion in which used the axial force in conjunction with moments in the occipital condyle to create an

injury measure. The formula for Nij is shown in Equation 3.8.

Fz My Nij = + (3.8) Fint Mint

Equation 3.8 is composed of two terms, one for the axial force, and one for the moment. Prasad and Daniel noticed that there is an interaction between the tension and bending of the neck, that severe injury can be caused by either excessive bending or excessive tension, and that an increase in one value would cause a decrease in tolerance to the other. A linear relationship between the two variables was assumed. Since the neck can experience two types of axial force - tension or compression, and two types of bending modes - ﬂexion and extension, a limit for each mode of axial force bending must be used. There are a total of four limit values for a ﬁftieth percentile male, which are shown in Table 3.2.

Table 3.2: Nij limits for ﬁftieth percentile male

Limit Type Critical Value Fint (tension) 6806 N Fint (compression) 6106 N Mint (ﬂexion) 310 N · m Mint (extension) 135 N · m

The limit values shown normalizes the force and moment in Equation 3.8 so that an

Nij value that exceeds 1 exceeds the limit for Nij. The Nij function is best demonstrated via a graphical representation as shown in Figure 3.4. In the figure, the y-axis Chapter 3. Development of Finite Element Injury Measures for THUMS 47 represents the axial force and the x-axis represents the moment. Each tip of the parallelogram represents the acceptable limit for that value. It can be seen that the linear relationship is used to define the interaction between forces and moments, which might be a simplification of the actual relationship. In the calculation for Nij, the values at the vertex of the parallelogram would all become one as the forces and moments are normalized in each axis by the limits. Tension

Extension Flexion

Compression

Figure 3.4: Graphical illustration of Nij, the area shaded green represents acceptable values.

Implementation of the Nij function in THUMS has is diﬀerent than the implementation in the HIII dummy. Although the THUMS technically has an occipital condyle, this is not the sole source of load transmission. There are some soft tissues surrounding THUMS’ neck which also transmit some loads, but these are not expected to be great and are ignored. Another source of load transmission between THUMS’ head and the body is various springs and dampers that are also part of the THUMS neck as seen in Figure 3.5. The forces and moment transmitted through these discrete elements cannot be ignored and also must be taken into account. Thus the use of a section plane taken across the occipital condyle measures the amount of force in the occipital condyle. In addition, the forces in the discrete elements are projected to the normal vector to the section plane and added to the section forces. The moments are taken by using the vector from the occipital condyle to the discrete element as the moment arm, then the cross product between the moment arm and the force vector along the element yields the moment that is added to the section plane moment. Chapter 3. Development of Finite Element Injury Measures for THUMS 48

The THUMS neck consists of spinal column, soft tissue, and connective tissues as shown in Table 3.3. The vertebral bodies consists of the vertebral discs and the vertebral bodies.

Table 3.3: THUMS neck properties for select parts of the neck. Some parts such as the muscles and ligaments are shown as a range because there are a variety of diﬀerent properties for these.

Layer Part Material Type Material Property Cortical Rigid E = 5 GP a Vertebral Body Spongy Rigid E = 70 MP a Nucleus Pulposus Elastic E = 198 kP a Vertebral Disc Annulus Fibrosis Elastic E = 13.3 MP a Fibers Seatbelt k = 2.67 N/strain Cartilage Elastic E = 12.6 MP a Connective Vertebral Ligaments Elastic 3 MP a < E < 31 MP a Muscle Spring (Muscle) 3 N < Fmax < 72.5 N Soft Tissue Skin & Misc Viscoelastic K = 2.30 MP a

Since there are around 400 discrete elements in THUMS neck area, the forces and moments are added via a computational routine. There will be some reduction in the comparability between this Nij value and the HIII, but an occipital condyle only measurement will underestimate the forces and moments. Computation of the Nij occurs for every time step of the simulation. Nodal positions for the discrete elements must be tracked along with the position of the occipital condyle. Forces and moments for the occipital condyle are diﬃcult to acquire since it is not implemented as an actual joint, but instead a contact between the head and neck. Contact forces could be acquired, but there are no moments with contact forces. A nodal force group could be deﬁned at the occipital condyle, but the interface between the head and the neck is quite complex and this would have been time consuming to implement. Furthermore, various approx- imations would have meant this would not be an optimal solution. For this study, a section force across C1 is used to approximate the forces at the occipital condyle. This solution has the advantage that forces and moments are calculated automatically based on element results. However, it is not accurate either. Chapter 3. Development of Finite Element Injury Measures for THUMS 49

Springs & Dampers

Figure 3.5: THUMS neck springs and dampers between head and shoulder

3.2.3 Chest Measures

Chest injury measures are useful to predict rib fractures and soft tissue injuries, depending on the measure. Chest spinal acceleration is one choice for an injury measure that has been used to correlate to injury [71, 72]. Extreme spinal acceleration is usually considered significant only if it lasts longer than 3 ms in duration, this ensures that infinitesimally small extreme accelerations are not considered as the human body can tolerate large accelerations in small durations. The chest 3 ms acceleration, also called the 3 ms clip, can be understood as correlating to the load the chest experiences. Since F = ma, then the chest 3 ms clip can be understood as proportional to the force experienced by the chest, normalized by the effective mass. This injury measure has been limited to 60 Gs as dictated by NHTSA [50].

Implementation of the chest 3 ms clip was simple as the acceleration can be directly measured at a node. A massless node was created at the centroid of T6 and its movement constrained by the average movement of all the nodes in T6. By using an interpolated node, noise from individual nodes are reduced as well as allowing the arbitrary placement Chapter 3. Development of Finite Element Injury Measures for THUMS 50 of the node. The acceleration of the node can then be measured with LS-DYNA’s nodout output.

Chest deflection is another major injury measure for the chest. Unlike chest acceleration, chest deflection calculation requires measuring the change in the distance between two positions between the sternum and the spine. This change in distance is in response to a compressive force on the chest. One caveat of the chest deflection is that it is affected by the chest depth. A person with a deeper chest depth is less likely to be affected by the same deflection as a person with a shallower chest depth. It is due to this reason that sometimes the chest deflection is normalized by the chest depth to get a percentage deflection value. For the fiftieth percentile male, the absolute deflection limit specified by NHTSA is 63 mm or 2.5 in [50], except for certain circumstances, in which the limit is 76 mm. Normalized deflections will also be calculated, but there is no regulation for this measure. However, the Chest depth for the fiftieth percentile male is defined as 9.3 in ± 0.2 in (236.2 mm ± 5.08 mm), which means that the deflection allowed is approximately 26.9%. The chest depth for the THUMS is approximately 232 mm, which is within the range defined by Federal Motor Vehicle Safety Standard (FMVSS) 208.

Implementation of the chest deflection is shown in Figure 3.1 as the distance between the two nodes. The deflection is measured as the distance between a node on the sternum and a node on the back of the spine. Normalized deflections are calculated by dividing the current chest depth by the original chest depth at each point in time.

An extension to the idea of chest deflection is the calculation of deflection for individual ribs. The rib deflection measurements will not only give an idea of the deformation for each individual rib, but also give a general picture of the pattern of deformation. Chest deflection itself can be oversimplified in the fact that deflection is only calculated at one location, which may or may not be the center of the compressive forces. Although since it is at a position where the seat belt passed through, and also in front of the heart, it is a valuable measurement position. Individual rib deflection values can give an overall pattern of the chest deflection which maybe valuable in determining the effect on the Chapter 3. Development of Finite Element Injury Measures for THUMS 51 underlying organs. Since certain ribs are covering certain organs, a higher deflection in those ribs could mean that the organ behind them experiences more force in them.

Implementation of the THUMS rib displacement has some ambiguities in it. From Figure 3.11, there is a difference between the THUMS rib cage and the HIII rib cage as can be seen by the direction of the ribs. In the HIII rib cage, the ribs are horizontal in the thorax and normal to the thorax wall. This is a simplified representation of the human rib cage. On the other hand, the THUMS thorax is geometrically more accurate with ribs that are not horizontal in the thorax. Therefore to create a comparable rib deflection measurement, the difference between the HIII and THUMS must be accounted for. The rib deflections are implemented in two ways in THUMS. As the deflections are horizontal in the HIII, horizontal deflections are also measured in THUMS. This is shown as horizontal lines in Figure 3.6.

Figure 3.6: THUMS horizontal rib deﬂection measure

While the horizontal deflection measurement means that the HIII and the THUMS are comparable in terms of the placement of the measurement positions, it also means that the deflections measured in THUMS are no longer for the same rib. Deflection measurements that start at one rib on the spine end up on a different rib on the chest. There the same rib deflection measurements are also taken, as shown in Figure 3.7. Chapter 3. Development of Finite Element Injury Measures for THUMS 52

Figure 3.7: THUMS parallel (to the rib) rib deﬂection measure

The acceleration and the deﬂection can be combined into a mixed chest injury measure, the CTI. The CTI uses an idea that is similar to Nij in the fact the it assumes a linear relationship between the acceleration and the deﬂection. The CTI can be seen in Equation 3.9.

A D CTI = max + max (3.9) Aint Dint

The CTI is similar to Nij in the equation form and the interpretation for Aint and Dint is also similar. The two terms both represent intercept values for the chest acceleration and deﬂection. Unlike the Nij, the CTI is not allowed to reach the full intercept value when only acceleration or deﬂection exists. Instead, separate critical values Dc and Ac are imposed, as seen in Figure 3.8.

The intercept and critical values for CTI are listed in Table 3.4. The critical values for CTI are equal to the critical values for chest deﬂection and chest 3 ms clip.

With the CTI, the more acceleration a person tolerates, the less deﬂection he/she can handle without sustaining injury. The origin of this injury measure comes from the Chapter 3. Development of Finite Element Injury Measures for THUMS 53

120

Dint 100

(mm) 80 Dc 60

Deﬂection 40

Chest 20 A A 0 c int 0 20 40 60 80 100 Chest Acceleration (G’s)

Figure 3.8: Illustration of the CTI. The dashed line is the original CTI while the solid line softened the limits for chest acceleration and chest deﬂection at the end points of the curve

Table 3.4: Key limit values for CTI, for the 50% male

Intercepts Critical Values Deﬂection (Dint) Acceleration (Aint) Deﬂection (Dc) Acceleration (Ac) 102 mm 85 G’s 76 mm 60 G’s

idea of accounting for both seat belt loading and airbag loading. The seatbelt applies a concentrated load that is more likely to cause chest deflection compared to the airbag. While the amount of force exerted by the airbag could be greater than that of the belt, the chest deflection could be much less since the airbag exerts the force over a greater area. Since the CTI accounts for both deflection and acceleration, then it is presumed that both the belt and the airbag are accounted for with this injury measure. This is also the reason for having different critical values for CTI than the intercept values, to favor safety designs that uses both seatbelt and airbags.

In addition to chest injuries and rib fractures, there are also soft tissue injury criteria that can be calculated based on existing measurement. One proposed injury criterion for soft tissue is the viscous criterion [73]. The viscous criterion is a result of the inadequacies of the 3 ms acceleration criterion and the compression criterion to account for rate effects of soft tissue. Lau and Viano argues that the acceleration criterion is for the skeletal system only while the deflection accounts only for the deformation Chapter 3. Development of Finite Element Injury Measures for THUMS 54 component. Adding a rate sensitive component to the deflection criterion will in effect account for both deformation and the rate effects. The viscous criterion thus uses both the velocity (V (t)) and the compression (C(t)) as shown in Equation 3.10. As noted by the authors, under extremely slow loading conditions, the compression criterion will be more effective than the VC as the velocity approaches zero. Under higher velocity loading, the compression will also be small so the velocity will play a bigger role. The viscous criterion can be calculated by extracting the nodal velocities in addition to the compression. In real life situations, the velocity can also be calculated by integrating the acceleration.

d[D(t)] V (t) ≡ dt D(t) C(t) ≡ Initial Chest Depth VC ≡V (t) · C(t) (3.10)

Ribs Viscus

Upper Abd.

Figure 3.9: General components of THUMS chest Chapter 3. Development of Finite Element Injury Measures for THUMS 55

THUM cortical properties 150

) 100 Modiﬁed a P Originial M (

50 stress

0 0 0.005 0.01 0.015 0.02 strain (unitless)

Figure 3.10: THUMS modiﬁed rib cortical properties, the dashed line shows the original rib cortical property that was in THUMS, the solid line represents the new cortical stress vs. strain behavior.

The THUMS chest consists of the outer skin and soft tissue, the rib cage, and internal viscus. The version of THUMS used in this study does not include actual chest organs but larger solid parts that are used as placeholders for the organs to ensure bioﬁdelity as seen in Figure 3.9. There are two major parts that are representative of the THUMS chest organs, the viscus - comparable to the lungs and heart, and the diaphragm. The diaphragm divides the thoracic cavity and the abdomen. The parts of the THUMS chest are shown in Table 3.5. The skin and outer soft tissue of THUMS are combined into one part as described by a viscoelastic material, which covers an elastic shell. The rib cage is covered by a layer of null elements, possibly to enhance stability.

The ribs were modiﬁed as shown in Figure 3.10. The ideal properties for the ribs cortical bone would include damage and subsequent failure that are realistic compared to an actual human rib. This was attempted in THUMS as can be shown by the piecewise stress vs. strain curve that contains yielding, damage, and subsequent failure. However, this seems to over predict the number of rib fractures and cause unrealistic behavior [60]. The curve was modiﬁed to remove the damage phenomenon as this was believed to cause the excessive failure in the ribs. The failure criterion was kept at the same strain so that elements past 2.13% plastic strain were deleted from the part. This allowed the Chapter 3. Development of Finite Element Injury Measures for THUMS 56

Table 3.5: Material properties for some of THUMS thorax

Layer Part Material Type Material Property Soft Tissue Viscoelastic K = 2.30 MP a Outer Inner Skin Elastic E = 22 MP a Ribcage Cover Null Null Rib Cortical Plasticity Ei = 10.59 GP a Rib cage Rib Spongy Piecewise Ei = 40 MP a Cartilage Piecewise Ei = 49 MP a Viscus Cellular Rubber 4.15 < E < 246 kP a Organs Diaphragm Elastic E = 1 MP a Serous Membrane Elastic E = 1 kP a

number of rib fractures to be more realistic. The THUMS sternum was given the same material properties as that of the ribs. The spongy bone was not modiﬁed.

Hybrid III Total HUman Rigid Model Model for Safety 50% Male (THUMS) 50% Male

Figure 3.11: HIII vs. THUMS rib diﬀerences - HIII has rib cage that consists of horizontal bars in the THUMS thorax while THUMS have more natural downward ribs

3.2.4 Abdominal Measures

Although not part of the regulation, abdominal and pelvic accelerations are also measured. The abdominal acceleration is measured at the CG of the THUMS mid abdominal organ. As can be seen in Figure 3.12, the THUMS abdomen is divided into three generalized abdominal organ, the upper abdomen, the middle abdomen, and the lower abdomen. These three parts represents all the organs in the THUMS abdominal cavity. Select parts of the THUMS abdominal organs are shown in Table 3.6. Chapter 3. Development of Finite Element Injury Measures for THUMS 57

Table 3.6: THUMS abdominal parts and material properties

Layer Part Material Type Material Property Membrane Elastic E = 1.73 MP a Upper abdomen Tissue Viscoelastic K = 2.00 GP a Membrane Elastic E = 22.0 MP a Mid abdomen Tissue Crushable Foam E = 38.4 MP a Membrane Elastic E = 20.0 MP a Lower abdomen Tissue Viscoelastic K = 2.30 MP a Outer Elastic E = 2.00 MP a Pelvic ﬂesh Tissue Elastic E = 138 MP a Cortical Piecewise Ei = 17.3 GP a Spongy Damage 2 E = 15.0 MP a Pelvic bone i Sacrum Rigid E = 70 MP a Ligaments Piecewise E ≈ 20 MP a

Upper Abdomen

Abdomen

Lower Abdomen

Pelvis

Figure 3.12: THUMS abdominal organs - the abdominal organs have been simpliﬁed into the upper, mid and lower abdominal organs.

The force exerted by the lap belt (Figure 3.13) on THUMS and the force exerted on the iliac crest (Figure 3.14) are also measured to compare with the shoulder belt forces. Both of these injury measures are measured using nodal force groups, which uses the summation of the force on the nodes to calculate the total force. Chapter 3. Development of Finite Element Injury Measures for THUMS 58

Figure 3.13: Lap bet force measurement nodes. The yellow triangles points to the nodes where the force is measured

Figure 3.14: Iliac crest measurement nodes. The yellow triangles points to the nodes where the force is measured

3.2.5 Vehicle Interior Measures

In addition to measuring the components of THUMS body that provides insight into injury mechanism, there are additional components in the interior which indirectly correlates to some injuries. Example of these are the forces of the lap belt against the lap, forces of the shoulder belt against the shoulder, and tension of the belt at various locations. Chapter 3. Development of Finite Element Injury Measures for THUMS 59

The seatbelt consists of the shoulder belt and the lap belt, which serves to evenly distribute the forces of a crash between the thorax and the pelvis. Nominally, the forces in the shoulder belt and the lap belt should be consistent. Therefore it is desirable to measure the force of THUMS against both the shoulder belt and the lap belt. Too much forces in the lap belt, and there is a danger of submarining. Too much forces in the shoulder belt, and there is a danger of triggering the load limiters and causing broader head excursions. Studies have shown that restraints that engage both the shoulder and lap belt reduces fatality much more than restraints that only contains shoulder belts.

Belt tension and elongation is also measured at various positions which are deemed high stress locations. These measurements are performed at the D-ring, buckle, and belt anchors.

3.2.6 Test scenario

The case with a standard three-point restraint, no load limiter, and buckle pretensioner was used to investigate the accuracy of the injury measures. The case is shown in Figure 3.15. No airbag was modeled to keep the scenario simple. The injury measures are measured for the scenario and detailed in the results. The test scenario represents a single case of a multiple case study in to analyze the output of the injury measure.

3.3 Results

Selected injury measures from the simulation are divided into several categories: head, neck, thorax, and abdomen. Each injury measure is plotted above the acceleration pulse applied to show the relationship between input and output. Minimum and maximum values of the graph along with the times of occurrences are shown in the title of the plot for both the injury measure and the acceleration pulse. Chapter 3. Development of Finite Element Injury Measures for THUMS 60

Retractor

Knee Bolster

Pretensioner

Floor board

Figure 3.15: Three-point case for testing injury measures

3.3.1 Head

Head acceleration, HIC15, and HIC36 are three injury measures that are shown below. Head acceleration is directly measurable from FE output as a kinematic variable while HIC must be calculated from the time history of the acceleration. Hence HIC contains more information as the time history is also factored in. Chapter 3. Development of Finite Element Injury Measures for THUMS 61

Resultant head acceleration for THUMS is shown in Figure 3.16 and contains a surge in acceleration that lags behind the crash pulse by approximately 40 ms and peaks at around 90 ms, around 20 ms after the peak in the crash pulse. The head acceleration then gradually drops down over the next 120 ms. However, the resultant acceleration does not break down the acceleration into the x, y, and z components, which is more informative (see Appendix). Chapter 3. Development of Finite Element Injury Measures for THUMS 62

Head Accel (CFC 1000)−res vs Time Min:[0.000000s,4.78102e−007Gs] Max:[0.090501s, 73.5446Gs] 80

20 Head Accel (CFC 1000)−res (Gs) 10

0 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 Time (s)

Pulse (Acceleration) vs Time Maximum Acceleration Min:[0.112200s,−23.3Gs] Maximum Acceleration Max:[0.068200s,346Gs] 350 x−accel 300

250

200

150

100

50 Pulse (Acceleration)−Res (Gs)

−50 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 Time (s)

Figure 3.16: THUMS resultant head acceleration vs. time Chapter 3. Development of Finite Element Injury Measures for THUMS 63

Resultant head acceleration for HIII is shown in Figure 3.17. The maximum acceleration (90 Gs) is higher than the THUMS maximum acceleration (74 Gs). The peak also occurs earlier than the THUMS peak by approximately 15 ms.

HIII head acceleration vs Time Min:[0.002200s,0.0327839Gs] Max:[0.076500s, 90.3134Gs] 100

30 HIII head acceleration (Gs) 20

0 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 Time (s)

Pulse (Acceleration) vs Time Maximum Acceleration Min:[0.112200s,−23.3Gs] Maximum Acceleration Max:[0.068200s,346Gs] 350 x−accel 300

250

200

150

100

50 Pulse (Acceleration)−Res (Gs) 0

−50 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 0.2 Time (s)

Figure 3.17: HIII resultant head acceleration vs. time Chapter 3. Development of Finite Element Injury Measures for THUMS 64

The HIC15 (shown in green) for THUMS is plotted on top of the resultant acceleration

(shown as blue points) to show the relationship between HIC15 and resultant acceleration. The instantaneous HIC15, calculated using 15 ms of data forward from the current time point (Figure 3.18) shows the maximum HIC beginning approximately 10 ms after the peak in the crash pulse. The HIC15 window (as shown by the beige band) includes the maximum acceleration. Chapter 3. Development of Finite Element Injury Measures for THUMS 65

Max HIC: 488.384 at Time: [0.078301 0.093300] 100 500 Maximum HIC Head Resul Accel

0 0 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18

Pulse (Acceleration) vs Time Maximum Acceleration Min:[0.112200s,−23.3Gs] Maximum Acceleration Max:[0.068200s,346Gs] 350 x−accel 300

250

200

150

100

50 Pulse (Acceleration)−Res (Gs)

−50 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 Time (s)

Figure 3.18: THUMS HIC15 vs. time Chapter 3. Development of Finite Element Injury Measures for THUMS 66

The HIC15 (shown in green) for HIII is plotted on top of the resultant acceleration (shown as blue points) to show the relationship between HIC15 and resultant acceleration. The maximum HIC15 of 866 is higher than the THUMS maximum HIC15 of 488. The peak in HIC for HIII occurs 5 ms before the peak for THUMS. Chapter 3. Development of Finite Element Injury Measures for THUMS 67

HIII HIC Max HIC: 866.229 at Time: [0.073701 0.088700] 100 1000

80 800

60 600

40 400 Maximum HIC Head Resul Accel

20 200

0 0 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16

Pulse (Acceleration) vs Time Maximum Acceleration Min:[0.112200s,−23.3Gs] Maximum Acceleration Max:[0.068200s,346Gs] 350 x−accel 300

250

200

150

100

50 Pulse (Acceleration)−Res (Gs) 0

−50 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 0.2 Time (s)

Figure 3.19: HIII HIC15 vs. time Chapter 3. Development of Finite Element Injury Measures for THUMS 68

The HIC36 (shown in green) for THUMS is plotted on top of the resultant acceleration

(shown as blue points) to show the relationship between HIC36 and resultant acceleration. The instantaneous HIC36, calculated using 36 ms of data forward from the current time point (Figure 3.20) shows the maximum HIC beginning approximately 5 ms before the peak in the crash pulse. The HIC36 window (as shown by the beige band) includes the maximum acceleration. Chapter 3. Development of Finite Element Injury Measures for THUMS 69

Max HIC: 941.23 at Time: [0.063500 0.099401] 100 1000

50 500 Maximum HIC Head Resul Accel

0 0 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18

Pulse (Acceleration) vs Time Maximum Acceleration Min:[0.112200s,−23.3Gs] Maximum Acceleration Max:[0.068200s,346Gs] 350 x−accel 300

250

200

150

100

50 Pulse (Acceleration)−Res (Gs)

−50 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 Time (s)

Figure 3.20: THUMS HIC36 vs. time Chapter 3. Development of Finite Element Injury Measures for THUMS 70

The HIC36 (shown in green) for HIII is plotted on top of the resultant acceleration (shown as blue points) to show the relationship between HIC36 and resultant acceleration. The maximum HIC36 of 1570 is higher than the maximum THUMS HIC36 of 941 and occurs 5 ms after the peak for THUMS. Chapter 3. Development of Finite Element Injury Measures for THUMS 71

HIII HIC Max HIC: 1574.59 at Time: [0.068500 0.104401] 100 2000

50 1000 Maximum HIC Head Resul Accel

0 0 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16

Pulse (Acceleration) vs Time Maximum Acceleration Min:[0.112200s,−23.3Gs] Maximum Acceleration Max:[0.068200s,346Gs] 350 x−accel 300

250

200

150

100

50 Pulse (Acceleration)−Res (Gs) 0

−50 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 0.2 Time (s)

Figure 3.21: HIII HIC36 vs. time Chapter 3. Development of Finite Element Injury Measures for THUMS 72

3.3.2 Neck

The neck measures include normal force, shear force, bending moment, lateral moment, and Nij. These measures are all measured across C1. The forces and moments are direct measurements from a combination of the section plane and the discrete elements. Nij is calculated from the normal forces and bending moments, which makes it an indirect measurement that is a cumulative and accounts for multiple direct measurements. Chapter 3. Development of Finite Element Injury Measures for THUMS 73

The neck normal force for THUMS measures the tension or compression in cervical spin (Figure 3.22). Initially there is a small amount of compression, but approximately at 42 ms, the normal forces in the neck become tensile forces. The tension forces gradually increase after that and peaks to 1153 N at 90 ms, around 22 ms after the peak in the crash pulse. Chapter 3. Development of Finite Element Injury Measures for THUMS 74

Neck Normal Force vs Time Min:[0.035000s,−87.3332N] Max:[0.090000s, 1153.25N] 1200

1000

800

600

400

Neck Normal Force (N) 200

−200 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 0.2 Time (s)

Pulse (Acceleration) vs Time Maximum Acceleration Min:[0.112200s,−23.3Gs] Maximum Acceleration Max:[0.068200s,346Gs] 350 x−accel 300

250

200

150

100

50 Pulse (Acceleration)−Res (Gs)

−50 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 0.2 Time (s)

Figure 3.22: THUMS neck normal force vs. time Chapter 3. Development of Finite Element Injury Measures for THUMS 75

The neck normal force for HIII measures the tension or compression in the neck joint(Figure 3.23). Both the maximum compression and maximum tension forces are more than 3 times higher than the THUMS forces.

HIII Neck Normal Force vs Time Min:[0.042200s, −1647.3N] Max:[0.095901s, 4369.34N] 5000

4000

3000

2000

1000

0 HIII Neck Normal Force (N)

−1000

−2000 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 Time (s)

Pulse (Acceleration) vs Time Maximum Acceleration Min:[0.112200s,−23.3Gs] Maximum Acceleration Max:[0.068200s,346Gs] 350 x−accel 300

250

200

150

100

50 Pulse (Acceleration)−Res (Gs) 0

−50 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 0.2 Time (s)

Figure 3.23: HIII neck normal force vs. time Chapter 3. Development of Finite Element Injury Measures for THUMS 76

The neck shear force for THUMS measures the resultant shear across the section of the neck (Figure 3.24). This force is about 20% that of the normal force in peak magnitude and has a less uniform trend in the slope. Due to the fact that this is a resultant shear force, the shear in individual directions along the plane is not available. The shear force is not used in the calculation of Nij. Chapter 3. Development of Finite Element Injury Measures for THUMS 77

Neck Shear Force vs Time Min:[0.000000s, 0N] Max:[0.170000s, 224.42N] 250

200

150

100 Neck Shear Force (N)

0 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 0.2 Time (s)

Pulse (Acceleration) vs Time Maximum Acceleration Min:[0.112200s,−23.3Gs] Maximum Acceleration Max:[0.068200s,346Gs] 350 x−accel 300

250

200

150

100

50 Pulse (Acceleration)−Res (Gs)

−50 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 0.2 Time (s)

Figure 3.24: THUMS neck resultant shear force vs. time Chapter 3. Development of Finite Element Injury Measures for THUMS 78

The neck shear force for HIII measures the shear across the neck joint (Figure 3.25). The shear forces for the HIII are higher than those of the THUMS by a factor of 10.

HIII Neck Shear Force vs Time Min:[0.100501s,−2762.26N] Max:[0.043400s, 244.019N] 500

−500

−1000

−1500

−2000 HIII Neck Shear Force (N)

−2500

−3000 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 Time (s)

Pulse (Acceleration) vs Time Maximum Acceleration Min:[0.112200s,−23.3Gs] Maximum Acceleration Max:[0.068200s,346Gs] 350 x−accel 300

250

200

150

100

50 Pulse (Acceleration)−Res (Gs) 0

−50 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 0.2 Time (s)

Figure 3.25: HIII neck resultant shear force vs. time Chapter 3. Development of Finite Element Injury Measures for THUMS 79

Neck bending moment is the extension-flexion motion of the neck where extension is positive and flexion is negative (Figure 3.26). The neck experiences a gradual increase in flexion that peaks at 43 N·m approximately 35 ms after the peak in the crash pulse. There are two peaks in the flexion that is separate by about 20 ms.

Neck Bending Moment (CFC 600) vs Time Min:[0.105301s,−43.3556N*m] Max:[0.199902s, 4.57592N*m] 5

−5

−10

−15

−20

−25

−30

−35

Neck Bending Moment (CFC 600) (N*m) −40

−45 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 Time (s)

Pulse (Acceleration) vs Time Maximum Acceleration Min:[0.112200s,−23.3Gs] Maximum Acceleration Max:[0.068200s,346Gs] 350 x−accel 300

250

200

150

100

50 Pulse (Acceleration)−Res (Gs)

−50 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 Time (s)

Figure 3.26: THUMS neck bending moment vs. time Chapter 3. Development of Finite Element Injury Measures for THUMS 80

Neck bending moment for HIII is shown in Figure 3.27. The positive moment (extension) is approximately 10 times higher than the THUMS positive bending moment, but the negative moment (ﬂexion) is less than twice that of the THUMS positive bending moment. Chapter 3. Development of Finite Element Injury Measures for THUMS 81

HIII Neck Bending Moment vs Time Min:[0.059800s,−69.4054N*m] Max:[0.152702s, 49.2359N*m] 60

−20

−40

HIII Neck Bending Moment (N*m) −60

−80 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 Time (s)

Pulse (Acceleration) vs Time Maximum Acceleration Min:[0.112200s,−23.3Gs] Maximum Acceleration Max:[0.068200s,346Gs] 350 x−accel 300

250

200

150

100

50 Pulse (Acceleration)−Res (Gs) 0

−50 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 0.2 Time (s)

Figure 3.27: HIII neck bending moment vs. time Chapter 3. Development of Finite Element Injury Measures for THUMS 82

The neck lateral moment is the side to side motion of the head where rotation towards the occupant’s left is positive. The maximum lateral moment is approximately 20% of the maximum bending moment. The lateral moment is not used in the calculation of Nij.

Neck Lateral Moment (CFC 600) vs Time Min:[0.106201s,−9.46332N*m] Max:[0.088500s, 8.91718N*m] 10

−2

−4

−6

Neck Lateral Moment (CFC 600) (N*m) −8

−10 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 Time (s)

Pulse (Acceleration) vs Time Maximum Acceleration Min:[0.112200s,−23.3Gs] Maximum Acceleration Max:[0.068200s,346Gs] 350 x−accel 300

250

200

150

100

50 Pulse (Acceleration)−Res (Gs)

−50 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 Time (s)

Figure 3.28: THUMS neck lateral moment vs. time Chapter 3. Development of Finite Element Injury Measures for THUMS 83

The Nij combines the normal forces and bending moments to create a unitless measure that can be scaled to diﬀerent size bodies (Figure 3.29). The Nij shares a similar shape to the neck normal force plot (Figure 3.22) while the shape of the bending moments are less visible. The maximum Nij value of 0.17 is much less than the critical value of 1. Chapter 3. Development of Finite Element Injury Measures for THUMS 84

NIJ Neck Injury Criterion vs Time Min:[0.000000s,8.86236e−005Unitless] Max:[0.090000s,0.169627Unitless] 0.18

0.16

0.14

0.12

0.1

0.08

0.06

0.04 NIJ Neck Injury Criterion (Unitless) 0.02

0 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 0.2 Time (s)

Pulse (Acceleration) vs Time Maximum Acceleration Min:[0.112200s,−23.3Gs] Maximum Acceleration Max:[0.068200s,346Gs] 350 x−accel 300

250

200

150

100

50 Pulse (Acceleration)−Res (Gs)

−50 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 Time (s)

Figure 3.29: THUMS Nij vs. time Chapter 3. Development of Finite Element Injury Measures for THUMS 85

The Nij for HIII is shown in Figure 3.30. The maximum Nij for HIII is almost 5 times that of the maximum Nij for THUMS.

HIII NIJ Neck Injury Criterion vs Time Min:[0.000100s,7.08758e−006Unitless] Max:[0.059800s,0.749855Unitless] 0.8

0.7

0.6

0.5

0.4

0.3

0.2

HIII NIJ Neck Injury Criterion (Unitless) 0.1

0 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 Time (s)

Pulse (Acceleration) vs Time Maximum Acceleration Min:[0.112200s,−23.3Gs] Maximum Acceleration Max:[0.068200s,346Gs] 350 x−accel 300

250

200

150

100

50 Pulse (Acceleration)−Res (Gs) 0

−50 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 0.2 Time (s)

Figure 3.30: HIII Nij vs. time Chapter 3. Development of Finite Element Injury Measures for THUMS 86

3.3.3 Chest

Many chest injury measures were taken from THUMS, but only a few are shown here to illustrate the overall trends. For the other injury measures, refer to the appendix. Chest acceleration and deflection are standard injury measures that are studied often. Horizontal and parallel rib deflections are two different ways of measuring rib deflections that provides more information than sternal deflection. The shoulder force measures the amount of force the shoulder belt exerts on the shoulder. The rib failure plot shows the shell elements in the cortical bone that exceeded the failure stress of the rib. Chapter 3. Development of Finite Element Injury Measures for THUMS 87

The chest acceleration for THUMS is the maximum acceleration of T6. It is roughly correlated to the amount of force experienced by the chest. The chest acceleration curve is roughly similar to the crash pulse although the initial acceleration is less and the tail is higher. The maximum chest acceleration occurs approximately 5 ms before the maximum crash pulse. The acceleration data is also noisier than the crash pulse. Chapter 3. Development of Finite Element Injury Measures for THUMS 88

Chest Accel (CFC 180)−res vs Time Min:[0.000000s,0.000672948Gs] Max:[0.062900s, 48.3841Gs] 50

10 Chest Accel (CFC 180)−res (Gs)

0 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 Time (s)

Pulse (Acceleration) vs Time Maximum Acceleration Min:[0.112200s,−23.3Gs] Maximum Acceleration Max:[0.068200s,346Gs] 350 x−accel 300

250

200

150

100

50 Pulse (Acceleration)−Res (Gs)

−50 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 Time (s)

Figure 3.31: THUMS chest resultant acceleration vs. time Chapter 3. Development of Finite Element Injury Measures for THUMS 89

The chest acceleration for HIII is the maximum acceleration at the center of the second rib. The maximum acceleration for HIII is approximately 20 Gs higher than THUMS and occurs about 35 ms after the peak in THUMS chest acceleration. The chest acceleration for HIII is also more sustained than the THUMS chest acceleration and drops oﬀ slower. Chapter 3. Development of Finite Element Injury Measures for THUMS 90

Chest acceleration vs Time Min:[0.002100s,0.0507752Gs] Max:[0.097901s, 66.4213Gs] 70

20 Chest acceleration (Gs)

0 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 Time (s)

Pulse (Acceleration) vs Time Maximum Acceleration Min:[0.112200s,−23.3Gs] Maximum Acceleration Max:[0.068200s,346Gs] 350 x−accel 300

250

200

150

100

50 Pulse (Acceleration)−Res (Gs) 0

−50 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 0.2 Time (s)

Figure 3.32: HIII chest resultant acceleration vs. time Chapter 3. Development of Finite Element Injury Measures for THUMS 91

The chest deflection for THUMS is inward compression of the chest from the original chest shape. The deflection is reaches minimum approximately 70 ms after the peak in crash pulse, which is different from the acceleration in which the maximum occurred before the crash pulse.

Sternal Deflection (CFC 600) vs Time Min:[0.139202s,−5.39408cm] Max:[0.002400s,0.00083842cm] 1

−1

−2

−3

−4

Sternal Deflection (CFC 600) (cm) −5

−6 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 Time (s)

Pulse (Acceleration) vs Time Maximum Acceleration Min:[0.112200s,−23.3Gs] Maximum Acceleration Max:[0.068200s,346Gs] 350 x−accel 300

250

200

150

100

50 Pulse (Acceleration)−Res (Gs)

−50 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 Time (s)

Figure 3.33: THUMS chest sternal deﬂection vs. time Chapter 3. Development of Finite Element Injury Measures for THUMS 92

The chest deflection for HIII is inward compression of the chest as measured by a spring. The deflection for HIII is approximately 1 cm less than the deflection for THUMS and also peaks 60 ms earlier.

Chest Deflection vs Time Min:[0.080000s,−4.15063cm] Max:[0.000000s, 0cm] 0

−0.5

−1

−1.5

−2

−2.5

−3 Chest Deflection (cm) −3.5

−4

−4.5 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 Time (s)

Pulse (Acceleration) vs Time Maximum Acceleration Min:[0.112200s,−23.3Gs] Maximum Acceleration Max:[0.068200s,346Gs] 350 x−accel 300

250

200

150

100

50 Pulse (Acceleration)−Res (Gs) 0

−50 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 0.2 Time (s)

Figure 3.34: HIII chest deﬂection vs. time Chapter 3. Development of Finite Element Injury Measures for THUMS 93

The normalized chest deflection is the chest deflection divided by the initial chest depth. The normalized chest deflection reaches a minimum of 23%, after the crash pulse reaches a minimum.

Sternal Deflection (CFC 600) vs Time Min:[0.139202s,−23.2723%] Max:[0.002400s,0.0036173%] 5

−5

−10

−15

Sternal Deflection (CFC 600) (%) −20

−25 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 Time (s)

Pulse (Acceleration) vs Time Maximum Acceleration Min:[0.112200s,−23.3Gs] Maximum Acceleration Max:[0.068200s,346Gs] 350 x−accel 300

250

200

150

100

50 Pulse (Acceleration)−Res (Gs)

−50 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 Time (s)

Figure 3.35: Normalized sternal deﬂection vs. time Chapter 3. Development of Finite Element Injury Measures for THUMS 94

The horizontal rib deflection measures the deflection of individual ribs in a similar manner to the HIII deflection method. The ribs deflections are measured across lines horizontal to the ground even if they traverse different ribs. Ribs 2 and 3 experiences the highest inward compression while ribs 6 and 7 experiences the least inward compression. The maximum in rib deflection for rib 2 occurs approximate 35 ms after the peak in crash pulse. Ribs 6 and 7 experience little compression but more outward deflection of the rib. Chapter 3. Development of Finite Element Injury Measures for THUMS 95

Rib Deflection (CFC 600) vs Time Global Rib Deflection Min:[0.105801s,−4.94401cm] Global Rib Deflection Max:[0.145402s,2.88629cm] 3 Rib 1 2 Rib 2 Rib 3 Rib 4 1 Rib 5 Rib 6 0 Rib 7

−1

−2

−3

Rib Deflection (CFC 600)−Res (cm) −4

−5 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 Time (s)

Pulse (Acceleration) vs Time Maximum Acceleration Min:[0.112200s,−23.3Gs] Maximum Acceleration Max:[0.068200s,346Gs] 350 x−accel 300

250

200

150

100

50 Pulse (Acceleration)−Res (Gs)

−50 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 Time (s)

Figure 3.36: Horizontal left chest rib deﬂections vs. time Chapter 3. Development of Finite Element Injury Measures for THUMS 96

The parallel rib deflection measures the deflection of individual ribs along the same rib, which slope downwards from back to front. The rib deflections show that the minimum deflection occurs at a similar time point as the horizontal method of measuring chest deflection, around 100 ms. However, there is less variation in the shape of the curve for different ribs. There is expansion (positive deflection) on the tail of the curve rather than on both sides of the minimum as seen in the horizontal deflections. Chapter 3. Development of Finite Element Injury Measures for THUMS 97

Rib Deflection (CFC 600) vs Time Global Rib Deflection Min:[0.103501s,−4.74683cm] Global Rib Deflection Max:[0.160902s,3.30352cm] 4 Rib 1 3 Rib 2 Rib 3 2 Rib 4 Rib 5 1 Rib 6 Rib 7 0 Rib 8 Rib 9 −1 Rib 10

−2

−3 Rib Deflection (CFC 600)−Res (cm) −4

−5 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 Time (s)

Pulse (Acceleration) vs Time Maximum Acceleration Min:[0.112200s,−23.3Gs] Maximum Acceleration Max:[0.068200s,346Gs] 350 x−accel 300

250

200

150

100

50 Pulse (Acceleration)−Res (Gs)

−50 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 Time (s)

Figure 3.37: Parallel left check rib deﬂection vs. time Chapter 3. Development of Finite Element Injury Measures for THUMS 98

The resultant shoulder force is the force of the shoulder belt on THUMS shoulder. The shoulder force is approximately 1100 N at the highest point, which is 1 ms earlier than the max in the crash pulse. The shoulder force has a single peak at the point of highest force and tails oﬀ in both directions roughly linearly.

Shoulder (CFC 600) Force−res vs Time Min:[0.004801s, 1.92371N] Max:[0.067100s, 1143.11N] 1200

1000

800

600

400

Shoulder (CFC 600) Force−res (N) 200

0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 Time (s)

Pulse (Acceleration) vs Time Maximum Acceleration Min:[0.112200s,−23.3Gs] Maximum Acceleration Max:[0.068200s,346Gs] 350 x−accel 300

250

200

150

100

50 Pulse (Acceleration)−Res (Gs)

−50 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 Time (s)

Figure 3.38: Forces exerted on the THUMS shoulder Chapter 3. Development of Finite Element Injury Measures for THUMS 99

The rib fracture is measured as the failure of the shell elements that cover the rib cage, it is deﬁned as a function of the plastic strain on the element. The rib fracture pattern shows a concentration of failures on the left side of THUMS (right side of the graph). There are also other areas of failure such as near the left collar bone and the right lower ribs. Chapter 3. Development of Finite Element Injury Measures for THUMS 100

Rib 1 11 2 3 3

Rib 2

Rib 3 4

Rib 4 2 2 6

Rib 5 6

Rib 6 4

Rib 7

Rib 8 4 11

Rib 9 26

Rib 10 31 3

Rib Failure as a function of Normalized distance from sternum (−50% to 50%)

Figure 3.39: Chest predicted rib fracture locations Chapter 3. Development of Finite Element Injury Measures for THUMS 101

3.3.4 Abdomen

Some of the abdominal injury measures in this study include the abdominal acceleration, pelvic acceleration, lap belt force, and left iliac crest force. Abdominal acceleration measures the soft tissue acceleration while the pelvic acceleration measures the acceleration of the bony structure. Lap belt force and left iliac force measures the force the lap belt exerts on the abdomen and the iliac crest, respectively. Chapter 3. Development of Finite Element Injury Measures for THUMS 102

The abdominal acceleration is measured in the center of the abdomen part. The abdominal acceleration reaches a peak of approximately 41 Gs 35 ms before the peak in crash pulse in the x-direction. The x-direction is the primary direction of force and thus has a much higher magnitude than the other directions. The abdominal acceleration drops oﬀ drastically after the two initial peaks. Chapter 3. Development of Finite Element Injury Measures for THUMS 103

Pelvis Abdominal acceleration (CFC 1000)−x vs Time Pelvis Abdominal acceleration (CFC 1000)−y vs Time Min:[0.148102s,−11.6338Gs] Min:[0.056600s,−5.63614Gs] Max:[0.034000s, 41.3499Gs] Max:[0.141602s, 3.88433Gs] 50 4

3 40 2

30 1

0 20 −1 10 −2

0 −3 −4 −10 −5

Pelvis Abdominal acceleration (CFC 1000)−x (Gs) −20 Pelvis Abdominal acceleration (CFC 1000)−y (Gs) −6 0 0.05 0.1 0.15 0 0.05 0.1 0.15 Time (s) Time (s)

Pelvis Abdominal acceleration (CFC 1000)−z vs Time Pelvis Abdominal acceleration (CFC 1000) vs Time Min:[0.127501s,−7.00453Gs] Resultant Min:[0.000000s,4.88151e−006Gs] Max:[0.133301s, 8.91826Gs] Resultant Max:[0.034000s,41.3994Gs] 10 60 X 8 50 Y Z 6 40 4 30 2 20 0 10 −2 0 −4

−6 −10

Pelvis Abdominal acceleration (CFC 1000)−z (Gs) −8 −20 0 0.05 0.1 0.15 Pelvis Abdominal acceleration (CFC 1000)−Res (Gs) 0 0.05 0.1 0.15 Time (s) Time (s)

Figure 3.40: Abdominal acceleration vs. time Chapter 3. Development of Finite Element Injury Measures for THUMS 104

The pelvic acceleration for THUMS is measured as the average acceleration of all the nodes in the pelvic bone. The pelvic acceleration is noisier than the abdominal acceleration, it also has a higher maximum value of 76 Gs in the x-direction. But the maximum value is a sharp peak in the curve and the rest of the curve never goes higher than 50 Gs. The acceleration in the y and z direction are similarly noisy. Chapter 3. Development of Finite Element Injury Measures for THUMS 105

Pelvis Interped Pelvic Accel (CFC 1000)−x vs Time Pelvis Interped Pelvic Accel (CFC 1000)−y vs Time Min:[0.128201s,−45.2028Gs] Min:[0.026899s,−24.5269Gs] Max:[0.026300s, 76.2014Gs] Max:[0.026000s, 21.0599Gs] 80 25

20 60 15

40 10

5 20 0 0 −5

−20 −10 −15 −40 −20 Pelvis Interped Pelvic Accel (CFC 1000)−x (Gs) Pelvis Interped Pelvic Accel (CFC 1000)−y (Gs) −60 −25 0 0.05 0.1 0.15 0 0.05 0.1 0.15 Time (s) Time (s)

Pelvis Interped Pelvic Accel (CFC 1000)−z vs Time Pelvis Interped Pelvic Accel (CFC 1000) vs Time Min:[0.025900s,−58.2643Gs] Resultant Min:[0.000000s,0.000111801Gs] Max:[0.026300s, 58.4421Gs] Resultant Max:[0.026300s,96.1976Gs] 60 100 X 80 Y 40 Z 60

40 20 20

0 0

−20 −20 −40

−60 −40 −80 Pelvis Interped Pelvic Accel (CFC 1000)−z (Gs)

−60 Pelvis Interped Pelvic Accel (CFC 1000)−Res (Gs) −100 0 0.05 0.1 0.15 0 0.05 0.1 0.15 Time (s) Time (s)

Figure 3.41: THUMS pelvic acceleration vs. time Chapter 3. Development of Finite Element Injury Measures for THUMS 106

The resultant pelvic acceleration for THUMS is shown in Figure 3.42. The sharp peak in the x direction is also visible in the resultant.

Pelvis Interped Pelvic Accel (CFC 1000)−res vs Time Min:[0.000000s,0.000111801Gs] Max:[0.026300s, 96.1976Gs] 100

Pelvis Interped Pelvic Accel (CFC 1000)−res (Gs) 0 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 Time (s)

Pulse (Acceleration) vs Time Maximum Acceleration Min:[0.112200s,−23.3Gs] Maximum Acceleration Max:[0.068200s,346Gs] 350 x−accel 300

250

200

150

100

50 Pulse (Acceleration)−Res (Gs)

−50 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 Time (s)

Figure 3.42: THUMS resultant pelvic acceleration vs. time Chapter 3. Development of Finite Element Injury Measures for THUMS 107

The resultant pelvic acceleration for HIII is measured near center of the pelvic bone. The peak acceleration is close to the THUMS value of 96 Gs, but the average acceleration is higher.

Pelvic Acceleration vs Time Min:[0.002000s,−0.00966117Gs] Max:[0.042700s, 91.6422Gs] 100

20 Pelvic Acceleration (Gs)

−20 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 Time (s)

Pulse (Acceleration) vs Time Maximum Acceleration Min:[0.112200s,−23.3Gs] Maximum Acceleration Max:[0.068200s,346Gs] 350 x−accel 300

250

200

150

100

50 Pulse (Acceleration)−Res (Gs) 0

−50 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 0.2 Time (s)

Figure 3.43: HIII resultant pelvic acceleration vs. time Chapter 3. Development of Finite Element Injury Measures for THUMS 108

The lap belt force measures the force the lap belt exerts on the outer surface of THUMS. The lap belt force is much higher than the shoulder belt maximum force of 1100 N at 21 kN, almost 21 times higher. Both the lap belt and the shoulder belt shares similar shapes in the curve, with the lap belt forces being less noisy. The lap belt peak force also occurs at a similar time point to the shoulder belt (approximately 5 ms). Chapter 3. Development of Finite Element Injury Measures for THUMS 109

Pelvis Lap Belt (CFC 600) Force−res vs Time Min:[0.000001s, 4.07619N] 4 x 10 Max:[0.072600s, 21373N] 2.5

1.5

0.5 Pelvis Lap Belt (CFC 600) Force−res (N)

0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 Time (s)

Pulse (Acceleration) vs Time Maximum Acceleration Min:[0.112200s,−23.3Gs] Maximum Acceleration Max:[0.068200s,346Gs] 350 x−accel 300

250

200

150

100

50 Pulse (Acceleration)−Res (Gs)

−50 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 Time (s)

Figure 3.44: Abdominal lap belt force vs. time Chapter 3. Development of Finite Element Injury Measures for THUMS 110

The left iliac crest force is the amount of force the lap belt exerts on the left iliac crest. The force is approximately 1100 N at the highest point and is much less than the lap belt force. The force also tails oﬀ suddenly before 100 ms.

Pelvis Left Iliac Crest (CFC 600) Force−res vs Time Min:[0.012600s,0.407594N] Max:[0.060500s, 1079.5N] 1200

1000

800

600

400

200 Pelvis Left Iliac Crest (CFC 600) Force−res (N) 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 Time (s)

Pulse (Acceleration) vs Time Maximum Acceleration Min:[0.112200s,−23.3Gs] Maximum Acceleration Max:[0.068200s,346Gs] 350 x−accel 300

250

200

150

100

50 Pulse (Acceleration)−Res (Gs)

−50 0 0.02 0.04 0.06 0.08 0.1 0.12 0.14 0.16 0.18 Time (s)

Figure 3.45: Abdominal lap belt force on iliac crest vs. time Chapter 3. Development of Finite Element Injury Measures for THUMS 111

3.3.5 Derived measures

Certain injury measures are derived as single numbers from the output, these are shown in Table 3.7.

The 3 ms clip is the maximum chest acceleration that is sustained for more than 3 ms. It is shown in addition to the maximum chest acceleration because the maximum chest acceleration can be inaccurate due to noise.

The CTI is a unitless metric of the chest that combines the acceleration and the deﬂec- tion. A value higher than 1, as in the current results, is above the safe threshold deﬁned by the metric.

The maximum rib stress is the maximum stress experienced by the cortical bone (shell layer) of the ribs.

Table 3.7: Single number injury measures derived from the output.

Location Measure THUMS HIII 3 ms clip 42.715 Gs 52.028 Gs Chest CTI 1.0314 1.0661 Max rib stress 175 MPa NA Chapter 3. Development of Finite Element Injury Measures for THUMS 112

3.4 Discussion

The majority of the injury measures are within the tolerable threshold defined by NHTSA. The results show that there are many measures that do not have good correlation with the crash pulse, indicating that the effect of the impact is transferred to different parts of the body at different times.

3.4.1 Results

The abdominal acceleration peaks much earlier than the chest acceleration, suggesting that the lap belt engagement occurs earlier than the shoulder belt engagement. However, it might be due to this oﬀset in peak acceleration that causes the discrepancy of forces between the chest and the abdomen. While the magnitude of the acceleration were similar, the force of on the abdomen by the lap belt was a 21 times larger. There are a few factors that could cause this eﬀect. First, the early peaking of the lap belt could cause a higher force to develop in the lap belt due to longer contact time. THUMS is accelerated against the lap belt with much higher forces than against the shoulder belt as a result of the contact time. Second, the shoulder belt nodal force group only covers the shoulder and part of the THUMS chest while neglecting the lower portion of the chest. The shoulder nodal force group was designed to measure the amount of force exerted against the shoulder and upper chest portion of the shoulder belt, and not across the entire belt length. It was anticipated that the bulk of the force exerted would be on the shoulder or upper chest, but this remains an unknown and further studies should incorporate measurement of the lower chest as well. Third, the should belt was anchored to a retractor that allowed belt payout while the lap belt was anchored to a node that constrained the seatbelt to it. While the slip ring in the buckle and the shoulder D-ring allowed belt payout to propagate throughout the belt system, it was hindered by the slip ring friction and redirection. Slip ring redirection required a belt element to be stretched by approximately 40% before the element is propagated to the other side. Thus the belt could have much more leeway in the shoulder portion than the lap portion due to payout to the shoulder belt. Chapter 3. Development of Finite Element Injury Measures for THUMS 113

The iliac crest force peaking before the lap belt force indicates that the seatbelt engages the iliac crest before it engages the abdomen. This is indicates that the seat belt is positioned well and engaging the pelvic bone as designed to prevent excessive with the soft abdomen. The magnitude of the iliac crest force is much lower than that of the lap belt force, however. The iliac crest nodal force group is much smaller than the lap belt nodal force group, so the actual stress distribution could be diﬀerent. So that the ultimately the amount of force experienced by the abdomen is still much higher than the pelvic bone.

The Nij value peaks later than both the chest and abdominal acceleration at 90 ms. Although the Nij value only comes to within 20% of the critical value of 1, there are several factors that could affect the Nij value. First, the Nij value is measured from a suboptimal position of C1 of THUMS rather than the occipital condyle. Given the differences in geometry between THUMS and HIII, it is difficult to judge how to effectively imitate the injury measures of HIII with THUMS as the occipital condyle forces and moments were shown to be difficult to measure. Second, the various discrete elements interfered with the correct measurement of forces and moments due to the arbitrary ori- entation of these elements and the sheer number of them. In addition to accounting for the forces and moments at the C1 plane, these discrete elements had to be accounted for, which is another deviation of the measurement methodology of THUMS. The forces and moments are also further distorted by the change in position of these spring elements throughout the simulation, which must be tracked by the tissue they are attached to. Since no automated way was developed to track more 200 of these discrete elements, their positions was always assumed to be constant relative to the neck. Third, the lack of airbag and instrument panel would create a different motion of the neck than if the head had impacted an airbag. The forces and moment traces would be different in magnitude. The component of the Nij that contributed most is the neck tension force. A head to airbag impact could result in high compressive forces or bending moments in the neck.

The peak in head acceleration occurs at almost the same time as the peak in neck tension, which suggests that the head acceleration and neck tension are connected through mostly rigid structures. In addition, since most of the internal structures of the head are rigid Chapter 3. Development of Finite Element Injury Measures for THUMS 114 bodies, propagation of stress into the brain is not modeled correctly by this model and the acceleration is merely a reﬂection of the outer surface of the head.

Most of the HIII results show a more sustained injury measure that falls oﬀ slower than the THUMS. This indicates that the HIII is better at distributing the kinetic energy over a greater time interval than the THUMS.

3.4.2 Injury measures

There are multiple ways to measure the injury measure, and each could prove to be more comparable in some ways to ATD injury measures.

Using ATD injury measures can be a draw back for geometrically accurate models such as THUMS as the injury measures for simplified ATDs do not always correspond directly to a measure in THUMS. Even if there is a direct correspondence, the underlying assumptions for that injury measure may be less relevant. Furthermore, most often injury measures for simplified ATDs are used as a surrogate measure because the real measure of interest would be difficult or impossible to measure. Ideas such as shear stresses and maximum tensile stresses are probably more relevant to tissue failure, but since these would be difficult to measure in an ATD, a less relevant but more simple measure is taken. Nevertheless, development of injury measures for a geometrically accurate FE model would require much research and time as well as experimental data and are beyond the scope of this study. The injury measure focused here are meant to correlate existing injury measures to the THUMS model, thus allow some degree of comparability to previous literature and established limits. A more important use for the injury measures will be for relative comparisons between different simulation scenarios utilizing THUMS, as an absolute value for the injury measures might be less comparable due to geometry.

The HIC as shown by the HIII dummy is higher by approximately 400 for HIC15 and

500 for HIC36. This large deviation between the HIII and THUMS suggests a deviation of behavior in the head that could be attributed to multiple sources. The head material for THUMS and HIII are both mostly rigid bodies with deviations in geometry, but Chapter 3. Development of Finite Element Injury Measures for THUMS 115 the overall response is more likely to be different due to the design of the HIII as matching that of a physical ATD rather than a human body. The head acceleration graphs (Figure 3.16 and Figure 3.17 show that the head acceleration of THUMS is lower than that of HIII as well as having less peaks, which suggests a stiffer response by the HIII to the crash pulse. Due to the many differences in geometry in the upper body between the THUMS and HIII it is difficult to assign a single source that creates the variation in the outputs. The more prominent sources are differences in the neck, which through direct connection to the head, has a large effect on the response of the head. The THUMS neck not only consists of many interconnected parts, but also has springs and dampers to correct the behavior of the neck and simulate muscle behavior. The HIII neck, on the other hand, consists of a series of discs that each simulates a single cervical disk that bears the full transmission of any force through the neck. This makes the means of force transmission different between the THUMS and HIII.

The HIC injury measure is especially controversial as it is difficult to interpret simply. Gadd offered an intuitive interpretation of the SI initially, it might be that the SI still serves the purpose of offering some insight into the creation of injury measures for biological tissues. While interpretation of HIC such as those proposed by Eppinger [66] are useful, they do not correlate to the actual development of HIC. They are attempts to explain HIC after its development rather than to provide a guide for its actual development. Thus it can be seen why HIC would be a controversial injury measure. Perhaps a better injury measure would take into account differing parts of the head, instead of a single acceleration vs. time trace. Such a measure could use skull measurements for skull fracture term, brain measurements for brain injury term, and perhaps an interaction term that accounts for the interaction of the brain within the skull. Regard- less of how an improvement to HIC can be implemented, it is desirable to incorporate more measurements into the head so that more information is available for a new injury measure. While this may mean more equipment in an ATD, in a FE simulation, it would be trivial to gain additional measurements. Such ease of measurement have prompted development of FE based injury measures for brain injury [74] such as the Cumulative Strain Damage Measure (CSDM), Dilatation Damage Measure (DDM), and Chapter 3. Development of Finite Element Injury Measures for THUMS 116 the Relative Motion Damage Measure (RMDM) in the development of Simulated Injury Monitor (SIMon).

Since THUMS brain is formulated with a rigid body deﬁnition, such that element processing is not done for the part, and no deformation or relative motion occurs in the THUMS brain. This means that some accuracy is sacriﬁced in the brain. Brain injury research has been shown to be a complex subject on its own. On the other hand, the skin and outer soft tissue are formulated with elastic and viscoelastic materials, so that deformation, stress, and strain can be measured for the outer head. This could help identify the location of impact if the THUMS head came into contact with an interior component.

The chest deflection proved to offer different information than the chest acceleration. An advantage of using chest deflection over chest 3 ms clip is that the deflection values are not only proportional to the compressive load experienced by the ATD, but are also affected by the material properties of the chest itself. A person with more elastic ribs will experience more chest deflection than a person with harder ribs. As a result, the person with more elastic ribs will likely exert more forces on their internal organs, assuming that the ribs do not break. Since the chest 3 ms clip cannot account for the material properties, it will predict the same risk of injury. Chest deflection can also be correlated more directly to strain in the ribs and hence more to the risk of rib fracture.

The THUMS chest deflections peaks approximately 60 ms after the peaks in HIII chest deflection, indicating a stiffer response by the HIII. The THUMS chest deflections are also greater by approximately 30% compared to the HIII. The THUMS chest ribs are more pliable than the HIII could indicate that the HIII has stiffer ribs possibly to reduce the error from not having chest organs, which might not work well for this crash pulse.

Since the THUMS model is geometrically more accurate, the method of deflection measurement could be less meaningful since the deflection measurement for dummies are meant for a simplified model. For THUMS, rib and the parts underneath the ribs can be measured directly for loads, whereas the HIII dummy is restricted by the measurement instrumentation. If the intent of the rib deflection measurement is meant to detect rib fracture as a result of over bending of a rib, the parallel method might be more Chapter 3. Development of Finite Element Injury Measures for THUMS 117 relevant. On the other hand, if the intent of the rib deflection measurement is to detect underlying organ injuries as a result of deflection of the chest in a specific area, then the horizontal deflection measurements could be more relevant.

The CTI has a drawback in that it assumes an independent linear relationship between tolerance for deflection and tolerance for acceleration. With higher forces exerted against the chest, there is usually larger deflection. The CTI value is meant to offer an injury criteria for a seatbelt and airbag setup, and may not be as generalizable as the chest deflection or chest acceleration injury criteria. Chapter 3. Development of Finite Element Injury Measures for THUMS 118

3.5 Conclusion

Injury measures correlating to current HIII injury measures was developed for the THUMS. It was found that in a standard 40 mph frontal crash simulation, most of the injury measures were within the critical limits with the exception of the CTI. Transla- tion of ATD injury measures for FE models do not take full advantage of the power of FE modeling, but can correspond to validated measures. Chapter 4

Injury Risk Functions for THUMS

4.1 Introduction

The behavior of human tissue under extreme loading conditions such as a MVC is often difficult to measure accurately and directly. The difficulty of the measurement often makes ATDs with built in instrumentation valuable. They offer a means of correlate existing kinematics in human tissue to a quantitative measure. While correlation to real human tissue is ideal, there is often a wide variability in human tissue response and a difficulty in creating predictable comparisons. On the other hand, ATD behavior is repeatable and also less variable. The current ATDs must have a narrow range of response to the same crash pulse . In addition, current traffic safety laws and risk of injury functions are based on the effects of a crash on ATDs. Recent development in numerical modeling also more detailed measurements of FE human body surrogates. These injury measures also must correlate to existing ATD injury measures as well as human tissue. Therefore it is also desirable to match FE injury measures to ATD injury measures.

Injury measures are measurements that oﬀers some correlation to injury in a human body. While they represent independent values and can be measured in any human body

119 Chapter 4. Injury Risk Functions for THUMS 120 surrogate provided that adequate measurement options are available, the values may not be comparable between different human body surrogates. Since the injury risk functions [55, 56, 75] developed are used most often for the HIII dummy, it is important to evaluate the accuracy of the injury risk function when a different human body surrogate is used. The THUMS is a FE model that offers anatomical accuracy of the body in the skeletal structure and the soft tissue. Inner organs for the current version (1.62c) are not modelled, but are planned for the future. THUMS offer a drastically different anatomical structure compared to an HIII FE model due to its adherence to anatomical accuracy rather than the HIII dummy adherence to simplicity. As a result, it is difficult to gauge if all of the injury risk functions that are normally applicable to the HIII dummy are comparable with the THUMS. While THUMS has been validated in multiple load configurations, the injury measurement are often ambiguously defined and not all the injury measures were validated. Since the primary purpose of the human body surrogate is to correlate the injury to the risk, then it becomes necessary to apply the injury risk functions for current ATDs to the THUMS model by verifying the comparability between THUMS injury measures and ATD injury measures.

There are many human body surrogate models in development currently, but standard measures for injury prediction typically still revolves around those injury measures currently set for ATDs. While it is possible to use the standard ATD injury measures on new models for the human body, it is difficult to ascertain the comparability of the models. Validation should be done with the same impact scenario but use of the HIII FE model for new FE human body surrogates. In addition, the current FE models sometime implement novel injury measures, which must be validated directly with cadaver data or volunteer. The ideal injury measure will also offer a well established injury risk function that predicts probability of injury for different levels of the injury measure without being overly affected by individual specific characteristic, such as mass, height, or body shape. Such normalized injury measure can then be used by an injury risk function that’s more accurate for a larger population. Such as the chest deflection criterion, which reports an injury risk in terms of a standard distance. But a person with a larger chest depth could be less susceptible to injury than a person with smaller chest depth with the same Chapter 4. Injury Risk Functions for THUMS 121 deflection. However, the normalized chest deflection criterion accounts for this by dividing the deflection by the total chest depth, thus minimizing the effect of the person to person variation. The normalization may never be perfect, in fact if the chest deflection is used to assess the risk of rib fractures, normalizing by the chest depth may not be the best normalization technique because the ribs bending are also affected by the width of the chest. It would seem reasonable that a person with a wider chest can suffer more deflection in the sternum because the lever arm for bending is longer.

Injury measures for detailed human body surrogates can also be divided into global and local injury measures, global injury measures include such measures as head acceleration, chest acceleration, etc. Local injury measures include such measures as stress and strain at specific parts of the brain, rib maximum stress, and pressure inside the head. Global injury measures treats different components of the body as a whole, such as head acceleration is meant to represent the acceleration of the head as a rigid body, even if the head acceleration is measured from a specific site and not an averaged value across the whole head. On the other hand, local injury measures tends to break up components of the body into different regions or layers, and creates an injury measure that is intended to assess the risk of injury at a specific site rather than the whole component. The idea of the global injury measure is applicable to the ATD, as it contains simplified components of the body and therefore attempts to extrapolate some of its measures to all the body. The local injury measure is more applicable to numerical models or cadaver experimentation, where the study can focus on measuring a specific part of the model or cadaver with much more localization than is possible with an ATD. While local injury measures can be often illuminating, they are not as well validated because of less literature values to compare to as well as the increased difficulty in measurements depending on the model composition. There is also less agreement on the use of local injury measures. On the other hand, global injury measures are easily measured on an ATD, but are difficult to measure in actual cadavers because they are meant to represent an average value and cadaver measurements are measured at a single site.

Therefore there are variabilities in the injury measures themselves that are diﬃcult to account for. Thus a simple pass or fail injury risk, such as an IARV, would be controversial to develop that can be declared accurate, although they do serve as a Chapter 4. Injury Risk Functions for THUMS 122

Table 4.1: IARV values for select injury criterion

Injury measure IARV value Head acceleration 60 Gs HIC15 700 HIC36 1000 Nij 1 Chest acceleration 60 Gs Chest deﬂection 60 mm CTI 1

baseline for deciding what is acceptable. A more sophisticated injury risk measure is to utilize probability theory and predict a chance of injury instead of the pass or fail criterion. The idea of the probability of injury comes from the distribution function. It is assumed that the risk of injury starts of at a minimal level that is close to 0%. As the injury metric rises, the risk of injury also increases according to the type of distribution chosen. For the commonly used distributions, there is often a range of values in which an increase in the injury measure has a dramatic effect on the risk of injury. Below or above the range, and the injury measure has less and less of an effect on the risk of injury. The risk of injury is often correlated to the Abbreviated Injury Scale (AIS), or some other standardized measure of injury, such as rib fractures. The risk of injury can be a less intuitive measure of injury as it lacks the deterministic results that are found in pass or fail injury risk criteria. However, they are a more generalized way of representing injury risk as there’s always an inherent uncertainty in the differences in the human body between different people. Injury risk functions are developed using the results of cadaver data. Since the result of cadaver data’s response is often deterministic, such as injury or no injury, the use of multiple data points are used to fit the distribution of injury to a probability distribution.

Currently there are probability distributions that are used to assess injury risk. Initially the normal distribution (Equation 4.1) was used to assess the probability of injury, as the normal distribution is widely used in statistics. The original HIC injury risk curves by Mertz were analyzed with a normal injury risk function . With more development in injury risk analysis, other distributions were also utilized. The Lognormal distribution Chapter 4. Injury Risk Functions for THUMS 123

(Equation 4.2), the Weibull distribution (Equation 4.3), and the Logistic distribution (Equation 4.4) have all been used as the injury risk function.

2 − x e 2 f(x) = √ (4.1) 2π 2 − (ln x) e 2σ2 f(x) = √ x ≥ 0; σ > 0 (4.2) xσ 2π γ f(x) = γx(γ−1)e−x x ≥ 0; γ > 0 (4.3)

−(x−m) e b f(x) = −(x−m) b 6= 0 (4.4) b[1 + e b ]2

These parametric injury risk functions are often useful because they assume a distribution of the injury risk and allow a definition of the injury risk function based on a few coefficients. Furthermore, parametric injury risk functions such as the Logistic function are easier to fit and have well defined shapes and curve fitting procedures. The Logistic function itself is especially useful due to the fact that it allows the use of a closed form Cumulative Distribution Function (CDF) to describe the injury risk function. While parametric injury risk functions assume a shape for the CDF that do not always correlate well to the actual shape, the Logistic CDF has been used extensively in automotive injury risk functions and is used to correlate injury measures to risk of injury. This provides a standard platform for injury risk analysis that can be used to validate the THUMS injury risk function.

This study will utilize the current injury risk functions to analyze the eﬀectiveness of injury risk functions for THUMS. While the injury risk functions have only been validated for the HIII, the THUMS injury measures have been developed to attempt to correlate well with the HIII injury measures. Chapter 4. Injury Risk Functions for THUMS 124

4.2 Methods

The THUMS injury measures which injury risk functions are assessed for are the HIC36,

HIC15, Nij, chest 3 ms clip, chest sternal deﬂection, and CTI. The injury risk functions are taken from existing literature and have been validated with current ATDs as well as PMHS. While the ideal validation criteria for injury risk functions would be to validate directly against PMHS, it would be less repeatable and also require physical experimentation. In addition, since the ATD used is a FE model, the same method for calculating the injury risks can be applied to both the THUMS and the HIII model.

The injury risk curves in this study all come from existing literature and some are part of NHTSA regulations. The injury risk functions in this study comes from various sources and all utilize a logistic distribution . The CDF form of the logistic distribution is shown in Equation 4.5, where m is the location parameter and b is the scale. m is sometimes also written as µ, which correlates to the mean of the distribution.

1 F (x) = −(x−m) (4.5) 1 + e b

Most injury risk functions are defined at a set AIS level. AIS is a system of categorizing the severity of injury. The AIS is a 7-digit code that uniquely identifies an injury based on its body region, anatomic structure, specific site, specific injury name, and severity.

Table 4.2: AIS code description for the AIS severity number

AIS Code Description 1 Minor 2 Moderate 3 Serious 4 Severe 5 Critical 6 Maximum

The injury risk functions for the head include the HIC36 as well as the newer HIC15. The

HIC36 injury risk function only accounts for AIS 4+ injuries while the HIC15 accounts Chapter 4. Injury Risk Functions for THUMS 125 for AIS levels between one and fatal. A caveat of the HIC injury measure in general is that the duration is meant to account for an actual impact between the head and the interior component. In this case, neither the dashboard nor the steering wheel are modelled, and hence it would be diﬃcult for the head to come into contact with the interior component. In this case, the HIC36 could provide a better predictor of head injury as it is more suited for long duration head acceleration while the HIC15 is suited for short duration head impact against hard vehicle structures. The head injury risk functions for the HIC36 comes from Viano’s formulation while the risk functions for the

HIC15 comes from the Expanded Prasad/Mertz curves . The injury risk curve for HIC36 is shown in Equation 4.6 and Figure 4.1.

1 Risk(AIS 4+) = (4.6) 1 + e5.02−0.00351HIC36

0.9

0.8

0.7

0.6

0.5

0.4

0.3 Probability of Head Injury

0.2

0.1 AIS 4+ 0 0 500 1000 1500 2000 2500 3000 HIC (Gs) 36

Figure 4.1: Injury risk function for HIC36, based on Viano’s formulation

The injury risk function for HIC15 is shown in Equation 4.7 - Equation 4.12. The

HIC15 injury risk function does not follow the form of the standard Logistic CDF, but has an extra term ( 200 ). The injury risk function is based on the Expanded Prasad/Mertz HIC15 Chapter 4. Injury Risk Functions for THUMS 126 curves that originally measured only one AIS level and was later expanded by NHTSA to all the AIS levels . The risk curves are shown in Figure 4.2.

1 Risk(MAIS1+) = 200 (4.7) [(1.54+ −0.0065HIC15)] 1 + e HIC15 1 Risk(MAIS2+) = 200 (4.8) [(2.49+ −0.00483HIC15)] 1 + e HIC15 1 Risk(MAIS3+) = 200 (4.9) [(3.39+ −0.00372HIC15)] 1 + e HIC15 1 Risk(MAIS4+) = 200 (4.10) [(4.90+ −0.00351HIC15)] 1 + e HIC15 1 Risk(MAIS5+) = 200 (4.11) [(7.82+ −0.00429HIC15)] 1 + e HIC15 1 Risk(F atal) = 200 (4.12) [(12.24+ −0.00565HIC15)] 1 + e HIC15

0.9

0.8

0.7

0.6

0.5

0.4

0.3 MAIS 1+ Probability of Head Injury MAIS 2+ 0.2 MAIS 3+ MAIS 4+ 0.1 MAIS 5+ MAIS 6 0 0 500 1000 1500 2000 2500 3000 HIC (Gs) 15

Figure 4.2: Injury risk function for HIC15, based on Expanded Prasad/Mertz curves

The injury risk function for Nij is shown in Equation 4.13 and plotted in Figure 4.3. Although multiple AIS levels exist for the Nij in FMVSS Supplemental Notice of Pro- posed Rule Making (SNPRM) 1999, only the AIS 4+ injury risk level is evaluated in Chapter 4. Injury Risk Functions for THUMS 127 this study. The AIS 4+ level is chosen because it is the most often used injury risk assessment for single AIS level risk curves. It also oﬀers a balance in predicting the severity of the injury. Multiple AIS level injury risk functions were not utilized due to the fact that there are some inconsistencies with the Nij injury risk function for AIS 3+, which overtakes the other more severe AIS levels at an Nij value of 1.5.

1 Risk = (4.13) (AIS4+) 1 + e(2.693−1.196Nij)

0.9

0.8

0.7

0.6

0.5

0.4

0.3 Probability of Neck Injury

0.2

0.1 AIS 4+ 0 0 1 2 3 4 5 Nij (unitless)

Figure 4.3: Injury risk function for Nij, based on FMVSS SNPRM 1999

The injury risk function for chest 3ms clip is shown in Equation 4.14 - Equation 4.17 and plotted in Figure 4.4. The injury risks are from FMVSS SNPRM 1999 and represent the chance of injury for a 50% male. The injury risk for chest acceleration via the 3 ms clip is useful to understand the amount of force experienced by the chest. Chapter 4. Injury Risk Functions for THUMS 128

1 Risk(AIS2+) = (4.14) 1 + e(1.2324−0.0576a3ms) 1 Risk(AIS3+) = (4.15) 1 + e(3.1493−0.063a3ms) 1 Risk(AIS4+) = (4.16) 1 + e(4.3425−0.063a3ms) 1 Risk(AIS5+) = (4.17) 1 + e(8.7652−0.0659a3ms)

0.9

0.8

0.7

0.6

0.5

0.4

0.3 Probability of Chest Injury

0.2 AIS 2+ AIS 3+ 0.1 AIS 4+ AIS 5+ 0 0 50 100 150 200 250 Chest 3ms clip (Gs)

Figure 4.4: Injury risk function for chest 3ms clip, based on FMVSS SNPRM 1999

The injury risk function for chest sternal deﬂection is shown in Equation 4.18 - Equa- tion 4.21 and plotted in Figure 4.5. The injury risk for sternal deﬂection is often associated with the risk of rib fractures and less with regards to soft tissue injury . Chapter 4. Injury Risk Functions for THUMS 129

1 Risk = (4.18) (AIS2+) 1 + e(1.8706−0.04439D) 1 Risk = (4.19) (AIS3+) 1 + e(3.7124−0.0475D) 1 Risk = (4.20) (AIS4+) 1 + e(5.0952−0.0475D) 1 Risk = (4.21) (AIS5+) 1 + e(8.8274−0.0459D)

0.9

0.8

0.7

0.6

0.5

0.4

0.3 Probability of Chest Injury

0.2 AIS 2+ AIS 3+ 0.1 AIS 4+ AIS 5+ 0 0 50 100 150 200 250 300 Sternal Deflection (mm)

Figure 4.5: Injury risk function for chest sternal deﬂection, based on FMVSS SNPRM 1999

The injury risk function for CTI is shown in Equation 4.22 - Equation 4.25 and shown in Figure 4.6. The CTI is an injury measure that combines the sternal acceleration and deflection to account for injuries caused by excessive force or excessive deflection. The CTI hinges on assumption that deflection and forces should be balanced. While this seems like a reasonable assumption, it is also important to realize that often times forces and deflections are positively correlated. So that CTI does not combine independent measures. CTI promotes balanced forces and deflection by explicitly restricting the maximum allowable force and maximum allowable deflection. Consider the case of deflection without acceleration, this would be similar to a static loading of the thorax Chapter 4. Injury Risk Functions for THUMS 130 against the belt or the steering wheel at a constant speed, a crushing type injury. On the other hand, the case of acceleration without deflection would be similar to a sudden and quick impact on the chest, a blast type injury. CTI allows both of these types of injury as opposite ends of the triangle.

1 Risk = (4.22) (AIS2+) 1 + e(4.847−6.036CTI) 1 Risk = (4.23) (AIS3+) 1 + e(8.224−7.125CTI) 1 Risk = (4.24) (AIS4+) 1 + e(9.872−7.125CTI) 1 Risk = (4.25) (AIS5+) 1 + e(14.242−6.589CTI)

0.9

0.8

0.7

0.6

0.5

0.4

0.3 Probability of Chest Injury

0.2 AIS 2+ AIS 3+ 0.1 AIS 4+ AIS 5+ 0 0 0.5 1 1.5 2 2.5 3 Combined Thoracic Index (CTI − no units)

Figure 4.6: Injury risk function for CTI based on FMVSS SNPRM 1999 Chapter 4. Injury Risk Functions for THUMS 131

4.3 Results

The HIC15 risks of injury is shown in Figure 4.7 for THUMS and Figure 4.8 for HIII. The risk of injury for THUMS is much less than the risk of injury for HIII as can be seen by the stacked bar graphs. While the THUMS has a negligible risk of MAIS 5+ injury, the HIII has approximately 2% chance of MAIS 5 + injury. The diﬀerence is more drastic with MAIS2+ and MAIS3+ risks with a 45% and 27% diﬀerence, respectively.

Fatal MAIS5+ MAIS4+ MAIS3+ MAIS2+ MAIS1+

Buck 3Pt No L. Limit

0% 10% 20% 30% 40% 50% 60% 70% 80% 90% 100%

Figure 4.7: THUMS HIC15 risk of injury

Fatal MAIS5+ MAIS4+ MAIS3+ MAIS2+ MAIS1+

HIIIOutput

0% 10% 20% 30% 40% 50% 60% 70% 80% 90% 100%

Figure 4.8: HIII HIC15 risk of injury

Table 4.3: HIC15 injury risk comparison between THUMS and HIII

Fatal MAIS5+ MAIS4+ MAIS3+ MAIS2+ MAIS1+ THUMS <1% <1% 3% 13% 37% 77% HIII <1% 2% 11% 40% 82% 98% Chapter 4. Injury Risk Functions for THUMS 132

The HIC36 risks of injury is shown in Figure 4.9 for THUMS and Figure 4.10 for HIII. The risk of injury for THUMS is much less than the risk of injury for HIII as similar to HIC15. Since the formulation for the injury risk function for HIC36 is diﬀerent than the formulation for HIC15, the values are not comparable to HIC15.

AIS 4+

Buck 3Pt No L. Limit

0% 10% 20% 30% 40% 50% 60% 70% 80% 90% 100%

Figure 4.9: THUMS HIC36 risk of injury

AIS 4+

HIIIOutput

0% 10% 20% 30% 40% 50% 60% 70% 80% 90% 100%

Figure 4.10: HIII HIC36 risk of injury

Table 4.4: HIC36 injury risk comparison between THUMS and HIII

AIS 4+ THUMS 14% HIII 63% Chapter 4. Injury Risk Functions for THUMS 133

The Nij risk of injury is shown in Figure 4.11 for THUMS and Figure 4.12 for HIII. The risk of injury for THUMS is much less than the risk of injury for HIII as similar to the HIC.

AIS 4+

Buck 3Pt No L. Limit

0% 10% 20% 30% 40% 50% 60% 70% 80% 90% 100%

Figure 4.11: THUMS Nij risk of injury

AIS 4+

HIIIOutput

0% 10% 20% 30% 40% 50% 60% 70% 80% 90% 100%

Figure 4.12: HIII Nij risk of injury

Table 4.5: Nij injury risk comparison between THUMS and HIII

AIS 4+ THUMS 8% HIII 13% Chapter 4. Injury Risk Functions for THUMS 134

The chest acceleration (3 ms clip) risk of injury is shown in Figure 4.13 for THUMS and Figure 4.14 for HIII. Similar to previous risks of injury, the risk of injury for HIII is greater than the risk of injury for THUMS.

AIS 5+ AIS 4+ AIS 3+ AIS 2+

Buck 3Pt No L. Limit

0% 10% 20% 30% 40% 50% 60% 70% 80% 90% 100%

Figure 4.13: THUMS chest 3ms clip risk of injury

AIS 5+ AIS 4+ AIS 3+ AIS 2+

HIIIOutput

0% 10% 20% 30% 40% 50% 60% 70% 80% 90% 100%

Figure 4.14: HIII chest 3ms clip risk of injury

Table 4.6: Chest 3ms clip injury risk comparison between THUMS and HIII

AIS 5+ AIS 4+ AIS 3+ AIS 2+ THUMS < 1% 16% 39% 78% HIII < 1% 31% 59% 88% Chapter 4. Injury Risk Functions for THUMS 135

The chest deﬂection risk of injury is shown in Figure 4.15 for THUMS and Figure 4.16 for HIII. Unlike previous injury risks, the risk of injury for HIII is less than the risk of injury for THUMS.

AIS 5+ AIS 4+ AIS 3+ AIS 2+

Buck 3Pt No L. Limit

0% 10% 20% 30% 40% 50% 60% 70% 80% 90% 100%

Figure 4.15: THUMS chest deﬂection risk of injury

AIS 5+ AIS 4+ AIS 3+ AIS 2+

HIIIOutput

0% 10% 20% 30% 40% 50% 60% 70% 80% 90% 100%

Figure 4.16: HIII chest deﬂection risk of injury

Table 4.7: Chest deﬂection injury risk comparison between THUMS and HIII

AIS 5+ AIS 4+ AIS 3+ AIS 2+ THUMS < 1% 7% 23% 63% HIII < 1% 4% 15% 49% Chapter 4. Injury Risk Functions for THUMS 136

The CTI risk of injury is shown in Figure 4.17 for THUMS and Figure 4.18 for HIII. Similar to most other injury risks, the risk of injury for HIII is more than the risk of injury for THUMS.

AIS 5+ AIS 4+ AIS 3+ AIS 2+

Buck 3Pt No L. Limit

0% 10% 20% 30% 40% 50% 60% 70% 80% 90% 100%

Figure 4.17: THUMS CTI risk of injury

AIS 5+ AIS 4+ AIS 3+ AIS 2+

HIIIOutput

0% 10% 20% 30% 40% 50% 60% 70% 80% 90% 100%

Figure 4.18: HIII CTI risk of injury

Table 4.8: CTI injury risk comparison between THUMS and HIII

AIS 5+ AIS 4+ AIS 3+ AIS 2+ THUMS < 1% 7% 29% 80% HIII < 1% 9% 35% 83% Chapter 4. Injury Risk Functions for THUMS 137

4.4 Discussion

NHTSA defines the injury limit for 35 mph frontal crashes to be 10% chance of serious injury (AIS 3+), utilizing the HIII ATD. While this study simulates a 40 mph frontal crash, the HIII values all exceed the 10% threshold, for the injury risk curves that have it. There are several factors that limits the interpretation of the results. First, the focus of this study is the relative difference on safety system variations, not the absolute safety performance of the interior components. Thus it is not guaranteed that the interior components would have met NHTSA frontal standard with the HIII ATD at the standard crash pulse. The vehicle interior simulated does not have an airbag, which can further increase the chance of injury for frontal crashes when complementing the seatbelt. Second, the risk of injury is designed for a physical ATD and not a simulated one. Although the simulated HIII ATD conforms to the physical dummy in a few validation tests, the behavior is not guaranteed for all conditions. Third, the non-linear nature of the injury risk curve makes it difficult to interpret differences in injury risk. While a 5 % increase in the injury risk when the risk is at 5 % may be due to a significant increase in the model response, a 5 % increase in the injury risk when the risk is at 40 % may be due to a less significant increase. Especially considering that the injury risk curves are not validated for the THUMS.

Despite these limitations on the interpretation, several implications can be formed from these results.

First, the risk of injury suggests that the THUMS is prone to underestimate the risk of injury in the neck and head. In both the head risk measures (HIC15 and HIC36), the risk

of injury for the THUMS head was much lower than for HIII. For example, for the HIC15, the risk of injury for MAIS 3+ injuries was 13%, less than 1/3 that of the HIII MAIS 3+ risk. In fact, this is only 3% over the NHTSA limit. The THUMS underestimation of the injury risk can come from a few different areas. The THUMS neck is structurally different than the HIII. Not only is the THUMS head more physically realistic and complex, it also utilizes external springs and dampers to stiffen up the response of the neck to prevent a loose neck. As a result, it is likely that the head acceleration trace, which controls all the head injury measures, would be different in response when Chapter 4. Injury Risk Functions for THUMS 138 compared to the HIII head acceleration trace. Since the THUMS head injury risk is lower, this imply that the THUMS is better at damping the forces transmitted to the neck, or the THUMS neck itself is better at damping compared to the HIII. This could be remedied by adjusting the stiffness on the neck springs and dampers to correlated well with the HIII head, but that would only work for this specific scenario. The structural accuracy of the THUMS should be preferred to that of HIII, and a better validation study would be to compare the results of THUMS in the simulated environment with a PMHS.

Second, the higher HIC36 value for AIS 4+ risk as compared to HIC15 suggests that the duration of high head accelerations are longer than 15 ms. This would imply that the head does not impact any rigid objects and that the HIC values are a result of head motion only. This seems reasonable as no airbag or dashboard contact was seen in the results. Furthermore, this implies that the 36 ms limit for HIC36 is sensitive to more types of acceleration than the HIC15. The HIC15 therefore oﬀers a more restrictive ﬁlter on the types of acceleration that leads to head injury. Since the HIC was originally designed to assess the risk of skull fracture, then it would seem reasonable that high head acceleration with no head impact such as in this scenario would not cause skull

fracture. And since HIC36 predicts a higher risk of injury, it would seem to be less

accurate than the HIC15 for predicting skull fractures.

Third, Nij values are much closer between the two models compared to the HIC value indicates that the HIII and THUMS have closer neck response than the head. This would seem reasonable as the neck can act as the fulcrum for head motion and that any deviation between the neck motion of the two models would be magnified when it reaches the head. Since the seatbelt will restrain the occupant’s chest, the neck will likely move further than the chest as it is unrestrained. Of interest is also that Nij is calculated from the forces and moments of the neck while the HIC is calculated purely from head acceleration history, which neglects the effective mass of the head. This could suggest that the forces and moments generated in the THUMS neck and head are more comparable to those in the HIII than the kinematics, as the mass difference could counter balance the acceleration difference. There are several factors at work here, the applied boundary condition in this case is the acceleration, a kinematic variable. The Chapter 4. Injury Risk Functions for THUMS 139 acceleration is then transmitted through the seat to the occupant and should create similar kinematics with both the HIII and the THUMS. However, the restraint system in this case restrains the occupant based on a force versus elongation curve, hence the kinematic pulse is overcome by the kinetic-based restraints. Therefore the differences in Nij and HIC is likely due to the different treatment of the restraints on the two different models. A model with higher mass in the upper parts would trigger the load limit sooner than a model with lower mass. Both the THUMS and the HIII were designed to be of similar mass, but the distribution could be different. Another factor that affects the different responses in the Nij and HIC is the rigidity of the models. The HIII is a mostly rigid model with a few elastic components, while the THUMS is a soft body model that has rigid components where is is realistic to have them.

Fourth, the chest injury risk values shows that the THUMS is more likely to overestimate the risk of chest injury from deﬂection while underestimating it from acceleration. The underestimation of risk with chest acceleration seems to follow the trend of HIC and Nij, although the diﬀerences in risk between the THUMS and HIII is less for the chest compared to the HIC. The chest acceleration can be seen to have a higher risk than the

HIC15 for the same AIS level. The interesting thing about the deflection measurement is that it is the first instance of the of an injury measure where the HIII has a lower risk than the THUMS. This deserves more detailed analysis to pinpoint the cause of this difference. Since the injury risk of the chest deflection is correlated with the amount of deflection the ATD experiences, it is necessary that the THUMS experiences more deflection than the HIII. In addition, since both ATDs share the same applied pulse and the same restraint properties, it is necessary that both experience the same change in velocity. It is also known that the maximum acceleration are different between the THUMS and the HIII. If the risk of injury is judged solely by the maximum acceleration experienced by the ATD, then the THUMS will perform better than the HIII for these injury risks and will underestimate the risk of injury caused by safety systems. However, if the risk of injury is judged by the deflection, it can be seen that the THUMS overestimate the risk of injury as compared to the HIII. This indicates that these injury risk measures are highly related to the type of measure. The overestimation of the chest injury risk from deflection likely comes from the softer body of the THUMS that deforms more Chapter 4. Injury Risk Functions for THUMS 140 for the same given impulse. If the THUMS is a more accurate representation of the human body, it raises the point that perhaps the chest deflection injury measure is in the current HIII ATD is underestimating the risk of injury. With incompatible nature of the acceleration based injury measure vs the deflection based injury measure as applied to THUMS, the CTI offers a compromise that combines both the injury measures into a single measure. The CTI value for the THUMS and HIII comes very close to each other. In this case the HIII shows a slightly higher risk of injury than the THUMS. The CTI could be a viable injury risk measure for both the THUMS and the HIII for the chest.

4.5 Conclusion

The study shows that the current injury risk functions do not output similar results between diﬀerent ATDs. Furthermore, THUMS tends to underestimate the risk of injury derived from acceleration based injury measures while it overestimates the risk of injury derived from deﬂection based injury measures as compared to the HIII. Chapter 5

Comparison of diﬀerent load limits using THUMS

5.1 Introduction

The standard three-point belt restraint has been shown to save many lives in MVCs . An eﬀective injury prevention safety system, the seat belt can still be improved. While the general operation of the seatbelt has not changed much, various enhancements have been shown to improve the performance of the seatbelt. One controversial enhancement to the seatbelt is the load limiter. While the load limiter idea itself has been shown to improve outcome of crashes [5–7, 76–79], the optimal value of the load limit is not universal.

The load limiter is usually implemented as an extra fold in the seatbelt that is stitched into the belt. The stitch is designed to break after reaching a certain load and thus allowing the extra belt material to unfold, reducing the force in the belt. Another implementation of the load limiter is a torsion bar in the belt spooling apparatus (retractor) that will start to twist when the force exceeds the limit. While limiting the load of the seatbelt can oﬀer a decrease in the amount of force exerted on the occupant chest by the seatbelt, it can also increase the risk of head injury due to extra excursion of the

141 Chapter 5. Comparison of different load limits using THUMS 142 upper body (when there is no airbag) and increase the force of the airbag on the occupant . There is an optimal force limit that still decreases the force of the belt without incurring too much increase in other chances of injury. The optimal force limit can vary as a function of the occupants chest material property, the severity of the crash pulse, and the geometric configuration of the restraint system. Occupants with more fragile rib cages would benefit more from a lower load limit than occupants with tougher rib cages. In addition, a lower load limit in the shoulder webbing can reduce the risk of submarining as it allows more upper thorax excursion.

Previous studies on the load limiter has shown that the three-point seatbelt and airbag should be balanced in a restraint system for optimal protection of the occupant. The airbag serves to protect mostly in frontal collisions while the seatbelt protects in most directions, with a notable exception in the far side crashes, in which it is fairly easy for the occupant to be thrown out of the seatbelt due to the lack of restraint on the far side. The downside to the three-point belt is that it can cause significant localized chest deflection along the belt path while in high severity crashes . On the other hand, airbags distributes the force over a uniform area and do not cause localized deflection. To increase the effect of the airbag which causes more desirable chest deflections and decrease the localized deflection of the three-point belt, load limits can be imposed that will allow greater excursion of the occupant into the airbag. This will lessen the effects of localized deflection at the cost of greater force from the airbag. In addition, load limiters may not be able to provide as much benefits in non-frontal collisions as side airbags are less common and load limiters will increase the excursion in these directions. In rollovers, seatbelts are still the best form of occupant protection against ejections and gross movements. A looser coupling with the occupant due to load limiters may allow the occupant to spring free of the belt system.

Loads in the shoulder belt have been typically shown to be in the 7 to 8 kN range without load limiters on a 50th percentile male. Load limiters have been tested at various ranges that centers around 4 to 5 kN. While a drastic decrease in the force is produced with a 4 kN load limiter, it typically results in about 0.5 m increase in the head excursion according to Kallieris for a 95th percentile male. The extra excursion for a 5th percentile female is negligible. The eﬀect of the load limiter on the 5th percentile Chapter 5. Comparison of diﬀerent load limits using THUMS 143 female is less pronounced due to the fact that the 5th percentile female is likely to experience less shoulder belt force to begin with. Thus proportionally, the load limit is at a much higher percent of the maximum belt force than for someone who is heavier. To prevent occupant excursion into the steering wheel or dashboard for heavier occupants, load limiters should either consider the worst case scenario, the 95th percentile male, or somehow accounts for the inertia of heaver occupants, such as a weight sensor.

Cadaver tests, field studies, ATDs, and analytical simulations all have been used to analyze the effectiveness of load limiters in different configurations. Crandall et al and Kent et al performed cadaver testing with load limiters and analyzed the effect on the cadaver chest. Crandall experimented with a pretensioned three-point belt assembly that contained a 2 kN load limit. Using a 35 mph (56 kph) crash pulse, Crandall compared both cadaver and ATDs with three restraints systems. Similarly, Kent et al used cadaver sled tests at 30 mph (48 kph) to analyze various restraint combinations with force-limiting belts. Field experimentation with load limiters were done as early as 1970, in which vehicles sold in France included a load limiter in the shoulder webbing that ranged in limit from 2.1 kN to 8 kN .

5.2 Methods

A FE study was performed using THUMS . THUMS is a realistic human body model that attempts to accurate replicate the geometry of a human body in terms of the skeletal structure and the outer tissue. The model used in this study is version 1.62c and represents the 50th percentile male. The model itself is lacking in internal organs but is instead filled with large parts that are meant to behave in a similar manner to the internal organs. The THUMS was placed in a standard three-point restraint seat with retractor, pretensioner, and D-ring. The vehicle interior is simulated with a floor pan and knee bolster. The simulated position is the passenger side. No airbag was modelled in this study as it drastically increases the complexity of the problem. A four-point restraint was also modelled. The resulting FE models can be seen in Figure 5.1. Chapter 5. Comparison of different load limits using THUMS 144

Figure 5.1: Model setup and belt conﬁgurations. Left: three-point restraint system. Right: four-point restraint system

The FE package used in this study is LS-DYNA (Livermore, CA). LS-DYNA is a general purpose nonlinear FE program that is widely used in MVC simulations. LS-DYNA can simulate a variety of diﬀerent scenarios due to its many built-in functionality. The explicit simulation method is well suited to simulating short duration events such MVCs as it has been extensively utilized for this purpose. There are built-in LS-DYNA functionality for implementing seatbelt, retractor, pretensioner, and airbags. In addition, there are multitudes of options available for most of these items.

The LS-DYNA seatbelt element consists of one general seatbelt type and four specialized types that are used in this study. The general seatbelt type element is a 2D beam-like element that exerts force only in tension and its material properties are deﬁned by loading and unloading curves based on force vs. strain. For this study, the belt material properties are shown in Figure 5.2. As can be seen from the ﬁgure, the belt is given simple linear elastic behavior that elongates about 0.1% at 9 kN load.

Pretensioners, sensors, and slip-rings are three types of specialized seatbelt element that are not the focus of this study but are implemented also. Pretensioners controls the seatbelt elements to remove initial slack, sensors acts as triggering mechanisms that activates the other seatbelt elements, and slip-rings help redirect the seatbelt with an Chapter 5. Comparison of diﬀerent load limits using THUMS 145

Belt properties 15 Loading Unloading

10 Force (N)

0 0 0.2 0.4 0.6 0.8 1 1.2 1.4 1.6 −3 Engineering Strain x 10

Figure 5.2: Seatbelt properties: loading and unloading force vs. engineering strain

option of adding friction. These elements often work in conjunction with the retractor to control the behavior of the seatbelt.

Load limiters can be implemented in either the seatbelt element, or in this case, the retractor. A seatbelt based load limiter would imitate the extra loop in the seatbelt that tears when the load limit is exceeded. The retractor based load limiter would imitate the torsion bar in the retractor that twists when the load limit is exceeded. In the LS-DYNA implementation, the retractor can exist in one of two state. The first state is an unlocked state in which in the retractor can freely pay out belt material and attempts to preserve an initial tension value. In this state, the retractor’s initial tension value is specified by a single value and it may not always be able to preserve this initial force. When a seatbelt sensor element fires and acts on the retractor, there is a user defined time delay and the retractor element goes into the locked state. In the locked state, the retractor element pays out belt according to a loading and unloading curve. The retractor will follow the loading curve in tension and follows the unloading curve if there is no tension in the belt. There are two ways for the retractor to pays out belt material, the length of the last element attached to the retractor is lengthened when Chapter 5. Comparison of different load limits using THUMS 146 there are no more elements inside the retractor. However, if there are elements defined that are initially inside the retractor, the elements can be pulled-out with a predefined threshold. No predefined elements are defined for this study as the retractor is far enough from the D-ring to prevent any unrealistic behavior from an oversized seatbelt element.

The retractor loading curves for the three-point restraint is shown in Figure 5.3. The retractor loading curve shows three separate load limit levels. The ﬁrst load limit level at 3 kN is a relatively low load limit and belt payout is expected to be more than the other load limit levels. The load limit level at 6 kN represents a medium level for the force limit as it is approximately 70% of the expected maximum belt tension. The highest level, with no limit, is the control case that simulates no load limit on the belt. There is expected to be a diﬀerence of 3kN between each of the load limit levels as the highest shoulder belt force is expected to be around 8-9 kN with no load limit as seen from previous studies . The force vs pull-out curve shows that there is an overshoot in the force above the load limiter level and then a sudden drop as the load limiter activates, and a sharp increase in the force as the load limiter engages. The drop for the higher load limit is sharper than the drop for the lower load limit. Note that the loading curves for the load limiter comes from restraint manufacturer’s direct input and are not fabricated or contrived examples.

The retractor loading curves for the four-point restraint as shown in Figure 5.4 is derived from the loading curves for the three-point restraint and only scaled by a factor of 0.5 for the force. The curves are identical in other aspects. The four-point retractor loading curves are scaled by half due to the fact that there are two shoulder belts in the four-point restraint instead of the one shoulder belt in the three-point restraint. The retractor slightly increases the force for extremely large pull-outs for both restraint types.

The triggering sensor for the retractor is a time based sensor and is set according to the peaking of the crash pulse. The crash pulse applied for the simulation is shown in Figure 5.5 and is approximately a 40 mph (64.37 kph) DV. The sensor is triggered at 9 ms, approximately the time the ﬁrst peak in acceleration occurs. After the retractor Chapter 5. Comparison of diﬀerent load limits using THUMS 147

Three−point retractor properties 10 3 kN Load Limit 9 6 kN Load Limit No Load Limit 8

Force (kN) 4

0 0 0.2 0.4 0.6 0.8 1 Pull−out (m)

Figure 5.3: Three-point retractor properties: loading and unloading force vs pull-out for the retractor at load limits of 3 kN, 6 kN, and no limit

locks, the pretensioner is engaged and pulls in enough belt material to create a 200 N tension in the belt for both three-point and four-point restraints. Then the pretensioner is disengaged and the retractor takes over.

Table 5.1: Restraint activation sequence timeline for three-point restraint. Note that load limiter engagement occurs only for speciﬁc cases.

Event Action 9 ms Retractor sensor ﬁres - retractor locks 10 ms Pretensioner sensor ﬁres - pretensioner locks 200 N tension reached Pretensioner disengages - retractor active ≈ 3 kN tension reached Load limiter 1 engages (for load limit 1 cases) ≈ 7.5 kN tension reached Load limiter 2 engages (for load limit 2 cases)

The case matrix for this study is shown in Table 5.2. Both three-point and four-point restraints were ran with each load limit level, with the three-point vs. four-point the focus of another study. The current load limit levels are analyzed with respect to some

standard dummy injury measures: HIC36, HIC15, Nij, chest deflection and 3 ms clip, rib deflections, abdominal accelerations, and pelvic accelerations. Chapter 5. Comparison of different load limits using THUMS 148

Four−point retractor properties 5 1.5 kN Load Limit 4.5 3 kN Load Limit No Load Limit 4

3.5

2.5

Force (kN) 2

1.5

0.5

0 0 0.2 0.4 0.6 0.8 1 Pull−out (m)

Figure 5.4: Retractor properties: loading and unloading force vs pull-out for the retractor at load limits of 1.5 kN, 3 kN, and no limit

Applied Crash Pulse

300 ) 2 s m

( 200

100 acceleration 0

0 0.05 0.1 0.15 0.2 0.25 time (s)

Figure 5.5: Crash pulse applied: acceleration vs. time

Table 5.2: Load limiter study case matrix

Load limit level Restraints 3 kN Three-point and Four-point 6 kN Three-point and Four-point No limit Three-point and Four-point Chapter 5. Comparison of diﬀerent load limits using THUMS 149

5.3 Results

5.3.1 Simulation snapshots

The simulation snapshots for the three-point belt is shown in Figure 5.6, Figure 5.7, and Figure 5.8. Qualitatively, the simulation snapshots show that the load limiter 1 case has more head displacement than the load limiter 2 case, which in turn has more displacement than the no load limiter case in the ﬁnal frame. As a result, the no load limiter case shows THUMS rebounding at an earlier time than the HIII.

Figure 5.6: Three-point belt with load limiter 1. Chapter 5. Comparison of diﬀerent load limits using THUMS 150

Figure 5.7: Three-point belt with load limiter 2.

Figure 5.8: Three-point belt with no load limiter. Chapter 5. Comparison of diﬀerent load limits using THUMS 151

The simulation snapshots for the four-point belt is shown in Figure 5.11, Figure 5.9, and Figure 5.10. Similar to the three-point belt, the load limiter 1 case has higher head displacement than the load limiter 2 case, which in turn has higher head displacement than the no load limiter case. The four-point belt scenarios also have lower head displacement in general than the three-point belt scenarios.

Figure 5.9: Four-point belt with load limiter 1. Chapter 5. Comparison of diﬀerent load limits using THUMS 152

Figure 5.10: Four-point belt with load limiter 2.

Figure 5.11: Four-point belt with no load limiter. Chapter 5. Comparison of diﬀerent load limits using THUMS 153

5.3.2 Head injury measures

Head injury measures are shown below. It can be seen that for most injury measures, the injury measures are highest for load limiter 2 case in the three-point belt restraint and the no load limiter case in the four-point belt restraint. The 300 N load limiter cases shows the lowest injury measures.

The maximum head accelerations are shown in Figure 5.12. It can be seen that for the three-point case, the load limiter 2 case shows an increase in head acceleration from the load limiter 1 case. The four-point scenarios shows the lowest head acceleration with load limiter 2.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Max Acceleration (Gs) for all cases 80

30 Max Acceleration (Gs)

0 3Pt Retractor 4Pt Retractor Pretensioner Type

Figure 5.12: THUMS maximum head acceleration (Gs) for diﬀerent load limits. Chapter 5. Comparison of diﬀerent load limits using THUMS 154

The maximum HICs36 are shown in Figure 5.13. Both the three-point and the four- point cases show that the load limiter 1 creates the lowest HIC36.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

HIC36 for all cases 1000

900

800

700

600

500 HIC36 400

300

200

100

0 3Pt Retractor 4Pt Retractor Pretensioner Type

Figure 5.13: THUMS maximum HIC36 for diﬀerent load limits.

The maximum HICs15 are shown in Figure 5.14. Both the three-point and the four-

point cases show that the load limiter 1 creates the lowest HIC15 also.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

HIC15 for all cases 600

500

400

300 HIC15

200

100

0 3Pt Retractor 4Pt Retractor Pretensioner Type

Figure 5.14: THUMS maximum HIC15 for diﬀerent load limits. Chapter 5. Comparison of diﬀerent load limits using THUMS 155

5.3.3 Neck injury measures

The neck injury measures show similar trends to that of the head injury measures. The load limiter 1 cases have the lowest injury measure. In addition, the three-point restraints have slightly higher load limiter 1 values than the four-point restraints.

The THUMS maximum neck tensions are shown in Figure 5.15. Both the three-point and the four-point restraint shows a lower neck normal force for the load limiter 1. The three-point load limiter 1 has a higher value than the four-point load limiter 1, however.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Max Normal Force (N) for all cases 1200

1000

800

600

Max Normal Force (N) 400

200

0 3Pt Retractor 4Pt Retractor Pretensioner Type

Figure 5.15: THUMS neck tension for diﬀerent load limits. Chapter 5. Comparison of diﬀerent load limits using THUMS 156

The THUMS maximum Nijs are shown in Figure 5.16. Both the three-point and the four-point restraints show a lower Nij value for the load limiter 1. Similar to neck tension, the load limiter 1 shows a higher value for the three-point restraint.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Max Nij for all cases 0.18

0.16

0.14

0.12

0.1

Max Nij 0.08

0.06

0.04

0.02

0 3Pt Retractor 4Pt Retractor Pretensioner Type

Figure 5.16: THUMS Nij for diﬀerent load limits. Chapter 5. Comparison of diﬀerent load limits using THUMS 157

5.3.4 Chest injury measures

The chest acceleration (3 ms clip) is shown in Figure 5.17. The load limiter 1 again shows a lower force than either the load limiter 2 or the no load limiter case. The three- point restraint shows a slightly higher chest acceleration than the four-point restraint.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

3ms Clip (Gs) for all cases 45

20 3ms Clip (Gs) 15

0 3Pt Retractor 4Pt Retractor Pretensioner Type

Figure 5.17: THUMS chest 3 ms clip for diﬀerent load limits. Chapter 5. Comparison of diﬀerent load limits using THUMS 158

The THUMS chest compression (sternal deﬂection) values are shown in Figure 5.18 for both the three-point and four-point restraints. The load limiter 1 shows a lower absolute compression than the other load limiter cases. In addition, the four-point restraint shows a much lower value for the load limiter 1 than the three-point restraint.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Min Sternal Deflection (%) for all cases 0

−5

−10

−15

−20

Min Sternal Deflection (%) −25

−30

−35 3Pt Retractor 4Pt Retractor Pretensioner Type

Figure 5.18: THUMS chest deflection for different load limits. Chapter 5. Comparison of different load limits using THUMS 159

The THUMS CTI values are show in Figure 5.19. The load limiter 1 again shows the lowest value compared to the other load limiter cases.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Max CTI for all cases 1.4

1.2

0.8

Max CTI 0.6

0.4

0.2

0 3Pt Retractor 4Pt Retractor Pretensioner Type

Figure 5.19: THUMS CTI for diﬀerent load limits.

The THUMS maximum rib stresses (across the front of the rib cage) are shown in Figure 5.20. There does not seem to be a consistent pattern across the diﬀerence cases. The three-point restraint has a higher load limiter 1 value than the other load limiters, while the four-point restraint has a slightly higher no load limiter value. Chapter 5. Comparison of diﬀerent load limits using THUMS 160

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Max Rib Stress (MPa) for all cases 180

160

140

120

100

Max Rib Stress (MPa) 60

0 3Pt Retractor 4Pt Retractor Pretensioner Type

Figure 5.20: THUMS rib stress for diﬀerent load limits. Chapter 5. Comparison of diﬀerent load limits using THUMS 161

The THUMS lumbar normal forces are shown in Figure 5.21. Most of the cases show similar forces at approximately 2800 N, except for the four point load limiter 1, which shows a higher force of approximately 3400 N. It should be noted that the forces in the lumbar discs are tensile forces.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Lumbar Force (N) for all cases 3500

3000

2500

2000

1500 Lumbar Force (N)

1000

500

0 3Pt Retractor 4Pt Retractor Pretensioner Type

Figure 5.21: THUMS lumbar force for diﬀerent load limits. Chapter 5. Comparison of diﬀerent load limits using THUMS 162

5.3.5 Pelvis injury measures

The pelvic injury measure results are shown below. Unlike the chest and head injury measures, there are some injury measures here that do not show a consistent trend in the load limiter cases between three-point and four-point restraints. In some cases, the load limiter 2 case show a much higher injury measure value than the other load limiter cases.

The maximum abdominal acceleration is shown in Figure 5.22. The four-point load limiter 1 case shows a lower abdominal acceleration than the other load limiter cases. This trend is not reﬂected in the three-point restraints, where the injury measure values are similar.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Max Abdominal Acceleration (Gs) for all cases 45

Max Abdominal Acceleration (Gs) 10

0 3Pt Retractor 4Pt Retractor Pretensioner Type

Figure 5.22: THUMS maximum abdominal acceleration for diﬀerent load limits. Chapter 5. Comparison of diﬀerent load limits using THUMS 163

The maximum pelvic acceleration is shown in Figure 5.23. Unlike the abdominal acceleration, the three-point restraint shows a much lower pelvic acceleration for the load limiter 1 case as compared to the other load limiter cases. The four-point restraint shows a lower pelvic acceleration for the load limiter 2 value.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Max Apelvic (Gs) for all cases 100

40 Max Apelvic (Gs)

0 3Pt Retractor 4Pt Retractor Pretensioner Type

Figure 5.23: THUMS maximum pelvic acceleration for diﬀerent load limits. Chapter 5. Comparison of diﬀerent load limits using THUMS 164

The maximum VC is shown in Figure 5.24. This criterion represents a direction of motion that is the same as the current state of the abdomen, such as a compressed abdomen that is being compressed. It can be seen that the load limiter 1 shows the lowest injury measure values for both the three-point and the four-point restraints.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Max Viscous from Avg deflections (m/s) for all cases 0.7

0.6

0.5

0.4

0.3

0.2 Max Viscous from Avg deflections (m/s) 0.1

0 3Pt Retractor 4Pt Retractor Pretensioner Type

Figure 5.24: THUMS maximum VC for diﬀerent load limits. Chapter 5. Comparison of diﬀerent load limits using THUMS 165

The minimum VC is shown in Figure 5.25. This criterion represents a direction of motion that is the opposite of the current state of the abdomen, such as a compressed abdomen that is being uncompressed. Unlike the maximum VC, the load limiter 1 seems to show a higher absolute value than the other load limiters for both three-point and four-point restraints.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Min Viscous from Avg deflections (m/s) for all cases 0

−0.2

−0.4

−0.6

−0.8

−1 Min Viscous from Avg deflections (m/s) −1.2

−1.4 3Pt Retractor 4Pt Retractor Pretensioner Type

Figure 5.25: THUMS minimum VC for diﬀerent load limits. Chapter 5. Comparison of diﬀerent load limits using THUMS 166

The THUMS maximum lap belt force as exerted on the abdomen can be seen in Fig- ure 5.26. The three-point restraint shows the load limiter 1 as having the lowest lap belt force while the four-point restraint shows the load limiter 1 as having the highest lap belt force. The four-point restraint load limiter 1 lap belt force is approximately 6000 N higher than the other cases.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

4 x 10 Max Lap Belt Force (N) for all cases 3

2.5

1.5

1 Max Lap Belt Force (N)

0.5

0 3Pt Retractor 4Pt Retractor Pretensioner Type

Figure 5.26: THUMS maximum lap belt force for diﬀerent load limits. Chapter 5. Comparison of diﬀerent load limits using THUMS 167

The maximum force of the lap belt on THUMS left iliac crest is shown in Figure 5.27. The three-point restraint load limiter 1 shows an approximately 25% reduction in forces compared to the other three-point cases. The four-point restraint load limiter 1 and no load limiter cases shows approximately an 80 % reduction in forces compared to the load limiter 1 case.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Max Left Iliac Crest Force (N) for all cases 1200

1000

800

600

400 Max Left Iliac Crest Force (N)

200

0 3Pt Retractor 4Pt Retractor Pretensioner Type

Figure 5.27: THUMS left iliac crest force for diﬀerent load limits. Chapter 5. Comparison of diﬀerent load limits using THUMS 168

5.3.6 Risks

The injury risks as calculated from the maximum of the injury measures are shown below. The head injury risks include the HIC36 and HIC15, the neck injury risk used is the Nij, the chest injury risks are the chest 3 ms clip, the chest deﬂection, and the CTI.

The HIC36 risk of injury is calculated using Viano’s formulation, which is based on a logistic distribution. It can be seen that the four-point load limiter 1 has the lowest risk

of injury at approximately 7% risk of AIS 4+ head injury as predicted by the HIC36. The risk of injury for the three-point load limiter 1 was also low. The load limiter 1 injury risks were the only cases below 10% risk of injury.

HIC36 Probability of AIS 4+ Injury

AIS 4+

Retr 3Pt L. Limit 1

Retr 3Pt L. Limit 2

Retr 3Pt No L. Limit

Retr 4Pt L. Limit 1

Retr 4Pt L. Limit 2

Retr 4Pt No L. Limit

0% 10% 20% 30% 40% 50% 60% 70% 80% 90% 100%

Figure 5.28: THUMS HIC36 risks of injury. Chapter 5. Comparison of diﬀerent load limits using THUMS 169

The HIC15 risk of injury is calculated using the expanded Mertz/Prasad curves. Unlike the HIC36, the HIC15 shows a lower risk of AIS 4+ injury for all the cases. All the cases show an AIS 4+ risk of injury at below 5%. The load limiter 1 cases show a lower risk of injury compared to the other cases.

HIC15 Probability of Injury

Fatal MAIS5+ MAIS4+ MAIS3+ MAIS2+ MAIS1+

Retr 3Pt L. Limit 1

Retr 3Pt L. Limit 2

Retr 3Pt No L. Limit

Retr 4Pt L. Limit 1

Retr 4Pt L. Limit 2

Retr 4Pt No L. Limit

0% 10% 20% 30% 40% 50% 60% 70% 80% 90% 100%

Figure 5.29: THUMS HIC15 risks of injury. Chapter 5. Comparison of diﬀerent load limits using THUMS 170

The Nij AIS 4+ risks of injury is shown in Figure 5.30. The risks of injury for all cases are close and below 10%.

Nij Probability of AIS 4+ Injury

AIS 4+

Retr 3Pt L. Limit 1

Retr 3Pt L. Limit 2

Retr 3Pt No L. Limit

Retr 4Pt L. Limit 1

Retr 4Pt L. Limit 2

Retr 4Pt No L. Limit

0% 10% 20% 30% 40% 50% 60% 70% 80% 90% 100%

Figure 5.30: THUMS Nij risks of injury. Chapter 5. Comparison of diﬀerent load limits using THUMS 171

The chest 3ms clip is shown in Figure 5.31. The four-point load limiter 1 case shows the lowest risks of injury at which the AIS 4+ risks is approximately 12%. The next lowest risks case is the three-point load limiter 1. The four-point restraint shows higher risks for the load limiter 2 while the three-point restraint shows a higher risk for the no load limiter case.

Chest 3−ms Clip Probability of Injury

AIS 5+ AIS 4+ AIS 3+ AIS 2+

Retr 3Pt L. Limit 1

Retr 3Pt L. Limit 2

Retr 3Pt No L. Limit

Retr 4Pt L. Limit 1

Retr 4Pt L. Limit 2

Retr 4Pt No L. Limit

0% 10% 20% 30% 40% 50% 60% 70% 80% 90% 100%

Figure 5.31: THUMS chest 3ms clip risks of injury. Chapter 5. Comparison of diﬀerent load limits using THUMS 172

The sternal deﬂection risks are shown in Figure 5.32. The risks of injury for the four- point load limiter 1 case is lower than the other cases at less than half of any of the other cases. The risks of injury for the three-point load limiter 1 case is the second lowest but does not show as drastic a reduction.

Sternal Deflection Probability of Injury

AIS 5+ AIS 4+ AIS 3+ AIS 2+

Retr 3Pt L. Limit 1

Retr 3Pt L. Limit 2

Retr 3Pt No L. Limit

Retr 4Pt L. Limit 1

Retr 4Pt L. Limit 2

Retr 4Pt No L. Limit

0% 10% 20% 30% 40% 50% 60% 70% 80% 90% 100%

Figure 5.32: THUMS chest deﬂection risks of injury. Chapter 5. Comparison of diﬀerent load limits using THUMS 173

The CTI risks of injury is show in Figure 5.33. The risks of the three-point restraint cases as a whole are less than the risks of the four-point restraint cases, with the three- point load limiter 1 being the case with the lowest risk.

Combined Thoracic Index (CTI) Probability of Injury

AIS 5+ AIS 4+ AIS 3+ AIS 2+

Retr 3Pt L. Limit 1

Retr 3Pt L. Limit 2

Retr 3Pt No L. Limit

Retr 4Pt L. Limit 1

Retr 4Pt L. Limit 2

Retr 4Pt No L. Limit

0% 10% 20% 30% 40% 50% 60% 70% 80% 90% 100%

Figure 5.33: THUMS CTI risks of injury. Chapter 5. Comparison of diﬀerent load limits using THUMS 174

5.4 Discussion

As can be seen from the results, the load limiter 1 shows the lowest values in most of the cases. This conforms with previous studies that shows a drop in the injury measures as the load limit on the seat belt is reduced. Especially for older occupants with more fragile ribs, the beneﬁt of a decreased load limit could possibly save more lives. The caveat of lower load limits is that there is more occupant upper body displacement as can be seen in the simulation snapshots. If the head would have come into contact with a rigid object such as the dashboard or steering wheel, the risks of head injury would likely increase dramatically. The overall excursion of THUMS is quite high as shown in the simulation snapshots. Although the snapshots were not timed with key events in the simulation, but merely taken at ﬁxed time points, it can still be seen that the maximum head excursion penetrates THUMS own thighs for the lower load limits as shown in pane 5.

The injury metrics between the three-point belt and the four-point belt shows similar values for most of the injury measures. Besides the fact that the load limiter 1 has the lowest values in most cases, the load limiter 2 and the no load limiter case does not show a trend that the lower load limit will result in a lower injury measure. This could indicate the existence of a load limiter saturation point at which load limits higher than the saturation point will have no eﬀect compared to a belt with no load limit.

For most cases, the increase in the injury measure from the load limiter 1 cases to the higher load limit levels are not drastic. However, for HIC, chest sternal compression, and abdominal viscous criterion, the values are much lower for the load limiter 1. The four-point restraint has particularly low values of chest deflection and abdominal viscous criterion. The THUMS lap belt force indicates that the four-point load limiter 1 has the highest forces, which could indicate that the lap belt exerted more force on the THUMS than the shoulder belt in the load limiter 1 case than the other cases. This could explain the lower chest deflection. The cause of the low abdominal VC is less clear, but the THUMS shoulder belt passes close to the calculation point of the VC at the center of the abdominal organs, so that minor variation in the belt behavior could cause larger changes in the VC. A more robust way of calculate the less stable injury Chapter 5. Comparison of different load limits using THUMS 175 measures could be to average across multiple points. As can be seen in Appendix A, the VC is noisy even after filtering.

It should be noted that the left iliac force for the three-point belt is much higher than the four-point belt. This again could be due to belt positioning. Another factor that could contribute to the diﬀerence is the geometry of the belt layout. The buckle in this case acts as a pulley that redirects the force of the shoulder belt into the lap belt. However, for the four-point belt, this pulley system is not present, and is replaced by a simple anchor. Thus it should be expected that the force is greater for the three-point belt.

Most of the risks of injury for THUMS are less than 20%. Given the severity of the crash pulse, this seems acceptable. However, it is important to note that the injury risks are not calibrated to an ATD or any standards.

5.5 Conclusion

The overall trend of the study indicates a notable decrease in the injury measure for the load limiter 1. There is less diﬀerentiation between the load limiter 2 and the no load limiter case. The caveat to the load limiter trend is that there was no airbag or dashboard to intercept the head. Chapter 6

Comparison of diﬀerent pretensioner placement using THUMS

6.1 Introduction

Car safety systems are often difficult to optimize due to lack of data on human injury response and the variability among different occupants. Safety systems are often designed to pass the current regulation tests and with the idea of satisfying the existing injury criteria as measured in dummies as the gold standard. Often the subtle effects on real humans are lost in the current crash test dummies, also called ATDs. New sub- stitutes in the form of FE total human models can supplement some of the downfalls of the ATD. The THUMS is a geometrically detailed FE model of the human body that has been validated in multiple studies. It can be used in a simulated frontal crash test to supplement or improve results from ATDs. FE models can be especially useful for optimization purposes as it is much cheaper than using ATDs in real crash tests.

One safety system that can be optimized is the pretensioner placement for the seatbelt. In MVCs, there is often a signiﬁcant amount of slack in the belt that shortens the actual contact time between the occupant and the belt. This slack can be between 40 mm to 176 Chapter 6. Comparison of diﬀerent pretensioner placement using THUMS 177

90 mm or even higher in the winter time . The role of the pretensioner is to prolong the contact duration between the occupant and the seatbelt. In a MVC, the occupant must be accelerated from an initial velocity to a ﬁnal velocity of around zero. How serious the injuries to the occupant depends on the way the acceleration occurs. Since chest acceleration is one of the criteria to assess injury, then the best way to reduce chest acceleration is to attempt to create a constant force between the seatbelt and the chest throughout the event. The seatbelt should then stay in contact with the chest as long as possible to reduce the belt force. The pretensioner plays an important role here as it removes the initial slack in the seatbelt prior to a crash and tightens the belt around the occupant.

Pretensioner [80] in automobiles generally consists of either a pyrotechnic pretensioner which activates by igniting gases, or a spring based pretensioner which locks or unlocks a stiﬀ spring. Pretensioners generally activate via the airbag control module.

There are a few studies on the effect of pretensioners in the vehicle. [81] found the addition of pretensioners to significantly lower HIC, chest acceleration, and chest deflection in New Car Assessment Program (NCAP) tests. [82] found pretensioners and load limiters to improve injury measures in rear seat passengers. [83] introduced a coupling criterion to assess the effectiveness of the pretensioner. [84] studied the effect seatbelt pretensioners on the abdomen.

There are several implementations of the pretensioner/retractor combination in the com- mercial FE solvers LS-DYNA. There are implementations that differentiate between pretensioner type and pretensioner/retractor interaction. The pretensioner types include the pyrotechnic pretensioner and lock spring pretensioner. The pyrotechnic pretensioner also requires a pull-in vs. time curve to define the pretensioner behavior, whereas the spring base pretensioner behavior is defined by the spring element. Pretensioner/retractor interaction can be configured differently depending on the desired behavior but it can often be difficult to decide on the most realistic type of interaction. Chapter 6. Comparison of different pretensioner placement using THUMS 178

6.2 Methods

A FE study was performed with THUMS using LS-DYNA with varying pretensioner placement on a three-point belt system. The general setup of the simulation is shown in Figure 6.2. A standard three-point belt is used along with knee bolster and ﬂoorboards. No airbags were modeled.

Applied Crash Pulse

300 ) 2 s m

( 200

100 acceleration 0

0 0.05 0.1 0.15 0.2 0.25 time (s)

Figure 6.1: Crash pulse applied: acceleration vs. time

Retractors, sensors, and slip-rings are three types of specialized seatbelt element that are not the focus of this study but are implemented also. Retractors can pull-in belt elements depending on its state, sensors acts as triggering mechanisms that activates the other seatbelt elements, and slip-rings help redirect the seatbelt with an option of adding friction. These elements often work in conjunction with the pretensioner to control the behavior of the seatbelt. In addition, some scenarios also have load limiters.

The pretensioner used in both the retractor pretensioner and the buckle pretensioner is a LS-DYNA type 5 pyrotechnic retractor with force limiter. The type 5 pretensioner accepts a pull-in vs time curve and the belt is drawn into the retractor exactly as specified. A force limit of 200 N restricts the maximum force in the belt such that anytime that force limit is exceeded, the pretensioning is aborted. This is prevent unrealistic strangling of the occupant. Pretensioner pull-in vs time data, as shown in Figure 6.3 was provided by TRW. The pretensioner type 5 was chosen due to the fact Chapter 6. Comparison of different pretensioner placement using THUMS 179 that it was the only type that followed the specified pull-in vs time exactly. The other types tried to maintain realistic forces at the cost not following a pull-in vs time curve. r e n o i s n e t e r B P u

r c o k t l e c

a P r t r e e t R e n s i o n e r

Figure 6.2: Pretensioners can be located in either the retractor, in which it pretensioners through the slip ring, or the buckle. Chapter 6. Comparison of diﬀerent pretensioner placement using THUMS 180

Pretensioner pull-in vs. time

0.1

0.08 ) m

( 0.06

0.04 pull-in

0.02

0 0 0.05 0.1 0.15 0.2 0.25 time (s)

Figure 6.3: Pretensioner pull-in vs. time Chapter 6. Comparison of diﬀerent pretensioner placement using THUMS 181

In addition to the pretensioner, load limiters are also implemented to see the interaction with pretensioners.

Table 6.1: Restraint activation sequence timeline for three-point restraint. Note that load limiter engagement occurs only for speciﬁc cases.

Event Action 9 ms Retractor sensor fires - retractor locks 10 ms Pretensioner sensor fires - pretensioner fires 200 N tension reached Pretensioner disengages - retractor active ≈ 3 kN tension reached Load limiter 1 engages (for load limit 1 cases) ≈ 7.5 kN tension reached Load limiter 2 engages (for load limit 2 cases)

The triggering mechanism for the pretensioner is a time based sensor that fires at the time specified. The 10 ms pretensioner firing time is specified by TRW.

The case matrix for this study is speciﬁed in Table 6.2. Three diﬀerent cases of pretensioners are modeled - buckle pretensioner, retractor pretensioner, and both. When both pretensioner are active, the pull-in vs time curve is scaled by a factor of 0.5 in the y-direction to maintain the total pull-in.

Table 6.2: Load limiter study case matrix

Pretensioner Cases Both no load limit, 3 kN and 6 kN load limiters Buckle no load limit, 3 kN and 6 kN load limiters Retractor no load limit, 3 kN and 6 kN load limiters Chapter 6. Comparison of diﬀerent pretensioner placement using THUMS 182

6.3 Results

A list of simulation snapshots is shown ﬁrst at ﬁxed time points in the simulation. Maximum injury measures values are then presented as bar charts grouped by the pretensioner type. Injury risk calculated from the injury measures are then presented as stacked bar charts.

6.3.1 Simulation snapshots

This section shows the visual snapshots of the simulation at ﬁxed time points. The outputs by diﬀerent pretensioners at the same load limit are similar.

The simulation snapshot for all pretensioner types for the load limiter 1 case is shown in Figure 6.4, Figure 6.5, and Figure 6.6. Visual inspection shows no drastic diﬀerences between the diﬀerent pretensioner types.

Figure 6.4: Three-point belt with both pretensioner and 3 kN load limit. Chapter 6. Comparison of diﬀerent pretensioner placement using THUMS 183

Figure 6.5: Three-point belt with buckle pretensioner and 3 kN load limit.

Figure 6.6: Three-point belt with retractor pretensioner and 3 kN load limit. Chapter 6. Comparison of diﬀerent pretensioner placement using THUMS 184

The simulation snapshots for all pretensioner types for the load limiter 2 case is shown in Figure 6.7, Figure 6.8, Figure 6.9. Visual inspection shows no drastic diﬀerences between the diﬀerent pretensioner types.

Figure 6.7: Three-point belt with both pretensioner and 7.5 kN load limit.

Figure 6.8: Three-point belt with buckle pretensioner and 7.5 kN load limit. Chapter 6. Comparison of diﬀerent pretensioner placement using THUMS 185

Figure 6.9: Three-point belt with retractor pretensioner and 7.5 kN load limit. Chapter 6. Comparison of diﬀerent pretensioner placement using THUMS 186

The simulation snapshots for all pretensioner types for the no load limiter case is shown in Figure 6.7, Figure 6.11, Figure 6.12. Visual inspection shows no drastic diﬀerences between the diﬀerence pretensioner types.

Figure 6.10: Three-point belt with both pretensioner and no load limit.

Figure 6.11: Three-point belt with buckle pretensioner and no load limit. Chapter 6. Comparison of diﬀerent pretensioner placement using THUMS 187

Figure 6.12: Three-point belt with retractor pretensioner and no load limit. Chapter 6. Comparison of diﬀerent pretensioner placement using THUMS 188

6.3.2 Head injury measures

The maximum head acceleration is shown in Figure 6.13. There is not a noticeable diﬀerence across diﬀerent pretensioner types. The three-point retractor pretensioner does show that the no load limiter case has a lower value than the load limiter 2 case whereas the other types of pretensioners have a higher no load limiter value. The three- point buckle pretensioner has the lowest maximum head acceleration by a slight amount.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Max Acceleration (Gs) for all cases 80

30 Max Acceleration (Gs)

0 3Pt Both 3Pt Buckle 3Pt Retractor Pretensioner Type

Figure 6.13: THUMS maximum head acceleration (Gs) for diﬀerent pretensioners. Chapter 6. Comparison of diﬀerent pretensioner placement using THUMS 189

The maximum HIC36 is shown in Figure 6.14. There is not a significant difference between different pretensioner types. The buckle pretensioner shows slightly higher values for the load limiter 1 and load limiter 2 cases.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

HIC36 for all cases 1000

900

800

700

600

500 HIC36 400

300

200

100

0 3Pt Both 3Pt Buckle 3Pt Retractor Pretensioner Type

Figure 6.14: THUMS maximum HIC36 for diﬀerent pretensioners.

The maximum HIC15 is shown in Figure 6.15. Both the buckle and the retractor pretensioner shows a higher load limiter 2 value than the no load limiter value while the both pretensioner case shows a slightly higher no load limiter value than the load limiter 2 value. Chapter 6. Comparison of diﬀerent pretensioner placement using THUMS 190

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

HIC15 for all cases 600

500

400

300 HIC15

200

100

0 3Pt Both 3Pt Buckle 3Pt Retractor Pretensioner Type

Figure 6.15: THUMS maximum HIC15 for diﬀerent pretensioners. Chapter 6. Comparison of diﬀerent pretensioner placement using THUMS 191

6.3.3 Neck injury measures

The neck tension force is shown in Figure 6.16. The values are close to each other for all pretensioner types.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Max Normal Force (N) for all cases 1200

1000

800

600

Max Normal Force (N) 400

200

0 3Pt Both 3Pt Buckle 3Pt Retractor Pretensioner Type

Figure 6.16: THUMS neck tension for diﬀerent pretensioners.

The Nij is shown in Figure 6.17. The both pretensioner case shows an increase in value as the load limit increases while the buckle pretensioner shows a decrease in value as the load limit increases. The diﬀerences are, however, minor. Chapter 6. Comparison of diﬀerent pretensioner placement using THUMS 192

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Max Nij for all cases 0.18

0.16

0.14

0.12

0.1

Max Nij 0.08

0.06

0.04

0.02

0 3Pt Both 3Pt Buckle 3Pt Retractor Pretensioner Type

Figure 6.17: THUMS Nij for diﬀerent pretensioners. Chapter 6. Comparison of diﬀerent pretensioner placement using THUMS 193

6.3.4 Chest injury measures

The chest 3ms clip is shown in Figure 6.18. The both pretensioner case shows the lowest 3ms clip for the load limiter 1 case, but also has the higher 3ms clip for the no load limiter case.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

3ms Clip (Gs) for all cases 50

3ms Clip (Gs) 20

0 3Pt Both 3Pt Buckle 3Pt Retractor Pretensioner Type

Figure 6.18: THUMS chest 3 ms clip for diﬀerent pretensioners.

The sternal deﬂection is shown in Figure 6.19. The retractor pretensioner shows the smallest amount of sternal deﬂection for the load limiter 1 case.

The CTI is shown in Figure 6.20. The retractor pretensioner shows the lowest CTI for the load limiter 1 case.

The maximum rib stresses are shown in Figure 6.21. The buckle pretensioner has the lowest value with the load limiter 2 cases. This is unlike most of the other injury measures where the load limiter 1 case has the lowest values.

The lumbar forces are shown in Figure 6.22. The lowest lumbar forces are seen in the both pretensioner case with no load limit. Chapter 6. Comparison of diﬀerent pretensioner placement using THUMS 194

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Min Sternal Deflection (%) for all cases 0

−5

−10

−15 Min Sternal Deflection (%)

−20

−25 3Pt Both 3Pt Buckle 3Pt Retractor Pretensioner Type

Figure 6.19: THUMS chest deﬂection for diﬀerent pretensioners.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Max CTI for all cases 1.4

1.2

0.8

Max CTI 0.6

0.4

0.2

0 3Pt Both 3Pt Buckle 3Pt Retractor Pretensioner Type

Figure 6.20: THUMS CTI for diﬀerent pretensioners. Chapter 6. Comparison of diﬀerent pretensioner placement using THUMS 195

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Max Rib Stress (MPa) for all cases 200

180

160

140

120

100

Max Rib Stress (MPa) 60

0 3Pt Both 3Pt Buckle 3Pt Retractor Pretensioner Type

Figure 6.21: THUMS rib stress for diﬀerent pretensioners.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Lumbar Force (N) for all cases 3000

2500

2000

1500 Lumbar Force (N) 1000

500

0 3Pt Both 3Pt Buckle 3Pt Retractor Pretensioner Type

Figure 6.22: THUMS lumbar force for diﬀerent pretensioners. Chapter 6. Comparison of diﬀerent pretensioner placement using THUMS 196

6.3.5 Pelvis injury measures

The maximum abdominal acceleration for different pretensioners is shown in Figure 6.23. There does not seem to be a significant difference for different pretensioners.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Max Abdominal Acceleration (Gs) for all cases 45

Max Abdominal Acceleration (Gs) 10

0 3Pt Both 3Pt Buckle 3Pt Retractor Pretensioner Type

Figure 6.23: THUMS maximum abdominal acceleration for diﬀerent pretensioners.

The maximum pelvic acceleration for diﬀerent pretensioners is shown in Figure 6.24. The both pretensioner with no load limit case and the retractor pretensioner with load limit 1 shows the lowest maximum pelvic acceleration.

The maximum abdominal VC is shown in Figure 6.25. The both pretensioner case has low values for the load limiter 1 and load limiter 2 case, but a high value for the no load limit case.

The minimum abdominal VC is shown in Figure 6.26. The buckle pretensioner has the lowest absolute value with the load limiter 2 case but also the highest absolute value with the load limiter 1 case.

The maximum lap belt force is shown in Figure 6.27. There is not a significant difference between different pretensioners. Chapter 6. Comparison of different pretensioner placement using THUMS 197

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Max Apelvic (Gs) for all cases 100

40 Max Apelvic (Gs)

0 3Pt Both 3Pt Buckle 3Pt Retractor Pretensioner Type

Figure 6.24: THUMS maximum pelvic acceleration for diﬀerent pretensioners.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Max Viscous from Avg deflections (m/s) for all cases 0.35

0.3

0.25

0.2

0.15

0.1 Max Viscous from Avg deflections (m/s) 0.05

0 3Pt Both 3Pt Buckle 3Pt Retractor Pretensioner Type

Figure 6.25: THUMS maximum VC for diﬀerent pretensioners.

The maximum left iliac crest force is shown in Figure 6.28. The retractor pretensioner has a slightly lower force than the other pretensioner types at the lowest load limit level. Chapter 6. Comparison of diﬀerent pretensioner placement using THUMS 198

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Min Viscous from Avg deflections (m/s) for all cases 0

−0.1

−0.2

−0.3

−0.4

−0.5 Min Viscous from Avg deflections (m/s) −0.6

−0.7 3Pt Both 3Pt Buckle 3Pt Retractor Pretensioner Type

Figure 6.26: THUMS minimum VC for diﬀerent pretensioners.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

4 x 10 Max Lap Belt Force (N) for all cases 2.5

1.5

1 Max Lap Belt Force (N)

0.5

0 3Pt Both 3Pt Buckle 3Pt Retractor Pretensioner Type

Figure 6.27: THUMS maximum lap belt force for diﬀerent pretensioners. Chapter 6. Comparison of diﬀerent pretensioner placement using THUMS 199

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Max Left Iliac Crest Force (N) for all cases 1200

1000

800

600

400 Max Left Iliac Crest Force (N)

200

0 3Pt Both 3Pt Buckle 3Pt Retractor Pretensioner Type

Figure 6.28: THUMS left iliac crest force for diﬀerent pretensioners. Chapter 6. Comparison of diﬀerent pretensioner placement using THUMS 200

6.3.6 Risks

The HIC36 risks are shown in Figure 6.29. The retractor pretensioner and the both pretensioner case shows risks below 10% for the load limiter 1 case.

HIC36 Probability of AIS 4+ Injury

AIS 4+

Both 3Pt L. Limit 1

Both 3Pt L. Limit 2

Both 3Pt No L. Limit

Buck 3Pt L. Limit 1

Buck 3Pt L. Limit 2

Buck 3Pt No L. Limit

Retr 3Pt L. Limit 1

Retr 3Pt L. Limit 2

Retr 3Pt No L. Limit

0% 10% 20% 30% 40% 50% 60% 70% 80% 90% 100%

Figure 6.29: THUMS HIC36 risks of injury. Chapter 6. Comparison of diﬀerent pretensioner placement using THUMS 201

The HIC15 risks are shown in Figure 6.30. The AIS 3+ risks are under 10% for all of the load limiter 1 cases. It should be noted that the risks of injury for AIS 4+ are much lower than the HIC36 risks of AIS 4+ risks.

HIC15 Probability of Injury

Fatal MAIS5+ MAIS4+ MAIS3+ MAIS2+ MAIS1+

Both 3Pt L. Limit 1

Both 3Pt L. Limit 2

Both 3Pt No L. Limit

Buck 3Pt L. Limit 1

Buck 3Pt L. Limit 2

Buck 3Pt No L. Limit

Retr 3Pt L. Limit 1

Retr 3Pt L. Limit 2

Retr 3Pt No L. Limit

0% 10% 20% 30% 40% 50% 60% 70% 80% 90% 100%

Figure 6.30: THUMS HIC15 risks of injury. Chapter 6. Comparison of diﬀerent pretensioner placement using THUMS 202

The Nij risks are shown in Figure 6.31. The risks of injury for AIS 4+ injuries are all less than 10%, these risks are similar to the AIS 3+ risks of injury for the HIC15 injury measures.

Nij Probability of AIS 4+ Injury

AIS 4+

Both 3Pt L. Limit 1

Both 3Pt L. Limit 2

Both 3Pt No L. Limit

Buck 3Pt L. Limit 1

Buck 3Pt L. Limit 2

Buck 3Pt No L. Limit

Retr 3Pt L. Limit 1

Retr 3Pt L. Limit 2

Retr 3Pt No L. Limit

0% 10% 20% 30% 40% 50% 60% 70% 80% 90% 100%

Figure 6.31: THUMS Nij risks of injury. Chapter 6. Comparison of diﬀerent pretensioner placement using THUMS 203

The chest 3ms clip risks of injury is shown in Figure 6.32. Unlike the HIC36 injury risks, there were no cases where the risk of AIS 4+ were less than 10%.

Chest 3−ms Clip Probability of Injury

AIS 5+ AIS 4+ AIS 3+ AIS 2+

Both 3Pt L. Limit 1

Both 3Pt L. Limit 2

Both 3Pt No L. Limit

Buck 3Pt L. Limit 1

Buck 3Pt L. Limit 2

Buck 3Pt No L. Limit

Retr 3Pt L. Limit 1

Retr 3Pt L. Limit 2

Retr 3Pt No L. Limit

0% 10% 20% 30% 40% 50% 60% 70% 80% 90% 100%

Figure 6.32: THUMS chest 3ms clip risks of injury. Chapter 6. Comparison of diﬀerent pretensioner placement using THUMS 204

The chest deﬂection risks of injury is shown in Figure 6.33. All the cases has less than 10% chance of AIS 4+ injury.

Sternal Deflection Probability of Injury

AIS 5+ AIS 4+ AIS 3+ AIS 2+

Both 3Pt L. Limit 1

Both 3Pt L. Limit 2

Both 3Pt No L. Limit

Buck 3Pt L. Limit 1

Buck 3Pt L. Limit 2

Buck 3Pt No L. Limit

Retr 3Pt L. Limit 1

Retr 3Pt L. Limit 2

Retr 3Pt No L. Limit

0% 10% 20% 30% 40% 50% 60% 70% 80% 90% 100%

Figure 6.33: THUMS chest deﬂection risks of injury. Chapter 6. Comparison of diﬀerent pretensioner placement using THUMS 205

The CTI risks of injury is shown in Figure 6.34. All cases except the both pretensioner with no load limit has less than a 10% chance of AIS 4+ injury.

Combined Thoracic Index (CTI) Probability of Injury

AIS 5+ AIS 4+ AIS 3+ AIS 2+

Both 3Pt L. Limit 1

Both 3Pt L. Limit 2

Both 3Pt No L. Limit

Buck 3Pt L. Limit 1

Buck 3Pt L. Limit 2

Buck 3Pt No L. Limit

Retr 3Pt L. Limit 1

Retr 3Pt L. Limit 2

Retr 3Pt No L. Limit

0% 10% 20% 30% 40% 50% 60% 70% 80% 90% 100%

Figure 6.34: THUMS CTI risks of injury. Chapter 6. Comparison of diﬀerent pretensioner placement using THUMS 206

6.4 Discussion

The results did not show a significant benefit to using any specific pretensioner set-up. There were minor variations in the injury measures, but the load limiter effect seems much more significant than the pretensioner effect. This suggests that the placement of the pretensioner is not as important as the load level of the load limiter. With the absence of safety concerns, other factors such as cost, aesthetics, and practicality could be considered.

The caveat of the pretensioner implementation in this simulation is that the pretensioner parameters may or may not be realistic. Specifically, the 200 N deactivation force applied to the pretensioning device, which specified that the pretensioner should be deactivated when the force exceeds 200 N, could cause premature abortion of the pretensioning sequence. The original design parameters specifies a 0.1 m pull-in, which was with respect to a Madymo HIII model that was designed with belt slack on the occupant. This was not easily translated to the current model. The 200 N limit was applied as it created the most stable stimulation result, but it might not have corresponded well to a real world pretensioning device.

The results of this study seems to suggest that the beneﬁts of optimizing the pretensioner placement is negligible, and that other means of improving the safety design, such as four-point restraint, should be explored to see if they provide better results.

However, it is also easy to miss effects of the pretensioning device by failure to measure the correct injury measures. The current simulation measures only at a few specific sites on the THUMS body, which may not be enough to discern the full effect of varying the pretensioner placement. Additional measures may prove to show more noticeable effects of the pretensioner placement. Thus the lack of a trend in the pretensioner placement could be the result of not enough data.

There are also more variations on this study that could be implemented in the future. One consideration is the use of double pretensioning (pretensioning at two specific points) could be fire sequentially rather than simultaneously. [83] performed a sequen- tial pretensioning with a delay of 5-11 ms to avoid interaction and found that the chest Chapter 6. Comparison of different pretensioner placement using THUMS 207 acceleration was reduced by 10% - 20%. In fact, pretensioning can be applied at as many points as the belt configuration, such that there could be triple pretensioning for the three-point belt, and quadruple pretensioning for the four-point belt. The deactivation of the pretensioner can also be adjusted. While the 200 N deactivation force was used in this study along with a pull-in curve, it could be changed to a specific pull-in, or a specific time, or a different force. Chapter 6. Comparison of different pretensioner placement using THUMS 208

6.5 Conclusion

The placement of the pretension device did not drastically affect the outcome of this simulation. This suggests that there is not a safety incentive to optimizing the pretensioner placement. However, the lack of significance could be due to an unrealistic simulation design parameters, or insufficient injury measures. Chapter 7

Comparison of four-point and three-point restraint using THUMS

7.1 Introduction

The seatbelt is a relatively old safety invention in the field of automotive safety. As one of the first safety devices on board an automobile, the seatbelt serves a crucial role in protecting the occupant and is still the most effective safety system on an automobile. However, there are still some problem areas with the three-point belt that could be improved. For example, in rollovers and far side crashes, the lap belt portion will usually prevent total ejection, but the occupant may still slip out of the shoulder belt and can suffer head or upper torso injuries. To remedy these problems different belt setups have been proposed. Instead of constraining only one shoulder, proposed seatbelt designs uses two shoulder restraints. These new designs are the four-point seat belt systems.

There have been many proposed modifications to the existing seatbelt systems. The automatic seatbelt systems that have been proposed utilizes an automatic track that fits the shoulder belt system onto the occupant during initial ignition. However, these belts provided conflicting results. While use of the automatic belt did seem to reduce 209 Chapter 7. Comparison of four-point and three-point restraint using THUMS 210 the risk of death, occupants with automatic belts often neglect to buckle the lap belt, which results in less than optimal performance compared to a standard three-point belt. The result is that the occupant should still buckle the lap belt for optimal safety. While the automatic belt does not change the overall design of the three-point belt system, it does emphasize the important of the synergy between the lap belt and the shoulder belt.

The reverse three-point belt utilizes a shoulder belt that originates from the medial aspect of the vehicle instead of the lateral. Preliminary tests shows better behavior in roll overs. However, this leaves the other shoulder unrestrained.

Even with four-point belts, there are design variations. The 3+2 point belt system, while technically not a four-point belt, uses a standard three-point belt as the base and adds an extra shoulder belt. The X4 four-point belt system uses a two criss crossed shoulder belt that meets in the middle of the chest with a lap belt. The V4 four-point belt system uses two shoulder belts that meets the lap belt in the center, creating a V shape in the occupant’s chest. A modiﬁcation of the V4 belt have the shoulder belt meet in the middle of the chest, creating a Y shape on the occupant’s chest. Rouhana has an extensive overview of the existing four-point restraint technologies.

The conclusion from Rouhana’s study implies that the V4 four-point belt has a much lower chest injury risk than the three-point belt and should be studied more in detail.

7.2 Methods

This study uses the THUMS to study the diﬀerence between a standard three-point restraint and the V4 four-point restraint. The simulation setup is shown in Figure 7.1. Load limiter, pretensioner, retractor, and slip rings are all implemented as described in the previous chapters.

The seatbelt material uses standard LS-DYNA 1D seatbelt material with loading and 6 N unloading properties shown in Figure 7.2. The stiﬀness is about 9.5 × 10 strain . The slack length deﬁned was 0 and no belt elements was initially stored in the retractor. Chapter 7. Comparison of four-point and three-point restraint using THUMS 211

Figure 7.1: Model setup and belt conﬁgurations. Left: three-point restraint system. Right: four-point restraint system

kg kg The mass of the belt is deﬁned to be 0.0548 m , which roughly correlates to 1080 m3 for standard belt geometry. The minimum length for the belt is 0.00117 m, or about 10% of the normal element length. The minimum belt length is use to calculate belt slip across slip rings or pull-in into retractors.

Belt Properties 15000 Belt Loading Belt Unloading

10000 Force (N)

5000

0 0 0.2 0.4 0.6 0.8 1 1.2 1.4 1.6 −3 Engineering Strain x 10

Figure 7.2: Seatbelt material force vs engineering strain Chapter 7. Comparison of four-point and three-point restraint using THUMS 212

The four-point belt is shown in Figure 7.3. The four-point belt has four sets of seatbelt elements and one shell element set attached to each. Each shell element set is connected at both ends via the seatbelt element sets. The retractors are connected to the top end of the shoulder belt and are located behind the seat. The sliprings are located at the top of the seat and are slightly in front of the seat back. For the four-point belt, only a shoulder pretensioner placement was modelled.

Sliprings Seatbelt Elements

Shell Elements

Retractors

Buckle

Figure 7.3: Four-point restraint with seat (isometric view)

Both the three-point and the V4 four-point belts were fitted in LS-PREPOST utilizing the belt fitting algorithm. The belt offset from the body was set to 1 mm. The 1D belt elements were connected to 2D shell elements that utilizes simple elastic material with a Young’s modulus of 200 GPa. The average belt element length was around 0.012 m. This was to allow contact between the THUMS and a 2D surface of belt. Preliminary simulations were ran to test that the 2D shells do not reach the slip rings.

The four-point buckle was implemented as a rigid body that attaches to 1D belt elements for the shoulder and the lap belt. It uses 2D shells and a custom inertia tensor:

h 0.2 0.0 0.0 i I = 0.0 0.2 0.0 0.0 0.0 0.2

The crash pulse applied is shown in Figure 7.4. Chapter 7. Comparison of four-point and three-point restraint using THUMS 213

Applied Crash Pulse

300 ) 2 s m

( 200

100 acceleration 0

0 0.05 0.1 0.15 0.2 0.25 time (s)

Figure 7.4: Crash pulse applied: acceleration vs. time

The case matrix for this study is shown in Table 7.1.

Table 7.1: Four-point belt study case matrix

Pretensioner 3-Pt belt 4-pt belt No load limit Buckle, retractor, both pretensioner Retractor pretensioner 3 kN load limit Buckle, retractor, both pretensioner Retractor pretensioner 7.5 kN load limit Buckle, retractor, both pretensioner Retractor pretensioner Chapter 7. Comparison of four-point and three-point restraint using THUMS 214

7.3 Results

7.3.1 Simulation snapshots

The simulation snapshots for the three-point and four-point belts are shown below. The general motion of the THUMS is similar for both belt types, but the three-point belt shows the THUMS veering towards its right as a result of the asymmetrical nature of the three-point belt. There does not seem to be a signiﬁcant diﬀerent in extent of the forward motion.

Figure 7.5: Three-point belt with both pretensioner and 3 kN load limit. Chapter 7. Comparison of four-point and three-point restraint using THUMS 215

Figure 7.6: Three-point belt with buckle pretensioner and 3 kN load limit.

Figure 7.7: Three-point belt with retractor pretensioner and 3 kN load limit. Chapter 7. Comparison of four-point and three-point restraint using THUMS 216

Figure 7.8: Three-point belt with both pretensioner and 7.5 kN load limit.

Figure 7.9: Three-point belt with buckle pretensioner and 7.5 kN load limit. Chapter 7. Comparison of four-point and three-point restraint using THUMS 217

Figure 7.10: Three-point belt with retractor pretensioner and 7.5 kN load limit.

Figure 7.11: Three-point belt with both pretensioner and no load limit. Chapter 7. Comparison of four-point and three-point restraint using THUMS 218

Figure 7.12: Three-point belt with buckle pretensioner and no load limit.

Figure 7.13: Three-point belt with retractor pretensioner and no load limit. Chapter 7. Comparison of four-point and three-point restraint using THUMS 219

Figure 7.14: Four-point belt with retractor pretensioner and 3 kN load limit.

Figure 7.15: Four-point belt with retractor pretensioner and 7.5 kN load limit. Chapter 7. Comparison of four-point and three-point restraint using THUMS 220

Figure 7.16: Four-point belt with retractor pretensioner and no load limit. Chapter 7. Comparison of four-point and three-point restraint using THUMS 221

7.3.2 Head injury measures

The maximum head acceleration is shown in Figure 7.17. The four-point belt shows the lowest value for the load limiter 2 case.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Max Acceleration (Gs) for all cases 80

30 Max Acceleration (Gs)

0 3Pt Both 3Pt Buckle 3Pt Retractor 4Pt Retractor Pretensioner Type

Figure 7.17: THUMS maximum head acceleration (Gs) for diﬀerent pretensioners. Chapter 7. Comparison of four-point and three-point restraint using THUMS 222

The maximum HIC36 is shown in Figure 7.18. The four-point belt shows the lowest value for the load limiter 1 case.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

HIC36 for all cases 1000

900

800

700

600

500 HIC36 400

300

200

100

0 3Pt Both 3Pt Buckle 3Pt Retractor 4Pt Retractor Pretensioner Type

Figure 7.18: THUMS maximum HIC36 for diﬀerent pretensioners.

The maximum HIC15 is shown in Figure 7.19. The four-point belt shows the lowest value for the load limiter 1 case by approximately 100 as compared to the next lowest value. Chapter 7. Comparison of four-point and three-point restraint using THUMS 223

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

HIC15 for all cases 600

500

400

300 HIC15

200

100

0 3Pt Both 3Pt Buckle 3Pt Retractor 4Pt Retractor Pretensioner Type

Figure 7.19: THUMS maximum HIC15 for diﬀerent pretensioners. Chapter 7. Comparison of four-point and three-point restraint using THUMS 224

7.3.3 Neck injury measures

The neck tension is shown in Figure 7.20. The four-point belt has the lowest neck tension with the load limiter 1 case. However, the other load limiter levels show similar values to the three-point cases.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Max Normal Force (N) for all cases 1200

1000

800

600

Max Normal Force (N) 400

200

0 3Pt Both 3Pt Buckle 3Pt Retractor 4Pt Retractor Pretensioner Type

Figure 7.20: THUMS neck tension for diﬀerent pretensioners. Chapter 7. Comparison of four-point and three-point restraint using THUMS 225

The maximum Nij is shown in Figure 7.21. The four-point belt has the lowest Nij with the load limiter 1 case.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Max Nij for all cases 0.18

0.16

0.14

0.12

0.1

Max Nij 0.08

0.06

0.04

0.02

0 3Pt Both 3Pt Buckle 3Pt Retractor 4Pt Retractor Pretensioner Type

Figure 7.21: THUMS Nij for diﬀerent pretensioners. Chapter 7. Comparison of four-point and three-point restraint using THUMS 226

7.3.4 Chest injury measures

The chest 3ms clip is shown in Figure 7.22. There does not seem to be a signiﬁcant diﬀerence between the three-point and the four-point restraints.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

3ms Clip (Gs) for all cases 50

3ms Clip (Gs) 20

0 3Pt Both 3Pt Buckle 3Pt Retractor 4Pt Retractor Pretensioner Type

Figure 7.22: THUMS chest 3 ms clip for diﬀerent pretensioners. Chapter 7. Comparison of four-point and three-point restraint using THUMS 227

The sternal deﬂection is shown in Figure 7.23. The four-point load limiter 1 case has a much lower chest deﬂection than the other cases.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Min Sternal Deflection (%) for all cases 0

−5

−10

−15

−20

Min Sternal Deflection (%) −25

−30

−35 3Pt Both 3Pt Buckle 3Pt Retractor 4Pt Retractor Pretensioner Type

Figure 7.23: THUMS chest deﬂection for diﬀerent pretensioners. Chapter 7. Comparison of four-point and three-point restraint using THUMS 228

The CTI as seen in Figure 7.24 shows the four-point restraint as having higher overall values than the three-point cases.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Max CTI for all cases 1.4

1.2

0.8

Max CTI 0.6

0.4

0.2

0 3Pt Both 3Pt Buckle 3Pt Retractor 4Pt Retractor Pretensioner Type

Figure 7.24: THUMS CTI for diﬀerent pretensioners. Chapter 7. Comparison of four-point and three-point restraint using THUMS 229

The maximum rib stresses are shown in Figure 7.25. The four-point has lower average values compared to the three-point. However, the load limiter levels has less of a diﬀerence.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Max Rib Stress (MPa) for all cases 200

180

160

140

120

100

Max Rib Stress (MPa) 60

0 3Pt Both 3Pt Buckle 3Pt Retractor 4Pt Retractor Pretensioner Type

Figure 7.25: THUMS rib stress for diﬀerent pretensioners. Chapter 7. Comparison of four-point and three-point restraint using THUMS 230

The lumbar forces are shown Figure 7.26. Unlike previous injury measures, the four- point has both the highest value, and the highest value is the load limiter 1 case.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Lumbar Force (N) for all cases 3500

3000

2500

2000

1500 Lumbar Force (N)

1000

500

0 3Pt Both 3Pt Buckle 3Pt Retractor 4Pt Retractor Pretensioner Type

Figure 7.26: THUMS lumbar force for diﬀerent pretensioners. Chapter 7. Comparison of four-point and three-point restraint using THUMS 231

7.3.5 Pelvis injury measures

The abdominal acceleration is shown in Figure 7.27. The four-point belt shows the lowest abdominal acceleration with the load limiter 1 case.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Max Abdominal Acceleration (Gs) for all cases 45

Max Abdominal Acceleration (Gs) 10

0 3Pt Both 3Pt Buckle 3Pt Retractor 4Pt Retractor Pretensioner Type

Figure 7.27: THUMS maximum abdominal acceleration for diﬀerent pretensioners. Chapter 7. Comparison of four-point and three-point restraint using THUMS 232

The maximum pelvic acceleration is shown in Figure 7.28. There does not seem to be a consistent pattern between belt restraint types.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Max Apelvic (Gs) for all cases 100

40 Max Apelvic (Gs)

0 3Pt Both 3Pt Buckle 3Pt Retractor 4Pt Retractor Pretensioner Type

Figure 7.28: THUMS maximum pelvic acceleration for diﬀerent pretensioners. Chapter 7. Comparison of four-point and three-point restraint using THUMS 233

The maximum abdominal VC is shown in Figure 7.29. The four-point load limiter 1 case has the lowest value by a wide margin.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Max Viscous from Avg deflections (m/s) for all cases 0.7

0.6

0.5

0.4

0.3

0.2 Max Viscous from Avg deflections (m/s) 0.1

0 3Pt Both 3Pt Buckle 3Pt Retractor 4Pt Retractor Pretensioner Type

Figure 7.29: THUMS maximum VC for diﬀerent pretensioners. Chapter 7. Comparison of four-point and three-point restraint using THUMS 234

The minimum abdominal VC is shown in Figure 7.30. The four-point restraint has much higher values than the three-point restraints.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Min Viscous from Avg deflections (m/s) for all cases 0

−0.2

−0.4

−0.6

−0.8

−1 Min Viscous from Avg deflections (m/s) −1.2

−1.4 3Pt Both 3Pt Buckle 3Pt Retractor 4Pt Retractor Pretensioner Type

Figure 7.30: THUMS minimum VC for diﬀerent pretensioners. Chapter 7. Comparison of four-point and three-point restraint using THUMS 235

The maximum lap belt force is shown in Figure 7.31. The four-point load limiter 2 case has the lowest value while the four-point load limiter 1 case has the highest value.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

4 x 10 Max Lap Belt Force (N) for all cases 3

2.5

1.5

1 Max Lap Belt Force (N)

0.5

0 3Pt Both 3Pt Buckle 3Pt Retractor 4Pt Retractor Pretensioner Type

Figure 7.31: THUMS maximum lap belt force for diﬀerent pretensioners. Chapter 7. Comparison of four-point and three-point restraint using THUMS 236

The maximum left iliac crest force is shown in Figure 7.32. The four-point restraint has signiﬁcantly lower values than the three-point restraints.

Load Limiter 1 (1500 N) Load Limiter 2 (3000 N) No Load Limiter

Max Left Iliac Crest Force (N) for all cases 1200

1000

800

600

400 Max Left Iliac Crest Force (N)

200

0 3Pt Both 3Pt Buckle 3Pt Retractor 4Pt Retractor Pretensioner Type

Figure 7.32: THUMS left iliac crest force for diﬀerent pretensioners. Chapter 7. Comparison of four-point and three-point restraint using THUMS 237

7.3.6 Risks

The HIC36 risks of injury is shown in Figure 7.33. The load limiter 1 cases all have less than 10% risk of AIS 4+ injury.

HIC36 Probability of AIS 4+ Injury

AIS 4+

Both 3Pt No L. Limit

Both 3Pt L. Limit 1

Both 3Pt L. Limit 2

Buck 3Pt No L. Limit

Buck 3Pt L. Limit 1

Buck 3Pt L. Limit 2

Retr 3Pt No L. Limit

Retr 3Pt L. Limit 1

Retr 3Pt L. Limit 2

Retr 4Pt No L. Limit

Retr 4Pt L. Limit 1

Retr 4Pt L. Limit 2

0% 2% 4% 6% 8% 10% 12% 14% 16% 18%

Figure 7.33: THUMS HIC36 risks of injury. Chapter 7. Comparison of four-point and three-point restraint using THUMS 238

The HIC15 risks of injury is shown in Figure 7.34. The four-point load limiter 1 case has the lowest risk of injury.

HIC15 Probability of Injury

Fatal MAIS5+ MAIS4+ MAIS3+ MAIS2+ MAIS1+

Both 3Pt No L. Limit

Both 3Pt L. Limit 1

Both 3Pt L. Limit 2

Buck 3Pt No L. Limit

Buck 3Pt L. Limit 1

Buck 3Pt L. Limit 2

Retr 3Pt No L. Limit

Retr 3Pt L. Limit 1

Retr 3Pt L. Limit 2

Retr 4Pt No L. Limit

Retr 4Pt L. Limit 1

Retr 4Pt L. Limit 2

0% 10% 20% 30% 40% 50% 60% 70% 80%

Figure 7.34: THUMS HIC15 risks of injury. Chapter 7. Comparison of four-point and three-point restraint using THUMS 239

The Nij risks of injury is shown in Figure 7.35. All the cases have similar risks of injury.

Nij Probability of AIS 4+ Injury

AIS 4+

Both 3Pt No L. Limit

Both 3Pt L. Limit 1

Both 3Pt L. Limit 2

Buck 3Pt No L. Limit

Buck 3Pt L. Limit 1

Buck 3Pt L. Limit 2

Retr 3Pt No L. Limit

Retr 3Pt L. Limit 1

Retr 3Pt L. Limit 2

Retr 4Pt No L. Limit

Retr 4Pt L. Limit 1

Retr 4Pt L. Limit 2

0% 1% 2% 3% 4% 5% 6% 7% 8%

Figure 7.35: THUMS Nij risks of injury. Chapter 7. Comparison of four-point and three-point restraint using THUMS 240

The chest 3ms clip risks of injury is shown in Figure 7.36. The risks of AIS 4+ injury for all cases were over 10%.

Chest 3−ms Clip Probability of Injury

AIS 5+ AIS 4+ AIS 3+ AIS 2+

Both 3Pt No L. Limit

Both 3Pt L. Limit 1

Both 3Pt L. Limit 2

Buck 3Pt No L. Limit

Buck 3Pt L. Limit 1

Buck 3Pt L. Limit 2

Retr 3Pt No L. Limit

Retr 3Pt L. Limit 1

Retr 3Pt L. Limit 2

Retr 4Pt No L. Limit

Retr 4Pt L. Limit 1

Retr 4Pt L. Limit 2

0% 10% 20% 30% 40% 50% 60% 70% 80%

Figure 7.36: THUMS chest 3ms clip risks of injury. Chapter 7. Comparison of four-point and three-point restraint using THUMS 241

The chest deﬂection risks of injury is shown in Figure 7.37. The four-point load limiter 1 case has the lowest risk of injury.

Sternal Deflection Probability of Injury

AIS 5+ AIS 4+ AIS 3+ AIS 2+

Both 3Pt No L. Limit

Both 3Pt L. Limit 1

Both 3Pt L. Limit 2

Buck 3Pt No L. Limit

Buck 3Pt L. Limit 1

Buck 3Pt L. Limit 2

Retr 3Pt No L. Limit

Retr 3Pt L. Limit 1

Retr 3Pt L. Limit 2

Retr 4Pt No L. Limit

Retr 4Pt L. Limit 1

Retr 4Pt L. Limit 2

0% 10% 20% 30% 40% 50% 60% 70% 80%

Figure 7.37: THUMS chest deﬂection risks of injury. Chapter 7. Comparison of four-point and three-point restraint using THUMS 242

The CTI risks of injury is shown in Figure 7.38. The three-point retractor pretensioner with load limit 1 has the lowest risk of injury.

Combined Thoracic Index (CTI) Probability of Injury

AIS 5+ AIS 4+ AIS 3+ AIS 2+

Both 3Pt No L. Limit

Both 3Pt L. Limit 1

Both 3Pt L. Limit 2

Buck 3Pt No L. Limit

Buck 3Pt L. Limit 1

Buck 3Pt L. Limit 2

Retr 3Pt No L. Limit

Retr 3Pt L. Limit 1

Retr 3Pt L. Limit 2

Retr 4Pt No L. Limit

Retr 4Pt L. Limit 1

Retr 4Pt L. Limit 2

0% 10% 20% 30% 40% 50% 60% 70% 80% 90% 100%

Figure 7.38: THUMS CTI risks of injury. Chapter 7. Comparison of four-point and three-point restraint using THUMS 243

7.4 Discussion

The four-point restraint cases, especially the load limiter 1 case, seems to perform better than the three-point restraint cases in most of the injury measures. However, the beneﬁt of the four-point belt is less pronounced compared to the varying the load limit level but better than varying the pretensioner placement.

The simulation snapshots do not show a noticeable decrease in the amount of head excursion, thus it does not seem that the four-point restraint constrains the occupant better. However, it should be noted that the four-point restraint does divide the force on the upper part of the occupant into two contact points rather than one. Such division will reduce the amount of force on each shoulder and the occupant’s forward movement will not be asymmetrical. The three-point snapshots shows that right shoulder point is acting as a pivot point for THUMS and is digging deeply into the shoulder. The four- point snapshots do not show the belt digging into the shoulder as seen in the three-point snapshots.

While the four-point restraint shows similar maximum head accelerations to the three- point restraints, the HIC values are lower for both the HIC36 and HIC15. This indicates that the energy delivered to the head is less for the four-point restraint.

The four-point results shows that the load limiter 1 case has the lowest neck tension and lowest Nij. However, the diﬀerences are not large compared to the load limiter diﬀerences.

For the chest injury measures, the four-point restraints did not show an overall improvement over the three-point restraints. However, both the 3-ms clip and the sternal deﬂection shows a lower value for the load limiter 1 case while the CTI and lumbar force did not show a lower value for any of the four-point cases. There could be an interaction between the load limit level and the restraint type such three-point restraints could perform better than four-point restraints at higher load limit levels.

The pelvic injury measures shows a similarly inconsistent trend. The abdominal acceleration, the maximum VC, and the left iliac crest force shows lower values for the Chapter 7. Comparison of four-point and three-point restraint using THUMS 244 four-point load limiter 1 case. However, the other load limiter cases do not show the same trend.

A problem with the four-point restraint is the relative novelty of the idea in seat belt design. While three-point belt design have been validated and tested extensively in the ﬁeld, there are far fewer four-point belt validation studies. Rouhana et al. did discusses some of the issues extensively with the four-point belt [85, 86] and some of the results presented here corresponds somewhat to those results. The V4 four-point design was shown to create less chest deﬂection and acceleration for the load limiter 1 case. But the load limiter 2 and no load limiter case did not agree with the results of Rouhana et al. The neck loads also did not appear to increase for the four-point restraint as noted in [86]. The lumbar forces did show a slight increase as seen in [86], but was only noticeable for the load limiter 1 case.

More thorough testing of the four-point restraints, along with more variations in the four-point design parameters, can pin point the speciﬁc causes of some of the higher injury measures.

The higher neck forces in the V4 belt as seen in [86] but absent from the results of this study should be investigated. As noted in [87], neck sprains are quite common but minor in professional motor racing that employs more restrictive restraints. As the head and neck are constrained more tightly in the racing car through the use of the Head And Neck Support (HANS), the head is less likely to come in contact with the vehicle interior. However, as a result of this, the tension forces in the neck are often quite high. The four-point restraint with higher load limits could have a similarly more constraining eﬀect on the occupant’s neck and could cause higher risks of neck injuries through tension. The Nij increases as the load limit increases. In [88], a wheel chair bound HIII ATD was subjected to three types of impacts with a curb, and it was found that the four-point restraint was not shown to be more eﬀective than a regular lap belt restraint through the measured HIC and neck forces and moments. The author also postulated that the four-point restraint’s more restrictive nature causes concentration of forces and moments in the neck. Chapter 7. Comparison of four-point and three-point restraint using THUMS 245

The counterpoint to this is that the occupant’s motion inside the vehicle interior is more constrained with higher load limits and the four-point restraint. So it seems that there maybe a balance here that should be tuned.

The four-point restraint also oﬀer beneﬁts in non-frontal crashes such as farside crashes and rollovers. The more restrictive nature of the four-point restraint can prevent un- wanted movement of the occupant in the vehicle interior.

The eﬀect of the four-point restraint on pregnant occupants are also less tested. There have been extensive studies on the eﬀect of the three-point belt on the pregnant occupant [89–99]. Unlike the three-point belt, the four-point belt has a buckle apparatus that is directly anterior to the fetus and could cause higher risks of placental abruption, although the abdominal injury measure do not show higher values than the three-point restraint in this study.

7.5 Conclusion

The four-point restraint shows mostly lower injury measures than the three-point restraint for the load limiter 1 case but not the other load limiter cases. The four-point restraint should be studied more extensively so as to understand the full eﬀect of a more restrictive belt design as well as its eﬀect on pregnant occupant and non-frontal crash scenarios. Bibliography

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