September 2016 

NASA/TM2016-9ROXPH,

Crew Exploration Vehicle (CEV) (Orion) 

Occupant Protection 

Nancy J. Currie-Gregg/NESC 

Langley Research Center, Hampton, Virginia 

Michael L. Gernhardt 

Johnson Space Center, Houston, Texas 

Charles Lawrence 

Glenn Research Center, Cleveland, Ohio 

Jeffrey T. Somers 

KBRwyle, Houston, Texas 
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Hampton, Virginia 23681-2199 

September 2016 

NASA/TM2016-9ROXPH,

Crew Exploration Vehicle (CEV) (Orion) 

Occupant Protection 

Nancy J. Currie-Gregg/NESC 

Langley Research Center, Hampton, Virginia 

Michael L. Gernhardt 

Johnson Space Center, Houston, Texas 

Charles Lawrence 

Glenn Research Center, Cleveland, Ohio 

Jeffrey T. Somers 

KBRwyle, Houston, Texas 
Available from:

NASA Center for AeroSpace Information

7115 Standard Drive

Hanover, MD 21076-1320

443-757-5802

Acknowledgments

The NESC team would like to acknowledge the following individuals who 

contributed to this report. Ms. Jennifer Jadwick, Ms. Jennifer Tuxhorn, Ms. 

Lesley Lee, and Mr. Nick Skytland contributed to the success of this effort as part 

of the Extravehicular Activity Physiology, Systems and Performance Project at 

JSC. In addition, the team would like to thank Dr. Jeff Jones, who contributed 

countless hours of support and was instrumental in the investigation during the 

first 2 years.

The authors would like to acknowledge the support of the HRP at JSC for their 

support and funding for this investigation.

The use of trademarks or names of manufacturers in the report is for accurate reporting and does not 

constitute an official endorsement, either expressed or implied, of such products or manufacturers by the 

National Aeronautics and Space Administration.
iii

Table of Contents

1.0 Executive Summary.......................................................................................................................1

2.0 Problem Description......................................................................................................................3

2.1 Background.........................................................................................................................3

2.2 Current Requirements and the Need for Additional Crew Protection Metrics...................4

2.3 Charter ...............................................................................................................................5

3.0 Data Analysis..................................................................................................................................6

3.1 Human Systems Requirements...........................................................................................6

BDR Model Acceleration Exposure Criteria Historical Background .................................6

Biodynamic Injury Criteria...............................................................................................23

3.2 Biodynamic Modeling ......................................................................................................34

Coordinate System Definitions.........................................................................................34

Model Development .........................................................................................................41

Filtering Specifications.....................................................................................................56

Model Correlation.............................................................................................................60

Conclusions and Recommendations.................................................................................79

3.3 BDR Model Refinements Summary .................................................................................80

Introduction and Background ...........................................................................................80

Methods ............................................................................................................................83

Results .............................................................................................................................88

Discussion.........................................................................................................................91

3.4 THUMS Analyses.............................................................................................................93

WPAFB Sled Test Analysis..............................................................................................93

4.0 Findings and NESC Recommendations.....................................................................................95

4.1 Findings ............................................................................................................................95

4.2 NESC Recommendations .................................................................................................96

5.0 Other Deliverables.......................................................................................................................97

6.0 Lessons Learned...........................................................................................................................98

7.0 Definition of Terms......................................................................................................................98

8.0 References.....................................................................................................................................98
iv

List of Figures

Figure 2.1-1. Six-crew Configuration and Local X-, Y-, and Z-axes..................................... 4

Figure 3.1-1. Photographs of a Volunteer Subject Prepared for an Impact Test 

Supported by Couch with a Form-fitting Liner and a Top View of the 

Liner Formed to Fit the Test Subject ................................................................ 7

Figure 3.1-2. Acceleration Vectors....................................................................................... 10

Figure 3.1-3. Multi-directional Impact Vehicle Showing Seat and Volunteer Subject 

(credit: USAF) ................................................................................................ 11

Figure 3.1-4. Restraint System Configuration...................................................................... 12

Figure 3.1-5. Seat and Restraint System............................................................................... 13

Figure 3.1-6. Photograph of Horizontal Deceleration Test Facility and Subject 

Restrained within the Multi-directional Test Apparatus................................. 13

Figure 3.1-7. Force Vectors Orientation Applied to the Volunteers .................................... 14

Figure 3.1-8. Volunteers Undergoing an Impact Test in a Developmental Seating 

System (credit: USAF)................................................................................... 15

Figure 3.1-9. A Volunteer Being Tested in a Developmental Pressure Suit (credit: 

USAF)............................................................................................................. 16

Figure 3.1-10. Probability of Spinal Injury Estimated from Operational Ejection Seat 

Experience....................................................................................................... 17

Figure 3.1-11. Seat-Back Geometry and Spine Alignment to the Ejection Thrust Vector .... 18

Figure 3.1-12. The F-4 Ejection Seat Injury Rates Associated with the DRI ........................ 19

Figure 3.1-13. FMVSS 208 Neck Compression Limit for Severe Injury............................... 30

Figure 3.2-1. WPAFB Horizontal Sled X (Frontal) Test Condition Coordinate System... 35

Figure 3.2-2. WPAFB Horizontal Sled +X (Rear) Test Condition Coordinate System....... 36

Figure 3.2-3. WPAFB Horizontal Sled Y (Lateral) Test Condition Coordinate System... 37

Figure 3.2-4. WPAFB Horizontal Sled +Z (Spinal) Test Condition Coordinate System .... 38

Figure 3.2-5. WPAFB LS-DYNA® Seat Model Coordinate System ................................... 39

Figure 3.2-6. MADYMO Model Coordinate System........................................................... 40

Figure 3.2-7. BDR Model Coordinate System ..................................................................... 41

Figure 3.2-8. Computer-Aided Design (CAD) Model of Seat Used for WPAFB Sled 

Testing............................................................................................................. 42

Figure 3.2-9. Sled with a 20-G Impulse in the Z-Direction. Sled Motion is to the Right. ...(a) 

Acceleration Pulse, (b) ATD Position ............................................................ 44

Figure 3.2-10. Seat with Side Impact Constraints Added and Oriented for Lateral Impact 

Loads............................................................................................................... 44

Figure 3.2-11. Sled Test with a 10-G Impulse in the X-Direction. Sled Motion Is to the 

Right. (a) Acceleration Pulse, (b) ATD Position ............................................ 44

Figure 3.2-12. LSTC Hybrid III FEM in WPAFB Seat Model.............................................. 46

Figure 3.2-13. FTSS Model in WPAFB Sled Test Seat ......................................................... 47

Figure 3.2-14. FTSS Hybrid III Re-positioned to Better Match Test Video.......................... 49

Figure 3.2-15. Physical ATD Hand and Feet Constraints...................................................... 49

Figure 3.2-16. Comparison between Three Preload Methods................................................ 50

Figure 3.2-17. MADYMO Hybrid III 50th-percentile ATD – Version 6.4 ............................ 53

Figure 3.2-18. WPAFB Configuration ................................................................................... 54

Figure 3.2-19. Isolated THUMS Model Prior to Integration with the Seat............................ 55

Figure 3.2-20. Seat Configuration for Lateral Loading Simulations...................................... 55
v

Figure 3.2-21. Seat Configuration for Anterior/posterior and Inferior/superior Loading 

Pulses.............................................................................................................. 55

Figure 3.2-22. Illustration of Seat Directions for Each Pulse................................................. 55

Figure 3.2-23. Location of the Node Used to Record Head Acceleration in the THUMS 

Model .............................................................................................................. 56

Figure 3.2-24. Limits for +Z-axis Acceleration as a Function of the Time-to-Peak 

Acceleration for Half-sine Acceleration Profile ............................................. 71

Figure 3.2-25. Comparison of +Z BDR Low and Moderate to ATD Lumbar Forces............ 71

Figure 3.2-26. EuroSID/Hybrid III Headrest Lateral Force Comparison............................... 76

Figure 3.2-27. EuroSID/Hybrid III Shoulder Support Lateral Force Comparison................. 77

Figure 3.2-28. EuroSID/Hybrid III Pelvis Support Lateral Force Comparison ..................... 77

Figure 3.2-29. EuroSID/Hybrid III Thigh Support Lateral Force Comparison...................... 78

Figure 3.2-30. EuroSID/Hybrid III Knee Support Lateral Force Comparison....................... 78

Figure 3.3-1. The Impedance Properties of a Volunteer Subject as a Function of 

Frequency........................................................................................................ 81

Figure 3.3-2. Probability of Concussion Based on the HIC ................................................. 83

Figure 3.3-3. The Acceleration-time History at the 6-G Impact Level ................................ 84

Figure 3.3-4. Sideward Impact Test Apparatus with Volunteer Subject.............................. 85

Figure 3.3-5. Side View of the Test Apparatus with a 95th-percentile ATD Subject........... 85

Figure 3.3-6. Positions of the Centers of the Force Cells are Defined with Respect to the 

Coordinate System Origin............................................................................... 86

Figure 3.3-7. Sled Acceleration and Forces Measured by the Upper Body Support Panel.. 88

Figure 3.3-8. Y-Axis Force Measured at Each Subject’s Hip as a Function of Sled 

Acceleration .................................................................................................... 89

Figure 3.3-9. Y-Axis Force at Each Subject’s Shoulder for Each Sled Acceleration 

Condition......................................................................................................... 89

Figure 3.3-10. Maximum Displacement of the Target over Each Subject’s Manubrium for 

Each Sled Acceleration Condition.................................................................. 90

Figure 3.3-11. Current Dynamic Response Acceleration Exposure Limits for Sideward 

Half Sine Impacts............................................................................................ 92

Figure 3.3-12. Dynamic Response Acceleration Exposure Limits Based on NASCAR 

Accidental Shoulder Complex Injuries........................................................... 92

List of Tables

Table 2.3-1. Steering Committee Members .......................................................................... 5

Table 3.1-1. Original HIC 15 msec (HIC15) HSIR Requirements..................................... 25

Table 3.1-2. Proposed Revisions to HIC15 HSIR Requirements........................................ 25

Table 3.1-3. Peak Head Acceleration HSIR Requirements................................................. 26

Table 3.1-4. Proposed Revisions to Peak Head Acceleration HSIR Requirements............ 26

Table 3.1-5. Peak Neck Bending Moment HSIR Requirements......................................... 27

Table 3.1-6. Proposed Revisions to Peak Neck Bending Moment HSIR Requirements.... 27

Table 3.1-7. Peak Neck Axial Tension HSIR Requirements.............................................. 28

Table 3.1-8. Proposed Revisions to Peak Neck Axial Tension HSIR Requirements ......... 28

Table 3.1-9. Peak Neck Axial Compression HSIR Requirements...................................... 29

Table 3.1-10. Proposed Revisions to Peak Neck Axial Compression HSIR Requirements. 30

Table 3.1-11. Peak Neck Shear Force HSIR Requirements.................................................. 31
vi

Table 3.1-12. Proposed Revisions to Peak Neck Shear Force HSIR Requirements............. 31

Table 3.1-13. Peak Lower Extremity Axial Compression HSIR Requirements................... 31

Table 3.1-14. Proposed Revisions to Peak Lower Extremity Axial Compression HSIR 

Requirements .................................................................................................. 32

Table 3.1-15. Chest Sternal-To-Spine Deflection HSIR Requirements................................ 32

Table 3.1-16. Proposed Revisions to Chest Sternal-To-Spine Deflection HSIR 

Requirements .................................................................................................. 32

Table 3.1-17. Restrained Body Movement HSIR Limits...................................................... 33

Table 3.2-1. WPAFB HIA Test Matrix............................................................................... 45

Table 3.2-2. Physical ATD Filter Specifications ................................................................ 58

Table 3.2-3. FTSS ATD Filter Specifications..................................................................... 58

Table 3.2-4. MADYMO ATD Filter Specifications ........................................................... 59

Table 3.2-5. THUMS Model Filter Specifications.............................................................. 59

Table 3.2-6. LSTC Hybrid III ATD Filter Specifications................................................... 60

Table 3.2-7. S&G Classifications........................................................................................ 61

Table 3.2-8. Relation between S&G Numerical Value and Qualitative Description of 

Test-predication Correlation ........................................................................... 61

Table 3.2-9. Lateral Impact Test Matrix ............................................................................. 75

Table 3.2-10. EuroSID Testing Matrix ................................................................................. 76

Table 3.3-1. Dynamic Response Parameters for the Shoulder Impact Model .................... 90

List of Equations

Equation 3.1-1. HIC Formula ................................................................................................... 25

Equation 3.2-1. WPAFB –X to Sled Coordinate Transformation............................................ 35

Equation 3.2-2. WPAFB Sled to SAE Coordinate Transformation ......................................... 35

Equation 3.2-3. WPAFB –X to SAE Coordinate Transformation............................................ 35

Equation 3.2-4. WPAFB +X to Sled Coordinate Transformation............................................ 36

Equation 3.2-5. WPAFB Sled to SAE Coordinate Transformation ......................................... 36

Equation 3.2-6. WPAFB +X to SAE Coordinate Transformation ........................................... 36

Equation 3.2-7. WPAFB –Y to Sled Coordinate Transformation............................................ 37

Equation 3.2-8. WPAFB Sled to SAE Coordinate Transformation ......................................... 37

Equation 3.2-9. WPAFB –Y to SAE Coordinate Transformation............................................ 37

Equation 3.2-10. WPAFB +Z to Sled Coordinate Transformation ............................................ 38

Equation 3.2-11. WPAFB Sled to SAE Coordinate Transformation ......................................... 38

Equation 3.2-12. WPAFB +Z to SAE Coordinate Transformation............................................ 38

Equation 3.2-13. WPAFB Seat SAE to Model Coordinate Transformation .............................. 39

Equation 3.2-14. WPAFB Seat CCIP to SAE Coordinate Transformation................................ 39

Equation 3.2-15. WPAFB Seat FTSS to SAE Coordinate Transformation ............................... 39

Equation 3.2-16. WPAFB Seat THUMS to SAE Coordinate Transformation........................... 39

Equation 3.2-17. SAE to WPAFB MADYMO Model Coordinate Transformation .................. 40

Equation 3.2-18. MADYMO Output to SAE Coordinate Transformation ................................ 40

Equation 3.2-19. SAE to WPAFB BDR Model Coordinate Transformation............................. 41

Equation 3.2-20. BDR Model Output to SAE BDR Model Coordinate Transformation........... 41

Equation 3.2-21. S&G Total Error ............................................................................................. 60

Equation 3.2-22. S&G Magnitude Error..................................................................................... 60

Equation 3.2-23. S&G Phase Error............................................................................................. 60
vii

Equation 3.2-24. S&G Iff Definition ........................................................................................... 60

Equation 3.2-25. S&G Igg Definition .......................................................................................... 60

Equation 3.2-26. S&G Ifg Definition .......................................................................................... 61

List of Appendices (stand alone volume parts 1 and 2)

Volume II Part 1

Appendix A. Occupant Protection Task 1 Orion Charter 

Appendix B. Sprague and Geers Model Correlation Results 

Appendix C. Wright-Patterson Sled Test Model Results 

Volume II Part 2

Appendix D. WPAFB Model Parameter Peak Value Comparison

Appendix E. THUMS Wright-Patterson Sled Test Comparison Results

Nomenclature

AFB Air Force Base

AFRL Air Force Research Library 

AIS Abbreviated Injury Scale

ANSI American National Standards Institute

ASCII American Standard Code for Information Interchange 

ATD Anthropomorphic Test Devices 

ATK Alliant Techsystems

ATLS Advanced Trauma Life Support 

BDR Brinkley Dynamic Response

BMD Bone Mineral Density

CAD Computer-Aided Design

CCD Crew Cursor Device

CCIP Constellation Crew Injury Prediction [ATD]

CEV Crew Exploration Vehicle

CFC Channel Frequency Class

CM Crew Module

CPU Central Processing Unit

CSDM Cumulative Strain Damage Measure

CSSS Constellation Space Suit System 

CxP Constellation Program

D Dimensional

DAC Design Analysis Cycle 

DoD Department of Defense

DOF Degree of Freedom

DRI Dynamic Response Index

DRx X-axis Dynamic Response

DRy Y-axis Dynamic Response

DSS Decelerator System Simulation

DU Display Unit

DXA Dual-energy X-ray Absorbance

EFA Exploratory Factor Analysis
viii

EOM End of Mission

EuroSID European Side-Impact Dummy

EVA Extravehicular Activity

F/B Fighter/Bomber

FAA Federal Aviation Administration

FACB Flight Activities Control Board 

FE Finite Element

FEM Finite Element Model

FMVSS Federal Motor Vehicle Safety Standard

FR Front Right

FS Flight Status

FTSS First Technologies Safety Systems

G Gravitational constant

GM General Motors

GRC Glenn Research Center

GSI Gadd Severity Index

HANS® Head and Neck Support®

HIA Horizontal Impact Accelerator

HIC Head Injury Criteria 

HRP Human Research Program 

HSIR Human-Systems Integration Requirements

IARV Injury Assessment Reference Values

IRC Injury Reference Criteria 

IRL Indy Racing League

IS Injury Severity 

IWE Independent Witness® Encrypted Format File

IWI Independent Witness® Unencrypted Binary File

JSC Johnson Space Center

LaRC Langley Research Center

lb pound

LM Lockheed Martin

LOC Loss of Consciousness

LS-DYNA® Finite Element Analysis Software by Livermore Software Technology 

Corporation

LSTC Livermore Software Technology Corporation

MADYMO MAthematical DYnamic Modeling

MATLAB® MATrix LABoratory data analysis program from The Mathworks

mm millimeter

MOB Medical Operations Board 

mph miles per hour

msec millisecond

MTSO Management and Technical Support Office

N Newton

NASA National Aeronautics and Space Administration

NASCAR National Association for Stock Car Auto Racing

NESC NASA Engineering and Safety Center
ix

NFS Not Further Specified

NHTSA National Highway Traffic Safety Administration

Nm Newton meter

NRB NESC Review Board

OOB Out-of-Board

OPRA Order Probit Regression Analysis

ORIS Operationally Relevant Injury Scale 

OSU Ohio State University

PMHS Post-Mortem Human Subject

PRA Probabilistic Risk Assessment

S&G Sprague-Geers

SAA Space Act Agreement 

SAE Society of Automotive Engineering

SAR Search and Rescue

SC Suit Connector

SE Self-Egress

SID Side-Impact ATD

SIMon Simulated Injury Monitor 

SMCCB Space Medicine Configuration Control Board 

SME Subject Matter Expert

SRD System Requirement Document 

SSP Space Shuttle Program

STD-DEV Standard Deviation

TASS TNO Automotive Safety Solutions

TBI Traumatic Brain Injury

THUMS Total HUman Model for Safety

TMA Transport Modified Anthropometric

US United States

USAF United States Air Force

USN United States Navy

USRA Universities Space Research Association

VDT Vertical Drop Tower

WP Wright Patterson

WPAFB Wright-Patterson Air Force Base

WSTC Wayne State Tolerance Curve
1

Abstract

Dr. Nancy J. Currie, of the NASA Engineering and Safety Center 

(NESC), Chief Engineer at Johnson Space Center (JSC), requested an 

assessment of the Crew Exploration Vehicle (CEV) occupant protection as 

a result of issues identified by the Constellation Program and Orion 

Project. The NESC, in collaboration with the Human Research Program 

(HRP), investigated new methods associated with occupant protection for 

the Crew Exploration Vehicle (CEV), known as Orion. The primary 

objective of this assessment was to investigate new methods associated 

with occupant protection for the CEV, known as Orion, that would ensure 

the design provided minimal risk to the crew during nominal and 

contingency landings in an acceptable set of environmental and spacecraft 

failure conditions. This documents contains the outcome of the NESC 

assessment. NASA/TM-2013-217380, “Application of the Brinkley 

Dynamic Response Criterion to Spacecraft Transient Dynamic Events.” 

supercedes this document.

1.0 Executive Summary

The NASA Engineering and Safety Center (NESC), in collaboration with the Human Research 

Program (HRP), investigated new methods associated with occupant protection for the Crew 

Exploration Vehicle (CEV), known as Orion.

Reliable injury predictive tools and injury criteria are required to ensure that human-rated 

spacecraft (e.g., Orion or a commercial space venture) be designed with the appropriate level of 

occupant protection [ref. 1]. It is important the occupants be protected from injury and the vehicle 

is not designed with excessive protection that leads to unnecessary vehicle weight and/or 

complexity. While conventional tools used in the aerospace community have led to safe designs, 

retirement of the Space Shuttle Program (SSP) and transition to a vehicle whose primary landing 

mode will probably be water based, with contingency land landings using parachutes, will 

introduce many new challenges. Approaches to ensure safety, such as those used in the commercial 

aviation and automotive industries, can provide a foundation for human-rated spaceflight for these 

new vehicles. However, their application to the Orion Project or a commercial venture requires 

modification and study.

The NESC, HRP, and the Constellation Program (CxP) developed injury criteria and methods for 

predicting injury for the Orion crew module (CM) and for future commercial space programs. This 

working relationship was somewhat different from the independent way in which the NESC 

normally operates. However, due to the complexity of problems and the demand for a strong and 

diverse skill set, it was determined that a combined effort was required.

The core of the occupant protection criteria is based on the Brinkley Dynamic Response (BDR) 

model and associated injury criteria limits. These criteria are being used by NASA and the United 

States (US) military to determine the injury risk or adverse physiological response for vehicle 

occupants based on seat acceleration. While the BDR is useful for generating an overall estimate 

of the probability of injury, the criteria have limitations. The BDR model and associated criteria 

assume a basic seat geometry, restraint, and head protection, and are therefore only an 

approximation for other seat designs and protection systems. Furthermore, the BDR model cannot 
2

be used to predict risk when improvements are made to the seated environment. Risk can only be 

lowered by reducing the driving loads into the seat. In addition, the BDR model cannot predict 

suit-induced injury (e.g., localized blunt trauma and point loading) related to rigid elements or 

interference with the restraints.

The US Air Force Research Laboratory (AFRL) at the Wright-Patterson Air Force Base (WPAFB) 

in Dayton, Ohio, evaluates body restraints and support systems using volunteer subjects exposed 

to anticipated sub-injury impact conditions. These evaluations are conducted by comparison with 

the baseline BDR body support and restraint systems and evaluation of the likelihood of injury 

using established biodynamic data. Systems proposed are assessed to reduce the likelihood of 

injury and to reduce tolerance to impact conditions. 

As a result of the present work, additional injury criteria specific to the head, neck, and legs have 

been incorporated into the Human System Interface Requirements (HSIR) [ref. 1]. These 

additional injury criteria complement the BDR criteria already incorporated in the requirements. 

These supplemental injury criteria were developed primarily for the automotive industry and are 

regularly used to ensure automotive safety. The injury criteria used by the automotive industry are 

designed for automobile accidents and considerably higher allowable probability of injuries than 

are acceptable for NASA and manned vehicle landings, so judgment was used to extrapolate the 

criteria for use for manned spaceflight. The automotive industry criteria provided a better 

understanding of the location and type of occupant injury and the effect of seating conditions and 

occupant protection. When used in conjunction with the BDR criteria it can provide a more 

complete assessment of occupant protection. 

Before the National Highway Transportation Safety Administration (NHTSA) for automotive 

safety and the Federal Aviation Administration (FAA) for commercial aircraft safety implement 

standards, they anchor their criteria with human test data collected from accidents and laboratorybased human studies. NASA is planning to leverage from these extensive bodies of work. 

However, NASA should perform human testing specific to the environment and loadings of 

planned manned vehicles to anchor the occupant protection injury criteria with relevant test data. 

In the past, NASA deemed this type of testing necessary and performed it for the Apollo Program; 

similarly, the Russians performed this type of testing for Soyuz. Currently, NASA has no plan to 

obtain human test data for Orion or other future planned manned vehicles.

In addition to the BDR criteria and occupant protection criteria, body movement requirements 

were added to the HSIR. These requirements complement the biodynamic injury criteria by 

establishing requirements that limit relative body motions that could result in injury. Limiting body 

movement is a sound design practice and will ultimately lead to the necessity for head and torso 

restraint systems, such as a Head and Neck Support (HANS®), side supports, and at least a partially 

conformal seat design. Furthermore, the BDR criteria were developed based on data collected for 

specific seat configurations that provided lateral support and body movement constraints. For the 

BDR criteria to be applicable to manned spacecraft, body movement requirements are a necessity 

to provide a similar seated configuration as was used to develop the BDR criteria. A similar seat 

configuration would be one that provides the same level of body constraint as is provided by the 

seat used during the development of the BDR model and associated injury criteria. 

Implementation of the occupant protection and body movement criteria requires the use of physical 

and numerical anthropomorphic test devices (ATD) to extract the occupant response in realistic 

seated environments and loading conditions. Numerical models of the Hybrid III ATDs were 
3

evaluated for assessing potential crew injuries through extensive sled testing at AFRL and 

numerical modeling. The First Technology Safety Systems (FTSS) model of the Hybrid III was 

recommended for use by NASA and the Orion Project since this model performed best among the

models assessed. However, the FTSS model did not consistently produce correlation results to 

within the desired 20 percent of the test data. Overall, the FTSS Hybrid III model performed best 

among the three ATD numerical models. It is recommended that before a numerical model is used 

to perform a final assessment of occupant injury, the model should be validated through physical 

testing. This testing should be performed in a similar physical environment, including similar or 

identical seat, harness, helmet, and suit and loading conditions, comparable to the loading 

conditions that will ultimately be used for the occupant protection assessments. 

2.0 Problem Description

2.1 Background

The CxP plans to replace the SSP with a CEV CM named “Orion.” Orion is similar to Apollo, but 

larger, and will carry a crew of four to six members while Apollo was designed for a crew of three. 

The Orion CM will descend to Earth with a three-parachute recovery system in the event of an 

emergency escape during the launch phase or during the final phase of the CM’s return to Earth. 

The CM is designed to be capable of a primary water landing and contingency land landings. 

Landing systems must be designed for multi-terrain impacts. Earth terrain is highly variable at the 

proposed landing sites; soil at Edwards Air Force Base (AFB) in California is typically hard, and 

soil at Carson Sink in Nevada is often soft. During an emergency escape from the launch pad, the 

CM will be lifted from the launch vehicle by a launch abort system with a trajectory designed for 

a water landing. Landing and recovery, and crew protection during water or land landings are a 

design challenge, and despite years of successful space missions cannot be considered routine. To 

keep the size of parachute systems reasonable and to optimize the weight of the landing system, 

the terminal velocity of the Orion CM parachute recovery systems was designed to provide a 

descent velocity at landing of approximately 24 ft/sec nominal to 33 ft/sec contingency (16 to 22.5 

mph, 7.3 to 10.1 m/s). In addition to the terminal descent (vertical) velocity, the parachutes and 

CM will move horizontally depending on the horizontal wind velocity. The CM touchdown 

velocity is typical of crash impact velocities of small aircraft and helicopters. Thus, the CM’s 

impact velocity may produce injuries without some type of mitigation for impacts onto land. With 

proper pitch attitude into low sea states, water landings do not require additional mitigation for 

nominal landings since the water acts as a natural impact attenuator. The Apollo capsule was 

designed to impact the water with a parachute hang angle of 27 degrees. Since the Apollo entered 

the water pitched at an angle, the acceleration experienced was nominally 10 G or less for a 

30-ft/sec (9.1-m/s) impact [ref. 2]. The Apollo spacecraft had crushable ribs and a crew seat pallet 

with stroking energy absorbers to alleviate contingency impacts. However, a land landing of 

Apollo in an abort condition or a flat 0-degree attitude water impact with one parachute fully 

opened was expected to result in occupant injuries. Since the probability of a pad abort with a land 

impact was considered low, the CxP accepted those risks. 

Under ideal nominal conditions, a land landing with retro-rockets or impact-attenuating airbags, 

or a water landing will result in tolerable accelerations to the crew if the seats and restraint systems 

are properly designed. For contingency landing conditions, a stroking crew seat pallet similar to 

the Apollo design is planned for the Orion CM to reduce the impact accelerations to conditions 

that are tolerable to the crew and that will prevent injury. However, some mitigation for more 
4

severe contingency impacts can be designed with little weight penalty if a systems approach is 

taken early in the design phase. 

The primary purpose of this assessment was to assess the use of finite element (FE) seat and 

occupant simulations for assessing injury risk for Orion. These FE seat and crash ATD models 

allow for varying the occupant restraint systems, cushion materials, side constraints, flailing of 

limbs, and detailed seat/occupant interactions. Also included in this assessment is a discussion of 

impact injury criteria used by the automotive industry to assess crash ATD response and how these 

criteria can be used for Orion occupant protection. With properly designed seats (side supports and 

restraints), relatively high accelerations can be tolerated with no or minimal injury, especially for 

impacts with the crew lying on their backs in a seat oriented as shown in 

Figure 2.1-1. The local inertial axis system has the +X-axis forward, the +Z-axis pointed toward 

the head, and the +Y-axis to the left. For accelerations along the +X-axis, which would occur for 

a flat impact, the inertial response to a spine-to-chest acceleration is called ‘eyeballs in.’ An inertial 

response to an acceleration along the –X-axis (chest-to-spine) is called ‘eyeballs out.’ An inertial 

response to an acceleration along the +Z-axis (seat) is called ‘eyeballs down’, and an inertial 

response to an acceleration pointing in the –Z-axis is called ‘eyeballs up.’ The seat coordinate 

system is a local system and is a non-inertial axis system. Hence, angular rotations with respect to 

an inertial system must be considered. 

Figure 2.1-1. Six-crew Configuration and Local X-, Y-, and Z-axes

2.2 Current Requirements and the Need for Additional Crew Protection Metrics

The CxP HSIR document [ref. 1] includes provisions for protecting the crew during dynamic 

phases of flight, including linear and rotational acceleration limits and BDR acceleration exposure 

limits. Although these requirements add a level of crew protection, it became apparent to the Orion 

Project Office that additional requirements may be necessary. In addition, the automotive and car 

racing industry has shown considerable improvements in occupant protection over the last 10 

years. In an attempt to leverage this expertise to improve crew protection during Orion launch and 

landings, a tiger team consisting of NESC and CxP personnel was formed to further investigate 

occupant protection principles and their applicability to the spaceflight environment. 

+Z(up)

+X (back)

+Y(right)
5

2.3 Charter

In 2007, the Orion Project was in the early stages of design analysis cycle (DAC) #2 and was 

challenged with evaluating the 606C vehicle design capability for nominal water landings and 

contingency land landings. To integrate this technical effort, an Orion Landing Strategy Tiger 

Team was established per Orion Project Guidance Memo ZV-07-034 (see Appendix A). 

Additionally, an Occupant Protection Industry Day was held in July 2007 to bring together experts 

in occupant protection principles, research, and analysis from leading industry, research 

institutions, and government Agencies. 

In the subsequent months, a steering committee (Table 2.3-1) was assembled to oversee the design 

and analysis activities being performed for the landing event during DAC #2. These activities 

included:

 assessing current vehicle design for water landing and contingency land landing events,

 assessing the vehicle design susceptibility to roll control and tip over,

 reviewing methods for assessing occupant injury during ascent/aborts/landings,

 developing an alternate seat/attenuation design solution, which improves occupant 

protection and operability, and

 testing the seat/attenuation system designs to ensure valid results. 

Table 2.3-1. Steering Committee Members

Chair Irene Piatek

Vehicle Integration 

Office/Operations

Angela Hart

Systems Engineering 

Integration Team

Chris Johnson

CB Lee Morin

Health and Medical 

Technical Authority

Mark Jernigan

Safety & Mission Assurance John Trainor

Orion Chief Engineer Julie Kramer White

Extravehicular Activity Miriam Sargusingh

NESC Nancy Currie

The steering committee chartered three tasks to coordinate these activities. The first team 

(Task 1) was chartered to address requirements and analysis techniques; this assessment 

encompasses the work of the Task 1 team. The Task 2 team examined alternate seat/attenuation 

solutions, and the Task 3 team performed the seat/attenuation system testing. 

The Task 1 team was chartered with the following leadership, description, and requirements:

 Lead: JSC Human Adaptation and Countermeasure Office  Mr. Mike Gernhardt.

 Team: NESC, industry and academic experts, key stakeholders (Crew Office, JSC 

Engineering, and JSC Space Medicine and Human Engineering).

 Steering Committee Point of Contact: Dr. Nancy Currie.

 Task Description: Determine appropriate methods for modeling and prediction of 

potential crew injuries. Activities include conducting data mining of injury databases 

(e.g., NASCAR, Crash Injury Research Engineering Network, military, etc.) and 
6

assessing impact simulation and injury prediction methodologies. Recommend 

techniques/changes to requirements appropriate for Orion, including stated limitations.

 Products:

1. Periodic briefings to Steering Committee/CEV Project Office.

2. Given driving load cases from the Orion Project Office, analyze Orion baseline 

design (606C) and alternate configurations developed by Task 2 team.

3. Final report included recommendations for:

a. Standard analysis techniques and models.

b. Impact acceleration/occupant injury criteria appropriate for use by Orion.

This document fulfills the final product listed above as the final report for Task 1 and includes the 

recommendations and outcomes of this effort.

3.0 Data Analysis

3.1 Human Systems Requirements

BDR Model Acceleration Exposure Criteria Historical Background

Early Use of Simplified Exposure Limits

The BDR criteria were developed as a result of an evolutionary process to define the human 

dynamic response to, and exposure limits for, short duration accelerations (i.e., <500 msec) 

associated with spacecraft landing and emergency escape system performance. During the era of 

the Mercury, Gemini, and Apollo Programs, the encapsulated ejection seats for the B-58, XB-70 

aircraft, and the cockpit crew escape system for the fighter/bomber (F/B)-111 aircraft for the 

USAF, occupant protection was a concern, and work was done to further develop and improve the 

impact exposure criteria. Acceleration limits were established [refs. 3 and 4] in terms of rate of 

onset, acceleration amplitude, and duration for areas known to be within voluntary tolerance and 

those known to cause moderate to severe injury. These acceleration limits were based upon the 

research of John P. Stapp and his contemporaries [ref. 4] using military volunteers, animal 

surrogates, and the results of accidental exposures of humans. However, the fitting of the 

trapezoidal acceleration-time histories was inadequate to assess the injury risk due to the complex, 

multi-directional landing impacts of these crew systems. The trapezoidal pulses were developed 

for ejection seats where the acceleration-time histories of the ejection catapult, rocket thrust, and 

deceleration due to aerodynamic forces could be controlled in terms of these acceleration-profiles.

Examples of other key research results include that of Beeding and Mosely [ref. 5], who conducted 

experiments using a horizontal deceleration facility to study the responses of volunteers to impact 

in forward-facing and rearward-facing seats with lap belt and shoulder harness restraints. They 

reported severe shock and repeated syncope with myalgia requiring one volunteer to be 

hospitalized for 5 days following his exposure to a 40-G sled acceleration 

with a rate of onset of 2,139 G/sec and a velocity change of 48.5 ft/sec in a rear-facing seat. 

Dr. Mosley reported that the subject may not have survived without immediate medical care. Prior 

exposures of volunteers in this position ranged from 25–40 G at onset rates from 

1,034–2,139 G/sec for durations of 50 to 190 msec. Impact velocities were below 50 ft/sec. 

The acceleration-time histories could be defined in terms of the existing trapezoidal limits, 

although a half-sine pulse shape more approximated the applied acceleration.
7

Addressing Complex Multi-directional Accelerations

Based upon survived accidents, experimental data from tests with volunteers, tests with animals 

[ref. 6], and theoretical work based on the laws of mechanics, AFRL laid out an initial plan to 

investigate the acceleration profiles with high rates of onset using experimental and analytical 

approaches. NASA joined this effort to investigate human tolerance to landing impact conditions 

with rates of onset higher than had previously been considered safe. 

To address this problem, Headley et al.[refs. 7 and 8] and Brinkley et al. [refs. 9 and 10] conducted 

experiments with volunteer subjects to explore the human responses to the complex waveforms 

associated with emergency escape capsule impact and the Project Mercury module launch abort 

landing impact. These experimental data demonstrated several relationships between the higher 

rates of acceleration onset, acceleration magnitude, and impact velocity change in the +Z- and +Xaxes. During these experiments, Brinkley et al. [ref. 9] investigated the beneficial effects of an 

individually contoured body support constructed of fiberglass compared to a body support liner 

that could be formed to fit the occupant (Figure 3.1-1).

Figure 3.1-1. Photographs of a Volunteer Subject Prepared for an Impact Test Supported by Couch 

with a Form-fitting Liner and a Top View of the Liner Formed to Fit the Test Subject

During this same period, Payne [ref. 11] used the extensive human and animal impact exposure 

data assembled by Eiband [ref. 4], and Goldman and von Gierke [ref. 12] to study the feasibility 

of developing simple, lumped-parameter dynamic models that might describe the human response 

and likelihood of injury based on the experimental data. Payne developed numerical models of 

restraint and support systems to show the effects of their mechanical properties on the human 

responses to complex acceleration profiles.

In 1962, NASA began to organize efforts to design the Apollo crew systems. One of the efforts 

included the development of CM landing impact system human tolerance criteria. Initially, crew 

protection system designers were forced to accept a maximum acceleration of 10 G and a rate of 

onset of 250 G/sec as the limiting crew acceleration exposure criteria. A major analytical and 
8

experimental program was organized by the NASA Manned Spacecraft Center (i.e., JSC) and was 

implemented by AFRL, the Naval Air Engineering Center, and the Armed Forces Institute of 

Pathology to identify potential critical modes of injury and likely adverse physiological responses, 

and to explore the effects of higher level multi-directional impacts “to assure astronaut 

functionality immediately after landing” [ref. 13].

The most immediate problem was to explore the human responses to sideward (Y-axis) impact. At 

that time, knowledge of human response to sideward acceleration was limited to the results of 

centrifuge experiments conducted by AFRL to sustained acceleration levels to 10 G [ref. 14]. In a 

study using a centrifuge, Hershgold examined the displacement of the volunteer subjects’ internal 

organs during steady-state acceleration. His radiographs showed extensive displacement of 

thoracic and abdominal viscera to the dependent side of the body with acceleration as small as 6 

G. Stapp [ref. 15] had not observed injuries to chimpanzees at calculated accelerations to 47 G 

with durations estimated to be 140 msec at the peak G level.

During this same time, impact tests of the B-58 encapsulated ejection seat were conducted with 

volunteer subjects [ref. 16]. The capsule was dropped with vertical velocities to 27 ft/sec (8.2 m/s) 

and landed on its back bulkhead with its occupant in a semi-supine position. The tests were 

conducted without detectable injury. During tests not reported in this article, the capsule was 

impacted with horizontal velocities to 34 ft/sec (10.4 m/s) in the sideward direction. An 

undetermined amount of energy was dissipated in tumbling and skidding of the capsule, making 

use and extrapolation of these data difficult. However, these tests demonstrated the tolerance to 

sideward impact was probably higher than the 10 G that had been investigated under steady-state 

acceleration of 10 G by Hershgold [ref. 14].

In preparation for a joint agreement with NASA to study the feasibility of higher impact exposure 

limits, Robinson et al. [ref. 17] conducted impact tests using Rhesus monkeys with sideward 

impacts to 75 G at terminal velocities to 32 ft/sec (9.8 m/s), with and without contoured body 

support systems. No post-mortem injuries were observed. Electrocardiographic evidence of 

transient abnormalities in conduction and rhythm were found following higher-level accelerations 

and impact velocities. Radiographs taken before and after the impacts showed an increase in the 

total heart shadow on the dependent side of the midline. Sequential radiographs revealed that the 

heart returned to normal within 3 hours after impact. 

Based on the agreement between NASA and the Department of Defense (DoD), Schulman et al., 

[ref. 18] explored the effects of downward acceleration (Z- axis) to 18.5 G with rates of onset to 

1,540 G/sec. This research supplemented and extended USAF research that had been conducted 

with animals and volunteers to define criteria for downward ejection seats [ref. 19]. 

AFRL conducted impact tests to investigate the dynamic responses of military volunteers using a 

vertical deceleration tower. Acceleration, forces, physiological reactions, and subjective responses 

of volunteers were measured for sideward (Y-axis) impacts ranging to 21.5 G with rates of onset 

to 1,350 G/sec and impact velocities to 25.5 ft/sec (7.8 m/s) [ref. 20]. An individually fitted, semirigid body support was used with torso and extremity restraints. A Mercury pressure-suit helmet 

weighing 4.4 lbs (21 kg) was worn by the subjects in these tests. Plastic foam (0.5-inch-thick, 

13-mm) pads were placed behind the helmet earphones and across the chin bar. The helmet was 

prevented from pitching motion, but not vertical motion by a 1.75-inch (44-mm) strap. The helmet 

microphone was removed and the impacts were tolerated by the male volunteer subjects.
9

A second series of impact tests jointly sponsored by the USAF and NASA were conducted to study 

the effects of seven impact vector directions and six acceleration profiles ranging in acceleration 

from 3–26 G, impact velocities ranging from 5–28 ft/sec (1.58.5 m/s), and onsets ranging from 

200–2,000 G/sec. The seven seat orientations, including forward, upward, and sideward 

(right and left) acceleration components in 45-degree increments were used on a vertical 

deceleration tower. A drawing of the orientations, designated as vectors A through G, is shown in 

Figure 3.1-2 [ref. 21]. Force, acceleration, and physiological data were collected and analyzed. 

The volunteers were contained by an experimental restraint system and supported by a flat seat 

and seat back, and side panels to support the pressure-suit helmet, shoulders, and upper legs. The 

liner filled with small beads was not used in this second series of tests. The subjects needed 

assistance to ensure that the liner was adequately formed about their bodies and they were more 

confident in the support provided by a rigid seat. A test subject is shown restrained within the 

multi-directional impact test vehicle in Figure 3.1-2. The seat and its frame were attached to the 

AFRL vertical deceleration tower by rods that were instrumented with force measurement cells 

(see Figure 3.1-3). The restraint system is shown in Figure 3.1-4.

The subject panel included 20 male subjects. The acceleration and force data were subjected to 

extensive mathematical analysis based upon the Fourier Transformation [ref. 22]. The relationship

between the measured forces and impact velocity, called mechanical impedance, was used to 

identify the human body dynamic response properties in terms of resonances similar to those of 

linear, second-order, spring-mass-damper systems. Broad, low-amplitude resonances were 

identified at 3.5, 5.5, 7.2, and 11.7 Hz. There was no gross distinction in the impedance magnitude 

or phase angle among the orientations studied. The subject impedance magnitude increased 

linearly with frequency to about 35 Hz. The analysis was not valid beyond this frequency because 

the velocity pulse did not contain significant frequency components beyond 35 Hz. Physiological 

and subjective response data that were collected indicated abrupt cardiac rhythm changes at the 

higher acceleration levels, wind knocked out, and various sites of transient pain, including the 

head. The investigators concluded the head would present a problem at higher impact levels. 

However, no voluntary tolerance limit was specified [ref. 21].
10

Figure 3.1-2. Acceleration Vectors
11

Figure 3.1-3. Multi-directional Impact Vehicle Showing Seat and Volunteer Subject (credit: USAF)
12

Figure 3.1-4. Restraint System Configuration

Further tests were accomplished by the USAF, as part of its joint effort with NASA, using a 

horizontal deceleration track at Holloman AFB, New Mexico. Stapp and Taylor [ref. 23] explored 

16 different seat orientations with 58 volunteers. A total of 146 tests were conducted to explore 

the effects of accelerations ranging in magnitude from 10–25 G, rates of onset from 1,000–2,000 

G/sec, and durations from 60 to 130 msec. A pneumatic piston accelerated a sled and subject, 

allowing it to coast and impact a water-filled decelerator at velocities ranging from 1846 ft/sec 

(5.514 m/s). The subjects wore Mercury pressure-suit helmets and were restrained by the same 

developmental restraint system used by Weis et al. [ref. 21] at the WPAFB. The seat was 

constructed of flat panels to support and constrain the subject’s movement in the rearward and 

sideward directions [ref. 23] (see Figure 3.1-5 and Figure 3.1-6). For smaller subjects, panels were 

used to ensure that there were no gaps between the subject and the seat.
13

Figure 3.1-5. Seat and Restraint System

Figure 3.1-6. Photograph of Horizontal Deceleration Test Facility and Subject Restrained within the 

Multi-directional Test Apparatus
14

All body positions and impact conditions were within voluntary tolerance, except the forwardfacing 45-degree reclining position at 25.4 G sled acceleration, with a rate of onset of 960 G/sec 

and a 97-msec duration. The subject experienced soft tissue pain and stiffness in the area of the 6th

through 8th thoracic vertebrae for 60 days. Bradycardia was experienced in 55 electrocardiograms 

immediately after impact. The bradycardia was triggered by headward impact vectors due to 

stimulation of the carotid sinuses, dropping the heart rate as much as 90 beats per minute for 10–

30 seconds. The restraint looseness was found to be a contributing factor. Persistent or severe pain 

was absent in 146 of 163 tests.

Brown et al. [ref. 24] extended the research of Weis et al. [ref. 21], and Stapp and Taylor [ref. 23] 

by conducting 288 impact tests with volunteers using the horizontal deceleration track to explore 

the responses of the nervous, cardiovascular, respiratory, and musculoskeletal systems. Twentyfour acceleration vector directions in 45-degree increments were studied using the same helmet, 

body support, and restraint in the two studies described above (see Figure 3.1-7). Seventy-nine 

subjects participated in this experimental study. At each seat orientation, the impact level was 

increased in increments of 2–5 G until the occurrence of an adverse reaction based upon the 

evaluation of the subjective, clinical, and physiological responses by the medical monitor, 

determined the maximum impact levels. Two tests were accomplished at each combination of 

position, acceleration level, and onset rate. The sled acceleration ranged from 5.5–30.7 G, the rate 

of onset was varied from 300–2,500 G/sec, and the impact velocity ranged from 9.3–45 ft/sec 

(2.813.7 m/s). The highest impact levels were achieved in positions 19 and 23, while lower levels 

were limited to 11.8 and 11.1 G in positions 9 and 13, respectively. The +Z- and –Z-axes 

orientations of positions 9 and 13 had been explored more thoroughly in the earlier tests by 

Schulman et al. [ref. 18], Weis et al.[ref. 21], and Stapp and Taylor [ref. 23]. 

Figure 3.1-7. Force Vectors Orientation Applied to the Volunteers

The neurological effects observed were momentary stunning and disorientation, which lasted no 

more than 2 minutes. In one case, this response was noted at an impact level as low as 17.4 G, 

where the impact force vector was in orientation 17. Headaches lasting several hours were noted 

in nine tests. Shortness of breath and chest pain were experienced by the subjects in more than 
15

one-half of the higher-level tests. Musculoskeletal complaints resulted primarily from muscle

spasms and strains of the neck, back, and lower extremities.

Subsequently, additional tests were conducted with volunteers in developmental seats, restraints, 

and pressure suits, as shown in Figure 3.1-8 and Figure 3.1-9. Records of these tests have not been 

located. However, a fracture of the seventh thoracic vertebra occurred as a result of one of two 

tests conducted with the pressure suit partly inflated [ref. 25].

Figure 3.1-8. Volunteers Undergoing an Impact Test in a Developmental Seating System 

(credit: USAF)
16

Figure 3.1-9. A Volunteer Being Tested in a Developmental Pressure Suit (credit: USAF)

Payne [ref. 26] further developed numerical models of human body dynamics and studied the 

effects of body support and restraint systems in a research effort jointly funded by the USAF and 

NASA. Important analytical results influencing the design of restraint systems included studies of 

the restraint system slackness and preload using a lumped parameter model of human response.

Focus on the Development and Validation of a Spinal Injury Model

Payne suggested that two lumped parameter models could be used to describe the human responses 

to spinal (+Z-axis) and transverse (X-axis) impact conditions. NASA and AFRL sponsored a more 

detailed study of these models [ref. 27] using data from tests with human cadavers, and impact and 

vibration tests with volunteers.

Concurrently, AFRL conducted tests with human subjects to measure human responses to spinal 

acceleration since the USAF was planning to develop an ejection seat and would update its military 

specification for ejection seats. At the same time, the USAF and USN were developing a two-man, 

emergency escape crew cockpit for the F/B-111. The spinal injury model was used on an 

experimental basis to evaluate the crew escape system performance during developmental and 

qualification ejections from a rocket-propelled sled. Brinkley estimated the ground landing to have 

a probability of injury greater than 20 percent. After the system was operational, the spinal injury 

rate was found to be 29.5 percent (23/78) by Hearon et al. [ref. 28].

Brinkley [ref. 29] and Brinkley and Shaffer [ref. 30] conducted laboratory impact studies and 

evaluation of the spinal injury rates associated with operational USAF ejection seats using the 

dynamic response index (DRI) model developed by Steck and Payne [ref. 27]. Using the 
17

operational ejection seat data, Brinkley and Shaffer validated the model using the injury 

probability from the work of Steck and Payne [ref. 27], but adjusted it to match the higher 

resistance to crew spinal injury shown in Figure 3.1-10 [ref. 30].

Figure 3.1-10. Probability of Spinal Injury Estimated from Operational Ejection Seat Experience

The DRI model estimates combined with ejection tests on volunteers were used to correct an 

extraordinarily high spinal injury rate associated with the ejection seat used in the F-4 aircraft. The 

complex curvature of the seat back and headrest caused the seat occupant to be forward of the 

ejection catapult thrust vector [refs. 30 and 31]. This caused the occupant’s head and upper body 

to flex forward during ejection, lowering the threshold of compression injury to the thoracic spine. 

To correct the high injury rate, the ejection catapult thrust was reduced and a rocket was added to 

the seat bottom to sustain the acceleration after the seat separated from the aircraft. The seat-back 

geometry and thrust vector alignment could not be accomplished without a major seat redesign or 

replacement (see Figure 3.1-11). Because of the feasible changes, the spinal injury rate was 

reduced from 34 to 8 percent, as shown in Figure 3.1-12 [ref. 30].

The ability of the spinal column to withstand +Z-axis acceleration without injury may be 

significantly increased with proper spinal column alignment prior to impact. 
18

Figure 3.1-11. Seat-Back Geometry and Spine Alignment to the Ejection Thrust Vector
19

Figure 3.1-12. The F-4 Ejection Seat Injury Rates Associated with the DRI

Evaluating the Acceptability of Multi-directional Acceleration Conditions

During the 1970s, the USAF used a method to evaluate multi-directional acceleration 

environments that incorporated the DRI combined with a method to assess the magnitude of 

accelerations in the other principle axes that occurred within a moving 62-msec analysis window. 

A computer program was developed to accomplish this analysis incorporating look-up tables to 

assess the effects of the X-, Y-, and –Z-axes. While this method provided a consistent means to 

evaluate the performance of relatively conventional ejection seats, it did not provide a means to 

evaluate a new generation of ejection seats. These new ejection seat designs incorporated an array 

of attitude, altitude, and airspeed sensors; digital flight control; ejection catapult thrust control; and 

rocket thrust vector control. In situations where the conditions at ejection were benign, the escape 

system performance would be controlled to produce a low injury risk. Where the conditions 

presented a higher risk to the seat occupant, the ejection seat would produce accelerations with a 

higher injury risk, but a higher likelihood of a successful escape. 

This concept led the USAF to generate a strategic plan for a development program to demonstrate 

the technologies required to demonstrate an advanced ejection seat that would have such flight 

control features, and be capable of safe escape at aircraft speeds to 700 knots (equivalent airspeed). 

The new technologies required to evaluate the performance of such an ejection seat included an 

ATD capable of measuring specific forces and moments at key internal locations and joints, a 

rocket sled capable of providing adverse attitudes and roll motion of an aircraft forebody at the 

time of an ejection test, and a means to sense and assess the injury risk associated with the escape 

system’s performance [ref. 32].
20

Development of a Method to Assess the Injury Risk Resulting from Multi-axial 

Accelerations

Brinkley [ref. 33] proposed a means to evaluate the performance of an escape system using a whole 

body acceleration exposure limit method. Its primary objective was to compute a set of dynamic 

responses that could be used to estimate the injury risk levels. The computations would be based 

on measurements of linear acceleration and angular velocities at a known point. Brinkley proposed 

that if the linear acceleration at a point on the seat and the angular velocity were known, then the 

seat motion would be uniquely defined and the linear acceleration at any point in the seat 

coordinate system could be calculated. Brinkley proposed that a single-degree-of-freedom (DOF), 

lumped-parameter model could be used to assess the dynamic response of the human body in each 

orthogonal axis. The dynamic responses of the three orthogonal axes could be used to calculate a 

general whole body injury risk in terms of an ellipsoidal approximation. 

Model parameters, natural frequency, and damping coefficients were selected on the basis of 

available laboratory data and existing lumped-parameter models. Injury risk levels were chosen 

on the basis of operational escape system experience and accidental injuries that had occurred in 

laboratory experiments. The DRI was adopted for the +Z-axis. A probability of spinal injury had 

been validated by ejection experience and non-injury and injuries experienced in the laboratory. 

The results of –Z-axis accelerations explored by investigators, such as Shaw [ref. 34] and 

Schulman et al. [ref. 18], and the USAF experience with downward ejection seats, were initially 

used to estimate low, moderate, and high risk levels. For simplicity, the DRI parameters were 

selected for the –Z-axis model.

Initial estimates of the X-axis models were derived from numerous reports of tests with volunteers 

by the USAF, USN, DoD contractors, and academic laboratories. The estimates supported the 

analytical work of Payne [ref. 26] and Steck and Payne [ref. 27].

Selection of model parameters for the Y-axis proved to be the most difficult due to the limited 

data. One set of sideward impact data was selected to be suitable for analysis [ref. 35]. Injury risk 

levels could not be developed with high confidence since few injuries had been observed beyond 

adverse cardiovascular responses [refs. 21, 23, 24 and 36] and syncope [ref. 36]. Unpublished 

experimental accidents with volunteer subjects conducted by aerospace contractors were added to 

estimate the high injury risk region. The influence of side support panels was also estimated. 

In subsequent publications by Brinkley [ref. 37] and Brinkley, et al. [ref. 38], the results of 

laboratory experiments were used to improve the lumped-parameter models. For example, 

Brinkley, et al. [ref. 38] reported the results of impact experiments and analyses to better define 

the parameters for the –Z-axis model. 

BDR Criteria Application

The BDR multi-axial dynamic response criteria have been used in numerous research and 

development applications. These include the:

 investigation of the Challenger accident, 

 the development, test, and evaluation of the Crew Escape Technologies escape system 

demonstration ejection seat, 

 evaluation of landing impacts of the Soyuz CM, 

 the design and assessment of an escape system concept for the National Aerospace Plane,
21

 test and evaluation of the Russian K-36D ejection seat performance [ref. 39],

 development, test, and evaluation of the X-38 assured crew recovery system, 

 development, test, and evaluation of the Soyuz TMA (Transport Modified 

Anthropometric) CM, and 

 advanced development, test, and evaluation of the K-36D-3.5A ejection seat [ref. 40].

Impact Exposure Limits for NASA Standard 3001

In December 2004, Dr. Smith Johnston convened a panel of sustained and impact acceleration 

experts to review a draft NASA standard for CxP. The BDR model criteria were recommended to 

be incorporated in NASA Standard 3001 (NASA-STD-3001), NASA Space Flight Human System 

Standard. A draft was prepared, reviewed, and incorporated. Rotational limits were proposed and 

recommended.

Assumptions and Requirements for the BDR Model Criteria

 Accelerations of less than 500 msec (e.g., during liftoff, launch abort, landing impacts, 

and parachute deployments).

 Seated crew where seat padding or cushions preclude amplification of transient linear 

accelerations transmitted to the occupant.

 Crew contained by a restraint system that included pelvic, torso, and anti-submarining 

restraints that provide occupant confinement similar to a conventional five-point harness.

 During the experimental efforts, the restraint system was adequately pre-tensioned to 

eliminate slack. Pyrotechnically powered inertial reels were used to position escape 

system occupants and to eliminate slack in the restraint during the operation ejection 

cases that were used.

 The +Z-axis limits assume that the seat cushion materials do not amplify the acceleration 

transmitted to the seat occupant. 

 The +X-axis limits assume that the seat occupant’s head is protected by a flight helmet 

with a liner adequate to pass the test requirements of American National Standards 

Institute (ANSI) Z-90 (latest edition) or equivalent. 

 The crew will be similarly restrained during all events that might require application of 

the BDR model criteria.

 The model cannot predict injury caused by rigid suit elements (i.e., localized blunt 

trauma, localized point loading, and interference with restraints). 

 The testing used to develop the lateral dynamic response had minimal gap between the 

subject and the seating support surfaces. Additional space between the seat and subject 

will increase the injury risk, but this increased risk will not be reflected in the model 

predictions.

Recent Additions to the Supporting Database

Since the original BDR criteria publications, research has been continued by the AFRL Human 

Effectiveness Directorate to address specific issues associated with the criteria use. These issues 

included expansion of the empirical database to study the influence of variations in impact 

duration, impact magnitude, off-axis impacts, and the influence of a specific complex waveform 

on the dynamic response properties of volunteer subjects. Topics of interest have included linearity 

of human dynamic response properties, such as acceleration and externally measured forces. 
22

Studies of the effects of helmet mass properties on the dynamic response of the head and neck 

have been of considerable interest to the USAF in view of the operational introduction of helmetmounted displays and night-vision equipment. Now that women are flying fighter aircraft equipped 

with ejection seats, more emphasis has been given to the study of the dynamic response of both 

genders to the accelerations and forces generated during emergency escape.

Impact studies have been conducted with military volunteer subjects to evaluate impact profiles 

with acceleration levels of 3, 6.9, and 12 G and impact rise times varying from 30–130 msec and 

to study the amplification or attenuation of subject accelerations and forces reacted by the seat 

structure. The axes studied to date include the +Z, Z, and X-Y and X-Z vectors, and vector 

combinations. More than 800 impact exposures have been accomplished with volunteers to explore 

their responses to these conditions. The +Z-axis study results have been analyzed and demonstrate 

that the measured seat pan force is linear over the range of acceleration levels explored. The impact 

force amplification or attenuation was found to be consistent with a simple, single-DOF, lumpedparameter model.

Strzelecki [ref. 41] experimentally studied the dynamic inertial response of volunteer subjects to

determine if the simple lumped-parameter model used within the BDR model could adequately 

simulate the forces measured at the seat pan. The profile studied was unique in that it consists of a 

short-duration impact followed by longer duration acceleration. The study was accomplished using 

the AFRL vertical deceleration tower. The facility was modified to produce an impulsive 

deceleration prior to the impact carriage deceleration-metering pin striking the water-filled 

decelerator normally used to control the deceleration magnitude, rise time, and duration. The 

measured responses of the volunteer subjects for seat pan force and chest and head acceleration 

confirmed the DRI model predictions and that the dynamic preload impulse was effective in 

reducing the dynamic overshoot and thereby the injury potential associated with the following 

deceleration pulse.

The study of gender influence included reviews of the literature and experimental investigations. 

Allnutt [ref. 42] conducted a review of the medical and safety literature exploring differences 

between the responses of men and women to traumatic vehicular events. Allnutt provided ample 

evidence that the gender-specific biological and anatomical characteristics affect a body’s 

likelihood of trauma. These characteristics include differences in stature, weight, tissue strength, 

and endocrine humoral factors. For example, bone from a female is different in density, structure, 

and strength, and changes with age from that of a male. Allnutt provided evidence showing that 

women’s bones are smaller when comparing them to men of the same height and weight. The 

difference is not limited to the size of boney structures; the microscopic thickness of trabecular 

(spongy) and cortical (compact) bone is thinner in women and becomes thinner with age. These 

findings demonstrate the use of the simple scaling methods employed in the automotive industry 

to adjust for the sensitivity of women to impact is a simplification of the more complex relationship 

between male and females and bone strength. The findings highlight BDR criteria limitations since 

the majority of the data used to develop the exposure limits were based upon laboratory studies 

primarily using male volunteers. 

With women now flying fighter aircraft, AFRL has undertaken literature reviews and experimental 

and analytical studies to investigate the influence of gender in sustained and impact acceleration 

conditions. Morris and Popper [ref. 43] studied the influence of gender and impact on neck motion. 

Buhrman and Wilson [ref. 44] conducted an initial study of gender and subject size on factors that 

might increase spinal injury risk during emergency escape from aircraft. The introduction of 
23

helmet-mounted displays and night-vision systems has increased the USAF’s concern for cervical 

spine injuries. Perry [ref. 45] conducted tests using a vertical deceleration facility along with male 

and female volunteers to study the influence of variations in the inertial helmet properties on the 

subjects’ dynamic responses to +Z-axis impacts. In 2000, Perry conducted experiments to compare 

the responses of males and females to –Gx impacts [ref. 46]. More recently, Gallagher et al. [ref. 

47] compared cervical stress and bone mineral density to subject size and gender during +Z-axis 

impacts conducted using the AFRL vertical deceleration tower. Gallagher et al. [ref. 48] conducted 

analyses of vertebral stress and bone mineral density during +Z-axis impacts.

In general, there was little difference between the physical responses of males and females. The 

natural frequencies that were determined revealed no statistically significant differences. The 

damping coefficient ratios were somewhat smaller. Computed stresses for +Gz studies showed 

that females experienced 15-percent greater cervical stress than males during 10-G impacts 

[ref. 47]. When normalized for weight, lumbar stress was 15 percent higher in the female subjects 

than in male subjects [ref. 49].

AFRL has conducted numerous experiments to evaluate the influence of factors, such as seat 

geometry, restraint system design features (including attachment position), and seat cushion 

properties on the likelihood of injury using vertical and horizontal impact facilities. Examples 

include: Perry et al. [ref. 50], a study of seat back angle effect on impact response; Hearon and 

Brinkley [ref. 51], a study of the influence of a negative-G strap; Hearon et al. [ref. 52], a study of 

factors influencing a high spinal injury rate during CM landings; and Hearon and Brinkley [ref. 

53], a study of seat cushion effects on subject impact responses.

Recommended Topics for Continued Research

The influence of gender has been studied by tests of human tissues, analysis of automotive 

accidents, and impact tests with volunteers. Additional research is required to determine tolerance 

decrements for body segments and whole body impact exposure limits. Available data should be 

adjusted for the accident victim’s age, or post-mortem human subject (PMHS).

Research on the effects of microgravity exposure on crew should be continued. Emphasis should 

be given to areas of the body that carry the highest impact loads during landing impact (e.g., hip, 

shoulder, and spinal column).

The influence of the distance between the load-carrying areas of the human body and the support 

structures should be evaluated analytically and experimentally.

The influence of pressure suits and their design features cannot be evaluated until specific 

developmental designs are defined. The topic of focused blunt trauma due to impact with 

mechanical design features of pressure suits remains relatively unexplored. The use of PMHSs 

remains a viable option for exploring this problem area.

Biodynamic Injury Criteria

For many years, the automotive industry has been interested in developing injury assessment 

methods to help design safer automobiles. Their goal is to relate the injury risk in a specific body 

region to measurements made in that corresponding region in an ATD or a computer-modeled 

ATD. Injury biomechanics research focuses on determining the appropriate injury mechanism 

parameters (e.g., acceleration, force, or structural deformation of body components) that can be 

measured in an ATD and determine limit values for those parameters for a certain level of injury 

risk in that body region. Accordingly, the design features of the ATDs and their limitations in 
24

representing the human body and its organs play a role in how injury assessments are performed. 

These limit values for injury criteria are referred to as injury assessment reference values (IARV) 

in the automotive industry. The ATD most commonly used in current automotive safety evaluation 

is the Hybrid III midsize adult male ATD. 

Mertz [ref. 54] provides a detailed injury risk assessments description based on ATD responses. 

Mertz describes the process as:

“Injury assessment reference values (IARVs) were developed to assess the 

efficacy of General Motors (GM) restraint system designs under a variety of 

simulated frontal accident conditions using the Hybrid III midsize adult male 

dummy as the vehicle occupant. These design limit values were chosen such that 

if an IARV was not exceeded in the prescribed test, then the risk of the associated 

injury would be unlikely. “Unlikely” was defined as risk levels less than 

5 percent. Since injury risk curves for the various dummy responses did not exist, 

the limit values that were chosen for the IARVs were simply conservative 

estimates of response values that would be consistent with the unlikely definition. 

These IARVs were published in 1984 as part of the GM petition of the NHTSA to 

allow the use of the Hybrid III midsize adult male dummy in Federal Motor 

Vehicle Standard (FMVSS) 208 testing. In 1993, IARVs were published for 

response measurements of the Hybrid III small female and large adult male 

dummies [refs. 55 and 56]. These values were obtained by applying constant 

failure stress scaling to the Hybrid III midsize adult male IARVs taking into 

account size differences.

“In 1998, the American Automobile Manufacturers Association, which is now the 

Alliance of Automobile Manufacturers (Alliance) proposed to NHTSA that the 

FMVSS 208 certification limits should be set at the 5-percent injury risk level. 

This would put all FMVSS 208 limits on an equal injury risk level that is 

consistent with the “injury unlikely” intent of the IARVs while allowing for 

reasonable compliance margin. The injury risk curves used by the Alliance were 

based on those proposed by Mertz et al. [ref. 57] for responses of the Hybrid III 

family of dummies.” 

The tables for injury criteria limit values for crew protection (see Section 3.1.2.1) are based on the 

FMVSS 208 IARVs for the small female, large male, and midsize male Hybrid III ATDs. The 

tables list the corresponding values for the 2- and 0.5-percent risk levels taken from the IARV risk 

curves presented in Mertz [ref. 54]. Deconditioned values for these parameters, based on the 

procedures given in Section 3.1.2.6, are listed below.

Head Injury

3.1.2.1.1 Head Injury Criteria 15 msec

Assessing the potential for head injury (e.g., brain and skull) has been a primary focus in injury 

biomechanics research for over 40 years. The pioneering efforts by researchers to predict brain 

injury from head acceleration measurements led to the Wayne State Tolerance Curve (WSTC) [ref. 

58]. Using the observation that most emergency room patients admitted with simple linear skull 

fractures were concussed (although not all concussed patients had skull fractures), they 

hypothesized that determining the translational head accelerations associated with the initiation of 
25

linear skull fractures in drop tests of embalmed cadaver heads would give an estimate of 

concussion occurrence. The extreme association of the limit accelerations with the durations of 

impact found in those tests for contact durations below 4 msec led Gadd [ref. 59] to plot the data 

on a logarithmic scale, which exhibited a linear relationship with a slope of –2.5 between the peak 

acceleration and time duration on a log-log plot. Gadd proposed the Gadd Severity Index (GSI), a 

weighted impulse criteria in which the integral of the measured head acceleration over time, raised 

to the 2.5 power was a measure of head injury potential based on the WSTC. Because the time 

interval for calculating the GSI was not specified, this led to confusion over the interpretation of 

more complex head acceleration waveforms. Versace [ref. 60] proposed an optimization method 

to search for the acceleration time interval, (t2-t1), which would maximize the GSI. This is referred 

to as the Head Injury Criteria (HIC) and codified by the NHTSA. All passenger cars sold in the 

U.S. must meet a designated level of HIC in a 35-mph barrier crash test for the vehicle occupants.

Equation 3.1-1. HIC Formula

HIC(t) = [

1

t2−t1

∫ ardt t2

t1

]

2.5

(t2 − t1)

While there have been many efforts to improve the HIC predictive features, there have been no 

methods that appear to have advantages over the HIC in routine crash protective system testing. 

The only change in the automotive standards has been a lowering of the HIC limit from 1,000 to 

700 for the midsize male. Since the HIC is felt to be most valid for head accelerations due to 

contact with a generally rigid surface, researchers [ref. 61] have selected a maximum HIC interval 

(t2-t1) of 15 msec or less based on biomechanical head impact research. Table 3.1-1 shows the 

original values included in the HSIR; and Table 3.1-2 lists the revised limit values for HIC for the 

small female, midsize male, and large male dummies at the 0.5- and 5-percent risk levels given in 

Mertz [ref. 54]. Crew deconditioning is not considered applicable to head injury assessment.

Table 3.1-1. Original HIC 15 msec (HIC15) HSIR Requirements

0.5 percent 2 percent 5 percent

Small 

Female

50th-

percentile 

Male

Large 

Male

Small 

Female

50th-

percentile 

Male

Large 

Male

Small 

Female

50th-

percentile 

Male

Large 

Male

HIC15 300 500 700

Table 3.1-2. Proposed Revisions to HIC15 HSIR Requirements

0.5 percent 5 percent

Small 

Female

50th-

percentile 

Male

Large 

Male

Small 

Female

50th-

percentile 

Male

Large 

Male

HIC15 278 250 239 779 700 670

3.1.2.1.2 Peak Head Acceleration

Mertz [ref. 54] presented risk curves for peak head acceleration without reference to pulse duration. 

The 5-percent risk limit was established as 180 G peak resultant head acceleration for the 50th-

percentile (or midsize) male. Table 3.1-3 shows the original values included in the HSIR and Table 

3.1-4 lists the revised limit values for the peak resultant head acceleration values for 0.5- and 5-

percent risk level values for the three ATD sizes. Crew deconditioning is not considered applicable 

to head injury assessment.
26

Recent analyses of racing car accidents using onboard crash recording and a mathematical driver 

model have indicated that peak head and head rotational accelerations may be more predictive of 

head injury than the HIC [ref. 62]. A number of extensive head finite element models (FEMs) have 

been developed as research tools for studying head injury assessment methods. Most are still in 

the validation stage of development and are complex and time consuming to run. An exception is 

the Simulated Injury Monitor (SIMon) model developed by the NHTSA. It is a simplified brain 

model in which measurements of 3-dimensional (D) motion in the form of head translational 

accelerations, and rotational velocities and accelerations, measured in an ATD test, can be used to 

assess the risk of several forms of brain injury. This model is in the development phase of its 

implementation.

Table 3.1-3. Peak Head Acceleration HSIR Requirements

0.5 percent 2 percent 5 percent

Small 

Female

50th-

percentile 

Male

Large 

Male

Small 

Female

50th-

percentile 

Male

Large 

Male

Small 

Female

50th-

percentile 

Male

Large 

Male

Peak Head 

Acceleration (G) 

(deconditioned)

119 112 109 151 142 138 166 155 151

Peak Head 

Acceleration (G) (nondeconditioned)

138 130 127 176 165 160 193 180 175

Table 3.1-4. Proposed Revisions to Peak Head Acceleration HSIR Requirements

0.5 percent 5 percent

Small 

Female

50th-

percentile 

Male

Large 

Male

Small 

Female

50th-

percentile 

Male

Large 

Male

Peak Head Acceleration (G) 138 130 127 193 180 175

Neck Injury Criteria

3.1.2.2.1 Peak Neck (Cervical) Bending Moments

The concept of predicting neck injuries due to bending has been studied for almost as long as head 

injury prediction. The present bending moment limit criteria for fore and aft bending 

(i.e., flexion and extension) are based on the work of Mertz and Patrick [ref. 63] in which the necks 

of volunteers and cadavers were inertially loaded by head motion in sled tests where the subjects 

were restrained by belt restraints or seatbacks without head support. The concept of neck injury 

due to bending is based on the concern for excessive bending motion (in flexion, extension, or 

lateral bending) causing ligamentous and boney damage as the vertebrae impinge in extreme 

motion. Since neck motion is difficult to measure with transducers in an ATD, the bendingmoment levels associated with excessive motion in cadaver tests was used by Mertz and Patrick 

as an indicator of potential impingement. Mertz [ref. 54] does not give risk curves for neck 

compressive loading or neck shear loading. Lacking that information, a conservative estimate of 

low risk (~0.5 percent) can be made by taking the data on human volunteer static neck loading 

from Patrick and Chou [ref. 64] as a surrogate for low-risk dynamic loading in neck flexion and 

lateral bending. Mertz [ref. 54] presents a risk curve for extension moment values related to serious 

injury for the case of minimum muscle tone in various sized dummies. 
27

Table 3.1-5 shows the original values included in the HSIR. Table 3.1-6 lists the revised limit 

values for the neck-bending moment limit values (i.e., 0.5- and 5-percent risk for the three ATD 

sizes for extension and 0.5- and 5-percent for flexion and lateral bending) for non-deconditioned 

and deconditioned crew. See Section 3.1.2.6 for information on the deconditioning factor.

Table 3.1-5. Peak Neck Bending Moment HSIR Requirements

0.5 percent 2 percent 5 percent

Small 

Female

50th-

percentile 

Male

Large 

Male

Small 

Female

50th-

percentile 

Male

Large 

Male

Small 

Female

50th-

percentile 

Male

Large 

Male

Peak neck (cervical) flexion 

bending moment (Newton 

meter (Nm)) (deconditioned) 

42 83 83 57 108 125 89 163 222

Peak neck (cervical) flexion 

bending moment (Nm) (nondeconditioned)

49 96 96 66 126 145 104 190 258

Peak neck (cervical) lateral 

bending moment (Nm) 

(deconditioned)

33 65 65 41 82 82 62 123 123

Peak neck (cervical) lateral 

bending moment (Nm) (nondeconditioned)38 75 75 48 95 95 72 143 143

Peak Neck (cervical) extension 

bending moment (Nm) 

(deconditioned)

15 34 42 27 49 67 28 56 75

Peak Neck (cervical) extension 

bending moment (Nm) (nondeconditioned

17 39 49 31 57 78 33 65 87

Table 3.1-6. Proposed Revisions to Peak Neck Bending Moment HSIR Requirements

0.5 percent 5 percent

Small 

Female

50th-

percentile 

Male

Large 

Male

Small 

Female

50th-

percentile 

Male

Large 

Male

Peak neck (cervical) flexion bending 

moment (Nm) (deconditioned) 22 43 57 82 163 215

Peak neck (cervical) flexion bending 

moment (Nm) (non-deconditioned) 25 50 66 95 190 250

Peak neck (cervical) lateral bending 

moment (Nm) (deconditioned) 27 53 71 62 123 162

Peak neck (cervical) lateral bending 

moment (Nm) (non-deconditioned) 31 62 82 72 143 188

Peak Neck (cervical) extension 

bending moment (Nm) 

(deconditioned)

23 46 61 34 66 88

Peak Neck (cervical) extension 

bending moment (Nm) (nondeconditioned27 54 71 39 77 102
28

3.1.2.2.2 Peak Neck (Cervical) Axial Tension

In analyzing the biomechanical database for neck tension, the Alliance noted there was one fatal 

neck lesion at the 5-percent risk level. As a result, the 3-percent risk level for neck tension was 

chosen for implementation in FMVSS 208 [ref. 54]. The worst case for neck tension loading is 

considered to be when the neck muscles are tensed only to the degree required to keep the head 

upright. Mertz [ref. 54] shows neck tension risk curves for serious neck injury limits with minimum 

muscle tone. Table 3.1-7 shows the original values included in the HSIR. Table 3.1-8 lists the 

revised limit values for the neck tension limit for the three ATD sizes with non-deconditioned and 

deconditioned values as described in Section 3.1.2.6. 

Table 3.1-7. Peak Neck Axial Tension HSIR Requirements

0.5 percent 2 percent 5 percent

Small 

Female

50th-

percentile 

Male

Large 

Male

Small 

Female

50th-

percentile 

Male

Large 

Male

Small 

Female

50th-

percentile 

Male

Large 

Male

Peak neck (cervical 

spine) axial tension 

(Newton (N)) 

(deconditioned)

631 943 1,138 1,753 2,781 3,363 2,161 3,440 4,326

Peak neck (cervical 

spine) axial tension (N) 

(non-deconditioned)

734 1,097 1,323 2,038 3,234 3,910 2,513 4,000 5,030

Table 3.1-8. Proposed Revisions to Peak Neck Axial Tension HSIR Requirements

0.5 percent 5 percent

Small 

Female

50th-

percentile 

Male

Large 

Male

Small 

Female

50th-

percentile 

Male

Large 

Male

Peak neck (cervical spine) axial 

tension (N) (deconditioned) 1,635 2,595 3,137 1,780 2,830 3,970

Peak neck (cervical spine) axial 

tension (N) (non-deconditioned) 1,901 3,017 3,647 2,070 3,290 3,970

3.1.2.2.3 Peak Neck (Cervical) Axial Compression

Neck compressive injury assessment focused on a number of possible injury modes. The most 

serious injuries occur when the spinal disruption compromises the spinal cord. These injuries are 

generally irreversible and can lead to weakness, paralysis, or death. Spinal structural element 

failure (i.e., vertebrae, discs, and ligaments) can lead to disruption of the cord from bone fragments 

impinging on the cord, or severe misalignment of the vertebrae causing cord disruption. There are 

many cervical spine compression failure modes. These modes depend on the spine configuration 

and the head attitude relative to the spine and torso when the load is applied to the head [ref. 65]. 

The normal cervical spine configuration in an upright, seated subject with the eyes looking forward 

is a concave to the rear curve (i.e., lordotic curve). As the head is rotated forward (flexion), the 

cervical spine lordotic curvature decreases and disappears as the spine becomes straight with 

forward bending. Continued forward bending of the head produces a concave to the front spinal 

curve (i.e., kyphotic curve). Research has shown that when the head is rotated (flexed) forward, 

the cervical spine becomes relatively straight with a slight kyphotic curvature, and a compressive 

load is applied to the top of the head, the spine is at risk of buckling with a subsequent dislocation 
29

of adjacent vertebrae near the center of the cervical spine. Termed a fracture-dislocation, this injury 

can produce paralysis at a relatively low compressive force on the head. This injury is commonly 

seen in headfirst diving situations (e.g., water sports, automobile rollovers, and football spearing 

contacts). When the cervical spine is greatly curved in the forward or rearward direction, and the 

top of the head is loaded in compression, other failure modes, such as stretched ligaments, chipped 

vertebra, or crushed vertebra usually occur at sufficiently high forces. When the spine is straight 

and a force is applied rapidly to the top of the head, it is possible to produce a basal skull fracture 

instead of a cervical spine injury.

The compressive force limits chosen for the HSIR have been derived from the automotive injury 

biomechanics and injury assessment literature and as codified in the NHTSA FMVSS 208, 

Occupant Crash Protection (see Figure 3.1-13). FMVSS 208 specifies the maximum allowable 

compression force on the neck to be 4000 N for the 50th-percentile male Hybrid III ATD as it 

interacts with the interior of a vehicle in a 35-mph frontal barrier crash. The injury risk associated 

with the FMVSS 208 injury assessment numbers is considered to be at the 5-percent risk level. 

The HSIR lists the compressive force limits for the 5th-percentile female and the 

95th-percentile Hybrid III ATDs. The HSIR document is concerned with injury assessment values 

for risk levels lower than 5 percent. However, Mertz [ref. 54] does not give risk curves for neck 

compressive loading or neck shear loading. Lacking that information, a conservative estimate of 

low risk (approximately 0.5 percent) can be made by taking the data on human volunteer static 

loading of the neck from Mertz [ref. 66] as a surrogate for low-risk dynamic loading in 

compression and shear. These values and deconditioned crew values (as described in Section 

3.1.2.6) are given in Table 3.1-9.

Table 3.1-9. Peak Neck Axial Compression HSIR Requirements

0.5 percent 2 percent 5 percent

Small 

Female

50th-

percentile 

Male

Large 

Male

Small 

Female

50th-

percentile 

Male

Large 

Male

Small 

Female

50th-

percentile 

Male

Large 

Male

Peak neck (cervical spine) 

compression (N) 

(deconditioned)

596 946 1,142 1,067 1,694 2,046 2,167 3,440 4,154

Peak neck (cervical spine) 

compression (N) (nondeconditioned)693 1,100 1,328 1,241 1,970 2,379 2,520 4,000 4,830
30

Figure 3.1-13. FMVSS 208 Neck Compression Limit for Severe Injury

Belt-restrained automobile occupants usually do not sustain loading to the top of the head in frontal 

crashes with the exception of some possible limited contact with the airbags. However, when a 

belt-restrained crew is in a spacesuit with an attached helmet, there is a possibility of compressive 

contact of the top of the head against the helmet interior during crash loading. A similar situation 

can occur in the development of head/neck restraints for racecar drivers. Some concepts use tethers 

that run vertically, rather than the horizontal tethers seen in the HANS® device. Vertical tethers 

can induce large compressive neck forces as the body moves forward into the restraint belts and 

the head interacts with the restrained helmet. To deal with this issue, the SFI 38.1 Standard [ref. 

67] for head/neck restraints limits the maximum compressive load to 2500 N in a 70-G frontal 

crash test with a midsize male Hybrid III ATD. Design countermeasures to minimize contact of 

the top of the head with the interior of the space helmet are necessary to ensure that neck 

compressive forces are controlled to acceptable levels as defined by HSIR Table N4-2 Transient 

Force Application Limits (see Table 3.1-9). Revisions to the table (i.e., removal of the 2-percent 

limits) are shown in Table 3.1-10. 

Table 3.1-10. Proposed Revisions to Peak Neck Axial Compression HSIR Requirements

0.5 percent 5 percent

Small 

Female

50th-

percentile 

Male

Large 

Male

Small 

Female

50th-

percentile 

Male

Large 

Male

Peak neck (cervical spine) 

compression (N) (deconditioned) 596 946 1,142 2,167 3,440 4,154

Peak neck (cervical spine) 

compression (N) (nondeconditioned)

693 1,100 1,328 2,520 4,000 4,830

3.1.2.2.4 Peak Neck (Cervical) Shear Force

Mertz [ref. 54] does not give risk curves for neck compressive loading or neck shear loading. 

Lacking that information, a conservative estimate of low risk (~0.5 percent) can be made by taking 

the data on human volunteer static neck loading from Mertz [ref. 66] as a surrogate for low-risk 

dynamic loading in compression and shear. These values and deconditioned crew values are 

0 10 20 30 40 50 60 0

500

1000

1500

2000

2500

3000

3500

4000

4500

5000

Duration [ms]

Neck Compression [N]
31

described in Section N4.1 [ref. 1] and are given in Table N4-2 [ref. 1] (see Table 3.1-11). Revisions 

to the Table (i.e., removal of the 2-percent limits) are shown in Table 3.1-12.

Table 3.1-11. Peak Neck Shear Force HSIR Requirements

0.5 percent 2 percent 5 percent

Small 

Female

50th-

percentile 

Male

Large 

Male

Small 

Female

50th-

percentile 

Male

Large 

Male

Small 

Female

50th-

percentile 

Male

Large 

Male

Peak neck (cervical spine) shear 

force (N) (deconditioned) 593 946 1,142 919 1,462 1,766 1,680 2,666 3,219

Peak neck (cervical spine) shear 

force (N) (non-deconditioned) 690 1,100 1,328 1,069 1,700 2,053 1,953 3,100 3,743

Table 3.1-12. Proposed Revisions to Peak Neck Shear Force HSIR Requirements

0.5 percent 5 percent

Small 

Female

50th-

percentile 

Male

Large 

Male

Small 

Female

50th-

percentile 

Male

Large 

Male

Peak neck (cervical spine) shear 

force (N) (deconditioned) 593 946 1,142 1,680 2,666 3,219

Peak neck (cervical spine) shear 

force (N) (non-deconditioned) 690 1,100 1,328 1,953 3,100 3,743

Lower Extremity Injury Criteria

Mertz [ref. 54] does not give risk curves for axial compressive loading for the lower extremities. 

Lacking that information, a conservative estimate of low risk (~0.5 percent) can be made by taking 

the data on knee long-duration or static-loading from Mertz [ref. 66] as a surrogate for low-risk 

compressive femur force limit and the peak knee force limit for short duration loading as the 5-

percent risk femur compressive force limit. 

Tibial plateau compressive force limits are given in Mertz [ref. 66] for the medial and lateral 

plateaus. A low-risk estimate of tibial axial compression can be estimated from assuming the single 

plateau limit force acting on the complete tibia and a 5-percent risk tibial axial compressive force 

limit as plateaus being simultaneously loaded, doubling the limiting force over the low-risk values. 

These values and deconditioned crew values are described in Section N4.1 [ref. 1] and are given 

in Table N4-2 [ref. 1] (see Table 3.1-13) Revisions to the table (i.e., removal of the 2-percent 

limits) are shown in Table 3.1-14.

Table 3.1-13. Peak Lower Extremity Axial Compression HSIR Requirements

0.5 percent 2 percent 5 percent

Small 

Female

50th-

percentile 

Male

Large 

Male

Small 

Female

50th-

percentile 

Male

Large 

Male

Small 

Female

50th-

percentile 

Male

Large 

Male

Peak femur axial compression (N) 

(deconditioned) 1,914 3,000 3,690 2,498 3,801 5,013 3,862 5,670 8,100

Peak femur axial compression (N) 

(non-deconditioned) 2,552 4,000 4,920 3,331 5,068 6,684 5,150 7,560 10,800

Peak tibia axial compression (N) 

(deconditioned) 1,914 3,000 3,690 2,490 3,900 4,800 3,825 6,000 7,380

Peak tibia axial compression (N) 

(non-deconditioned) 2,552 4,000 4,920 3,320 5,200 6,400 5,100 8,000 9,840
32

Table 3.1-14. Proposed Revisions to Peak Lower Extremity Axial Compression HSIR Requirements

0.5 percent 5 percent

Small 

Female

50th-

percentile 

Male

Large 

Male

Small 

Female

50th-

percentile 

Male

Large 

Male

Peak femur axial compression (N) (deconditioned) 3,867 5,670 7,212 4,640 6,803 8,653

Peak femur axial compression (N) (nondeconditioned) 5,156 7,560 9,616 6,186 9,070 11,53

7

Peak tibia axial compression (N) (deconditioned) 1,914 3,000 3,690 3,825 6,000 7,380

Peak tibia axial compression (N) (nondeconditioned) 2,552 4,000 4,920 5,100 8,000 9,840

Restrained Body Movement and Deflection Injury Criteria

3.1.2.4.1 Chest Sternal-To-Spine Deflection

The Hybrid III ATD chest deflection has been shown to be a good indicator of chest injury 

potential in belt-restrained vehicle occupants [ref. 68]. When two parallel shoulder belts are used 

to restrain the upper torso, the loading from the belts should be considered distributed loading. 

Mertz [ref. 54] presents risk curves for normalized deflection. The HSIR 0.5-, 2-, and 5-percent 

risk limit values of sternal deflection for the three ATD sizes are given in Table 3.1-15 (not 

distributed load), with the proposed revisions shown in Table 3.1-16 (revised to reflect two 

shoulder belts instead of one).

Table 3.1-15. Chest Sternal-To-Spine Deflection HSIR Requirements

0.5 percent 2 percent 5 percent

Small 

Female

50th-

percentile 

Male

Large 

Male

Small 

Female

50th-

percentile 

Male

Large 

Male

Small 

Female

50th-

percentile 

Male

Large 

Male

Chest Sternal to 

Spine Deflection 

(millimeter (mm))

28 31 35 36 44 49 41 50 55

Table 3.1-16. Proposed Revisions to Chest Sternal-To-Spine Deflection HSIR Requirements

0.5 percent 5 percent

Small 

Female

50th-

percentile 

Male

Large 

Male

Small 

Female

50th-

percentile 

Male

Large 

Male

Chest Sternal to Spine 

Deflection (mm) 33 41 45 39 48 53

Body Movement Limits

Dynamic overshoot occurs when the acceleration, forces, and moments measured on or within 

the human body or body segments exceed the input acceleration or forces that are imparted to the 

body. Dynamic overshoot is adverse only when the human body responses exceed the specified 

injury limits. The body movement limits are intended to control the amount of free space that 

will be permitted for body motion until the accelerating seat and restraint system contacts the 

occupant. Minimization of these movements reduces dynamic overshoot by coupling the 

occupant to the seat and restraint system. The indicators of acceptability are the Hybrid III ATD 

response parameters. Displacement measurements may be helpful in understanding the 
33

relationship between free space and injury metrics, but the influence of free space on the 

potential for injury is a function of additional factors (e.g., impact acceleration, velocity change, 

and viscoelastic properties) and the injury tolerance of the impacted body segments. The 

restrained body movement limits are given in 

Table 3.1-17 for the various directions of loading. 

Table 3.1-17. Restrained Body Movement HSIR Limits

Lateral 

(±Gy)

Anterior 

(+Gx)

Posterior 

(-Gx)

Head Movement (mm) 75 125 25

Chest Movement (mm) N/A 63 25

Pelvic Movement (mm) 37 50 25

Shoulder Movement (mm) 50 N/A N/A

Caudal Pelvic Movement (+Gz) (mm) 50

Upward Head Movement (-Gz) (mm) 50

Derivation of Deconditioning Factor

Nominal landing limits are set for critical parameters (e.g., loads and deflections) to keep 

significant injury risk at the 0.5-percent risk level, with additional tolerance reductions resultant 

from the crew being deconditioned due to long-duration spaceflight and increased body movement 

presumed due to the launch/entry/landing suit. Therefore, a scaling factor has been applied to the 

non-launch abort landing scenarios in the form of a deconditioning coefficient that adjusts for the 

reduced crew capacity to endure flight/landing loads. It is assumed that the crew in all launch abort 

landing scenarios will not be deconditioned. The added body deflections due to wearing the flight 

suit cannot be estimated at this time. However, the body deflection constraints defined in CxP 

70024 HS3130 will apply to the seat and the suit system. Therefore, a separate suit deflection 

coefficient is not required [ref. 1].

A deconditioning factor was estimated to adjust the maximum allowable loading at the femoral 

neck and the lumbar spine during a severe Orion reentry landing. The maximum allowable skeletal 

loading that can be experienced by an able-bodied person (i.e., no deconditioning) at the femoral 

neck, tibia, and lumbar spine during an Orion launch abort landing has been estimated from 

terrestrial crash impact data and other sources. This deconditioning factor is a function of measured 

human body physiological changes associated with dwell time away from the Earth’s surface (i.e., 

reduced gravitational environment). The data for the deconditioned crew were derived from bone 

mineral density (BMD) measurements by dual-energy X-ray absorbance and quantitative 

computerized tomography, which can provide volumetric BMD changes in the trabecular bone 

compartment over a mission.

The deconditioning factor can be multiplied by the able-bodied loading estimates to account for 

the BMD loss. For purposes of this analysis, the deconditioning factor was assumed to be a 

proportionality factor relating the allowable pre- and post-flight skeletal loading after 

deconditioned BMD loss. It was further assumed that the same probability of injury should exist 

in pre- and post-flight cases. It should be noted that the highest likelihood contingency landing 

scenarios were drivers for establishing the following loading conditions: spine axial compressive 

loading in a seated crew; axial tibial compressive forces occurring during a seat stroke or crush in 

which the foot pan for a seated crew is pushed in the cephalic direction relative to the trunk; hip 
34

loading due to a pure lateral blow to the greater trochanter of a seated crew; and hip loading due 

to a blow to the kneecap of a seated crew. 

Therefore, the following deconditioning factors, mainly driven by spaceflight-induced 

musculoskeletal changes, have been applied to the values in Appendix N of CxP 70024 [ref. 1], 

Head Acceleration Limits, Neck Protection Criteria, and Transient Force Applications Limits as 

noted:

Φ - deconditioned crew coefficient for femur and tibia = 0.75

ξ - deconditioned crew coefficient for spinal elements and head = 0.86

Application of Occupant Protection Principles

The following are examples of industry application of restraint principles:

 The body may be restrained to maintain accelerations and loads to within the 

allowable limits.

 The torso may be restrained with a multiple attach point harness.

 The body may be supported with a conformal seat that is essentially rigid and fits the 

contours of the back and bottom of the torso, shoulders, pelvis, and legs.

 The sides of the head, shoulders, hips, and legs may be supported to reduce sideimpact movement with close-fitting, conformal surfaces if required to maintain loads 

within the limits of human tolerance.

 The head and neck are supported such that loads and accelerations are within the 

limits of human tolerance.

Effective torso restraint in automobile racing (e.g., NASCAR and Indy Racing League (IRL)) and 

in space vehicle landing (e.g., Soyuz TM and TMA) has been accomplished with seats that 

conform to the crew’s back and sides to provide continuous support for the pelvis, lumbar and 

thoracic spine, and helmeted head. Such a seat conforms to the shape of the back and is sufficiently 

stiff to support the back shape and spine during impact accelerations (i.e., the back of this seat is 

conformal to the body). 

Effective car racing seats provide lateral support for sides of the pelvis, shoulders, and head. In the 

past, racing seats have commonly provided close-fitting pelvis lateral support, but such seats 

provided no motion control for the upper torso and head. The addition of panels that restrain the 

shoulders and head dramatically reduces injury potential in side impacts from automobile crashes.

3.2 Biodynamic Modeling

Coordinate System Definitions

This section describes the various coordinate systems employed. Appropriate data transforms were 

applied in all instances to ensure that the correct input and output orientations and polarities were 

obtained. All data were converted to Society of Automotive Engineering (SAE) J211/1 coordinate 

conventions to allow data comparisons [ref. 69]. The purpose of this section is to show the 

coordinate systems used and to provide the transformation matrices used to convert: input data 

from the environment specific coordinates to the SAE coordinate system; input data from SAE 

coordinates to the model specific coordinate system; and the model responses from the model 

specific coordinates back to the SAE coordinate system.
35

Environment Specific Coordinate Systems

3.2.1.1.1 WPAFB Horizontal Impact Accelerator (HIA) Coordinate System

The WPAFB HIA facility employs a Right-Hand Rule thumbs up, +X forward coordinate system 

for the sled. In addition, the impactor accelerates the sled from an initial velocity of zero, so no 

additional consideration is needed to correct for initial velocity. Because the coordinate system is 

fixed to the sled, the seat coordinate system varied depending on the seat orientation to the sled.

3.2.1.1.1.1 WPAFB –X (Frontal) Impact Runs

For this test condition, the ATD was facing the ram with its back toward the end of the track 

(Figure 3.2-1). Equation 3.2-1 to Equation 3.2-3 allow conversion from the seat coordinate system 

to the SAE coordinate system.

Equation 3.2-1. WPAFB –X to Sled Coordinate Transformation

a⃗ sled= [

−1 0 0

0 −1 0

0 0 1

] a⃗ -X 

Equation 3.2-2. WPAFB Sled to SAE Coordinate Transformation

a⃗ SAE= [

1 0 0

0 −1 0

0 0 −1

] a⃗ sled 

Equation 3.2-3. WPAFB –X to SAE Coordinate Transformation

a⃗ SAE= [

−1 0 0

0 1 0

0 0 −1

] a⃗ -X

Figure 3.2-1. WPAFB Horizontal Sled X (Frontal) Test Condition Coordinate System

+x 

(forward)

+z (up)

+y (left)
36

3.2.1.1.1.2 WPAFB +X (Rear) Impact Runs

For this test condition, the ATD was facing the end of the track (Figure 3.2-2), so the seat and sled 

coordinate systems are aligned. Equation 3.2-4 to Equation 3.2-6 allow conversion from the seat 

coordinate system to the SAE coordinate system.

Equation 3.2-4. WPAFB +X to Sled Coordinate Transformation

a⃗ sled= [

1 0 0

0 1 0

0 0 1

] a⃗ +X 

Equation 3.2-5. WPAFB Sled to SAE Coordinate Transformation

a⃗ SAE= [

1 0 0

0 −1 0

0 0 −1

] a⃗ sled 

Equation 3.2-6. WPAFB +X to SAE Coordinate Transformation

a⃗ SAE= [

1 0 0

0 −1 0

0 0 −1

] a⃗ +X 

Figure 3.2-2. WPAFB Horizontal Sled +X (Rear) Test Condition Coordinate System

+x 

(forward)

+z (up)

+y (left)
37

3.2.1.1.1.3 WPAFB Y (Lateral) Impact Runs

For this test condition, the ATD was facing to the right with the left shoulder toward the end of the 

track (Figure 3.2-3). Equation 3.2-7 to Equation 3.2-9 allow conversion from convert the seat 

coordinate system to the SAE coordinate system.

Equation 3.2-7. WPAFB –Y to Sled Coordinate Transformation

a⃗ sled= [

0 −1 0

1 0 0

0 0 1

] a⃗ -Y 

Equation 3.2-8. WPAFB Sled to SAE Coordinate Transformation

a⃗ SAE= [

1 0 0

0 −1 0

0 0 −1

] a⃗ sled 

Equation 3.2-9. WPAFB –Y to SAE Coordinate Transformation

a⃗ SAE= [

0 −1 0

−1 0 0

0 0 −1

] a⃗ -Y 

Figure 3.2-3. WPAFB Horizontal Sled Y (Lateral) Test Condition Coordinate System

+x 

(forward)

+z (up)

+y (left)
38

3.2.1.1.1.4 WPAFB +Z (Spinal) Impact Runs

For this test condition, the ATD was facing up with the head toward the end of the track 

(Figure 3.2-4). Equation 3.2-10 to Equation 3.2-12 allow conversion from the seat coordinate 

system to the SAE coordinate system. 

Equation 3.2-10. WPAFB +Z to Sled Coordinate Transformation

a⃗ sled= [

0 0 1

0 −1 0

1 0 0

] a⃗ +Z 

Equation 3.2-11. WPAFB Sled to SAE Coordinate Transformation

a⃗ SAE= [

1 0 0

0 −1 0

0 0 −1

] a⃗ sled 

Equation 3.2-12. WPAFB +Z to SAE Coordinate Transformation

a⃗ SAE= [

0 0 1

0 1 0

−1 0 0

] a⃗ +Z

 

Figure 3.2-4. WPAFB Horizontal Sled +Z (Spinal) Test Condition Coordinate System 

+x 

(forward)

+z (up)

+y (left)
39

LS-DYNA® Specific Coordinate Systems

Because LS-DYNA® allows for a custom coordinate system definition, additional consideration is 

needed for the numerical models run with each LS-DYNA® seat model. 

3.2.1.2.1 WPAFB LS-DYNA® Seat Model Coordinate System

Because the WPAFB testing has an initial velocity of zero, accelerations were from zero, so no 

additional consideration was needed to correct for initial velocity. Equation 3.2-13 shows the 

transformation employed to convert from SAE coordinates to the WPAFB LS-DYNA® model 

coordinates (Figure 3.2-5). Equation 3.2-14 shows the transformation to convert the Constellation 

Crew Injury Prediction (CCIP) ATD output to SAE coordinates. Equation 3.2-15 shows the 

transformation to convert the FTSS ATD output to SAE coordinates. Equation 3.2-16 shows the 

transformation used to convert the THUMS output to SAE coordinates.

Equation 3.2-13. WPAFB Seat SAE to Model Coordinate Transformation

a⃗ model= [

1 0 0

0 −1 0

0 0 −1

] a⃗ SAE 

Equation 3.2-14. WPAFB Seat CCIP to SAE Coordinate Transformation

a⃗ SAE= [

1 0 0

0 1 0

0 0 1

] a⃗ CCIP 

Equation 3.2-15. WPAFB Seat FTSS to SAE Coordinate Transformation

a⃗ SAE= [

1 0 0

0 −1 0

0 0 −1

] a⃗ FTSS

 

Equation 3.2-16. WPAFB Seat THUMS to SAE Coordinate Transformation

a⃗ SAE= [

−1 0 0

0 1 0

0 0 −1

] a⃗ THUMS

 

Figure 3.2-5. WPAFB LS-DYNA® Seat Model Coordinate System

+x 

(forward)

+z (up)

+y (left)
40

MADYMO Specific Coordinate Systems

MADYMO employs a Right-Hand Rule thumbs up, +X forward coordinate system 

(Figure 3.2-6). For WPAFB, no initial velocity condition must be accounted for, so 

Equation 3.2-17 was used. To convert the MADYMO output to SAE coordinates, 

Equation 3.2-18 was used.

Equation 3.2-17. SAE to WPAFB MADYMO Model Coordinate Transformation

a⃗ WPAFB = [

1 0 0

0 −1 0

0 0 −1

] a⃗ SAE 

Equation 3.2-18. MADYMO Output to SAE Coordinate Transformation

a⃗ SAE = [

1 0 0

0 −1 0

0 0 −1

] a⃗ MADYMO 

Figure 3.2-6. MADYMO Model Coordinate System

+x 

(forward)

+z (up)

+y (left)
41

BDR Model Specific Coordinate System

The BDR model employs a Right-Hand Rule thumbs up, +X forward coordinate system 

(Figure 3.2-7). For WPAFB sled testing, Equation 3.2-19 was used. To convert the BDR output to 

SAE coordinates, Equation 3.2-20 was used.

Equation 3.2-19. SAE to WPAFB BDR Model Coordinate Transformation

a⃗ WPAFB= [

1 0 0

0 −1 0

0 0 −1

] a⃗ SAE 

Equation 3.2-20. BDR Model Output to SAE BDR Model Coordinate Transformation

a⃗ SAE= [

1 0 0

0 −1 0

0 0 −1

] a⃗ Brinkley

 

Figure 3.2-7. BDR Model Coordinate System

Model Development

WPAFB Configuration

3.2.2.1.1 Introduction and Description of Test Facility 

The purpose of the WPAFB sled tests was to provide ATD sled test data under a variety of loading 

conditions so a comparison could be made between measured ATD impact responses and 

predictions made with numerical models. The Hybrid III and side-impact dummy (SID) ATDs 

were used for this study. As currently envisioned by NASA, the Hybrid III ATD will be used to 

evaluate the effectiveness of Orion CM crew protection systems. The Hybrid III ATD was 

originally designed to evaluate frontal impacts (X-axis impacts ±45 degrees) for the automotive 

industry. The primary impact vectors that are expected during an Orion CM landing are 

combinations of +X- and +Z-axes acceleration vector components and lateral Y-axis components. 

Therefore, concerns have been raised regarding the Hybrid III ATD response adequacy and 

corresponding numerical simulations to these loading directions. Furthermore, since the numerical 

+z (up)

+y (left)

+x 

(forward)
42

models are validated to represent the Hybrid III ATD rather than the specialized dummies used by 

the automobile industry to evaluate sideward and rearward impacts, concerns have been raised 

regarding its usefulness when the impact vectors are in these other directions. No dummies have 

been validated for all the combinations of loads, which are expected for Orion landings. While the 

sled testing did not determine the overall usefulness of using the Hybrid III for CxP, it did 

determine the numerical model accuracy for predicting the ATD response in loading directions 

other than frontal impact.

The test impacts varied in acceleration vector direction and magnitude, impulse duration, and body 

support and restraint configuration. The basic study approach was to collect the ATD’s dynamic 

accelerations and loads in an impact environment and compare these responses to the predicted 

ATD responses in the identical environment using the numerical models. Sled testing was 

performed using the AFRL HIA facility. The HIA is a multi-axial, high-G test facility that has 

been used for numerous human- and ATD-impact studies. For this study, a flat seat-pan/seat-back 

fixture with added leg support fixture and optional side support fixture was used. Two-seat 

configurations were used and are referred to as the Baseline 2 seat and the Baseline 2 plus sidesupports seat (Baseline 2 + S). The Baseline 2 seat has a rigid backrest, a rigid 

seat pan, a rigid footrest, and a leg support panel. The Baseline 2 + S seat is identical to the Baseline 

2 seat with the addition of side supports that restrict the motion of the head, shoulders, hips, and 

legs during lateral impact.

A five-point restraint system was used to confine the ATD in the seat. The restraint system 

consisted of two shoulder straps, a lap belt, and a negative-G strap. Provisions were made to 

tether/restrain the ATD’s hands and feet to designated attachment points on the seat structure and 

footrest. Instrumented ATDs and associated fixtures were provided. Drawings for the seat 

configurations for FEM purposes and the mechanical properties of the restraint system materials 

for numerical modeling were provided (Figure 3.2-8). Padding used for the ATD’s head protection 

was tested to determine the material’s mechanical properties.

Figure 3.2-8. Computer-Aided Design (CAD) Model of Seat Used for WPAFB Sled Testing

The test setup included two instrumented ATDs: a 50th-percentile male Hybrid III automotive ATD 

and a 50th-percentile male Hybrid III aerospace ATD. The majority of testing was performed using 

the Hybrid III automotive ATD. The automotive ATD was used for testing in all four directions 
43

(i.e., frontal, rearward, spinal, and lateral) and the aerospace ATD was used for comparison to the 

automotive ATD for spinal direction impacts. The aerospace ATD is identical to the automotive 

ATD except for the mechanical components used to represent the spine.

Follow-on tests were later performed using a European side-impact dummy (EuroSID) and a 50th-

percentile male Hybrid III automotive ATD exposed to lateral loadings so that a comparison could 

be made between the Hybrid III response in the lateral direction and the EuroSID. The current plan 

for assessing crew injuries during Orion landings is to use the Hybrid III FE ATD. The crew 

response will be obtained by loading the ATD with landing-load acceleration obtained from 

vehicle-landing simulations, which are currently being performed using LS-DYNA®. The landingload accelerations, which are combinations of X-, Y-, and Z-axes accelerations, will be applied to 

LS-DYNA® models consisting of the Hybrid III FEMs placed in the Orion seat designs. Injury 

criteria will be extracted from the simulations (e.g., neck forces, head accelerations, pelvic motion) 

and compared against the recommended injury criteria established in the HSIR requirements. An 

issue related to this approach is that the Hybrid III is not validated for lateral impact. While Orion 

landings generally produce primarily rear and spinal direction loading, there are landing 

conditions, particularly when vehicle roll is prevalent, where lateral loading occurs. Furthermore, 

the HSIR requirements include lateral direction-related criteria (e.g., neck loadings and body 

movements) which must be satisfied. 

The principle objective of the SID tests is to compare the Hybrid III lateral response to the EuroSID 

response with particular emphasis on comparing the HSIR lateral-loading criteria between the 

Hybrid III and EuroSID for identical seats and loading conditions. If it is determined that the 

Hybrid III produces similar results to the EuroSID for similar lateral loading, the Hybrid III can 

then be used for the above approach involving multi-axial landing simulations and data extraction. 

If the Hybrid III does not produce similar results to the EuroSID, then the above planned approach 

may not be workable and an alternate approach will need to be developed. It is expected that the 

results will be different and some adjustments will need to be made to the injury criteria. 

The ATD, seat system, and restraints were instrumented sufficiently to provide reliable data 

collection over the duration of the load application and subsequent ATD dynamic response. In 

addition to standard Hybrid III measurement response, forces were measured at the restraint 

attachment points and along the interface between the lateral seat supports and the instrumented 

ATD. Acceleration measurements were taken on the sled and seat. High-speed video and 

photographic documentation of the test series was collected for all test runs. The pre-impact 

positions of specific points located on the head, shoulders, hips, and knees were measured prior to 

each test with respect to the seat to ensure uniformity of the location of these anatomical points 

from test to test.

Two test acceleration profiles were selected based on predicted nominal and contingency landing 

conditions of the Orion capsule, and best acceleration properties to elicit a range of ATD response 

properties for validating the numerical models. Preliminary predictions of peak acceleration 

conditions are 10 G nominal and 20 G contingency. The acceleration profile used for most test 

conditions will have a time-to-peak of approximately 70 msec. The time-to-peak will be varied in 

several cells to evaluate the frequency sensitivity of the ATDs and the numerical models. Example 

pulses and configurations are shown in Figure 3.2-9 to Figure 3.2-11.
44

Figure 3.2-9. Sled with a 20-G Impulse in the Z-Direction. Sled Motion is to the Right. 

(a) Acceleration Pulse, (b) ATD Position

Figure 3.2-10. Seat with Side Impact Constraints Added and Oriented for Lateral Impact Loads

Figure 3.2-11. Sled Test with a 10-G Impulse in the X-Direction. Sled Motion Is to the Right. 

(a) Acceleration Pulse, (b) ATD Position

Table 3.2-1 shows the four directions of impact loading that were applied to the automotive ATD 

and the spinal and front impact loadings that were applied to the aerospace ATD. The identical 

seat configuration was used for all impact directions except for the lateral direction, where some 

of the tests were performed without side supports, some with side supports at a full gap distance, 

and some with 50-percent gaps. Peak G impact and the rise time-to-peak G levels were varied.
45

Table 3.2-1. WPAFB HIA Test Matrix

tp = time to peak G tp = 70ms unless otherwise noted Metering Pin

Baseline 2+S is baseline 2 seat with added side supports 9 (tp = 40ms, Vel. = 24 ft/sec, DRx = 14.4, DRz = 14.7)

Baseline 2+SG is Baseline 2 seat with gap as specified in revised HSIR document 11 (tp = 70ms, Vel. = 33 ft/sec, DRx = 13.4, DRz = 12.3)

Dummies removed from the seat and instrumentation rezeroed after each test 19 (tp = 105ms. Vel. = 43 ft/sec, DRx = 10.6, DRz 10.1)

a - test without brace at footrest

b - headrest one inch aft of dummy head

c - Test redone because acceleration input exceeded requirement

Test 8203 redone because foot tethers broke 

d - Lost dummy data cable

e - Lost lumbar z accelerometer

f - Cell Q1 add to test lateral impact with no lateral head support plate

Test 8240 conducted with upper limbs restrained to seat fixture with velcro straps

Test 8241 condcuted with upper limbs restained with velcro strap run beneath thigh; felt pad (0.5"thk) added to shoulder load plate

g - Cell U tests conducted with 50% recommended Gaps between manikin and lateral reaction load plates

h - cross bar added to lateral reaction plate for head; tie down straps added to lower seat for additional motion control during impact

i - Tests conducted with 100% recommended gaps 

j - Headreast modified to provide 1" offset from plane of headrest to plane of seatback; 

new bronze lower neck bracket added to allow additional rearward rotation of manikin head to provide contact between back of head and plane of headrest
46

Four numerical models were used in the present study for comparison to the data obtained in the 

sled tests at WPAFB. Results from each model were compared to the test data and to each of the 

other models to determine the strengths and weaknesses and their applicability to CxP and the 

types of landings expected during Orion landings. Three of the models are numerical models of 

the Hybrid III automotive ATD: the Livermore Software Technology Corporation (LSTC) Hybrid 

III FEM, the FTSS Hybrid III FEM, and the TNO Automotive Safety Solutions (TASS) 

MADYMO multi-rigid body model. The fourth model was the THUMS human body FEM.

3.2.2.1.2 CxP Crew Injury Protection ATD Model 

The Hybrid III automotive FE ATD used in this assessment was developed and validated by LSTC 

(see Figure 3.2-12). A full set of LSTC Hybrid III FE dummies has been modeled in 

LS-DYNA® format and set up according to the SAE calibration procedures for Hybrid III ATDs. 

The FE ATD used in this study was modeled using a combination of rigid and deformable parts. 

This ATD was validated by LSTC using standard impact tests in four different locations and 

positions, including head impact, neck flexion and extension, chest impact, and knee impact. These 

particular tests were selected because the intended use for this model is for frontal impact 

automobile crashes. However, this ATD was not validated for impacts in all possible orientations. 

The intent of the present testing was to identify how this model emulates the Hybrid III ATD for 

loadings other than the frontal impact direction. The LSTC Hybrid III FE ATD model can be used 

without charge, which provides an advantage over other numerical models. 

Figure 3.2-12. LSTC Hybrid III FEM in WPAFB Seat Model

The FE ATD was imported into the preprocessor LS-PrePost® to orient it in the desired positions 

so that the numerical model matched the position of the test ATD. In addition, limb operations 

were available to position the arms and legs properly. The neck was in a standard position that 

could not be altered. In the new ATD models currently in development by LSTC, the neck can be 
47

rotated to better match the test setup. However, this model has not been validated and was not used 

for this study. Since the validated FE neck model cannot be positioned arbitrarily, for the model 

to be entirely consistent with the test, it was necessary to position the test ATD’s neck to match 

the FEM.

Seatbelts were modeled with the preprocessor LS-PrePost®. The FE ATD was imported into the 

preprocessor; then “segment sets” were defined on the FE ATD where the seatbelts were to be 

positioned. Once these sets were defined, the seatbelts were positioned on the set and tensioned to 

make the belts fit to the contour of the ATD. The belts can be modeled with seatbelt elements (1-

D beamlike elements) or 2-D shell elements. The shell model is the most appropriate model for 

capturing the seatbelt/ATD interaction. The pre-tension in the seatbelt is modeled by defining a 

local coordinate system on the belt ends. A load is defined in this coordinate system with a 

magnitude equal to the preload applied in the test setup. The load must be ramped to the full 

magnitude. Applying full-magnitude tension at time zero can lead to unrealistic oscillations in the 

system response and to inaccurate simulation results.

3.2.2.1.3 FTSS Hybrid III 50th-percentile LS-DYNA® Model

Following a peer review conducted by the NESC [ref. 70], the Orion Project coordinated a pilot 

study using the FTSS1 LS-DYNA® Hybrid III 50th-percentile v7.0 ATD model for numerical 

comparisons with the WPAFB sled test data (see Figure 3.2-13). FTSS is a primary provider of 

physical test ATDs and their numerical models to the automotive industry ATDs. FTSS has 

extensive history and experience with test and numerical ATDs, including the Hybrid III. The 

FTSS Hybrid III was selected for the pilot study to identify this model’s performance for 

replicating the test data and to perform a comparison between the FTSS model and the other 

models investigated in this study.

Figure 3.2-13. FTSS Model in WPAFB Sled Test Seat

 

1 Note that since the original writing, First Technology Safety Systems has merged with Denton to become 

Humanetics.
48

The FTSS pilot study was conducted in two phases. In the first phase, predictive simulations were 

performed without having access to the test time history data as only test videos were provided. 

After that, the test data was supplied and the models were further refined to match the test as 

closely as possible. The purpose of the first phase was to identify how well the numerical models 

can predict test results (pre-test) without the use of the test data to guide model development. The 

second phase was to determine how well the models could match test results when modeling 

refinement is performed post-test. Both phases were valuable as they show how the models 

performed when test data are not available, or when test data are available and allow for modeling 

refinement. In some cases, no test data may be available and in other cases test data are used to 

improve modeling fidelity and the model is subsequently used for similar configurations or for 

alternate loading conditions.

From the test videos and provided data it was evident that the simulation was less constrained and 

allowed more ATD motion in the pelvis and chest. The tests showed similar trends, but key 

behaviors were delayed and the associated data signals contained large spikes. Visual indicators 

like the belt exit angle at the shoulder and deformation of the shoulder suggested that the FE ATD 

needed to be moved down and belt tension introduced to reduce allowable ATD motion. Situations 

such as these are common modeling complications and further exemplify the challenge of using 

these complex models without the benefits of having test data specific to the application to anchor 

the model.

To improve the correlation between the observed test response and the simulation, several 

improvements were made to the model. First, the ATD position was updated to achieve more 

contact with the back of the fixture.

The pelvis was rotated allowing additional torso rotation. The torso rotation allowed additional 

neck rotation, which allowed the pelvis to be moved rearward allowing a shorter lap belt. The 

shoes were removed from the baseline ATD model allowing the legs to be rotated down. The legs 

were spread apart as evident in the test video (see Figure 3.2-14). The arms and hands were 

repositioned accordingly and pre-loading was added prior to the start of the input acceleration.
49

Figure 3.2-14. FTSS Hybrid III Re-positioned to Better Match Test Video

The test videos showed a black band restraining the hands to the leg (see Figure 3.2-15). Also 

visible were similar bands restraining the lower legs to the test fixture. Six-DOF directional springs 

were defined to constrain the wrists to the legs. The spring was set to have an axial direction that 

moved with the tibia axis. For the lower leg, a simple unidirectional spring was used to constrain 

the lower leg to the seat structure.

Figure 3.2-15. Physical ATD Hand and Feet Constraints

Three methodologies were utilized to develop three models to match the initial conditions viewed 

in the videos (see Figure 3.2-16). 

FEM 1—Belt pre-tensioners were used to pull in the belts with loads approximated from the 

provided test data. This required that the applied input accelerations be delayed or offset in time 

while the ATD was pulled into the seat. A locking retractor was used to lock the belt length just 

prior to the start of the applied accelerations.

FEM 2—As in method FEM 1, pre-tensioners and retractors were used. Additionally, the pelvis 

was displaced downward (as a result of gravity) into the fixture over a period of 62 msec. The 

retractor was modified to lock at 62 msec.
50

FEM 3—The ATD mesh was pre-deformed to a position that was similar to those achieved in 

FEMs 1 and 2 methods. The pre-tensioners and retractors were still utilized to keep the belts as 

tight as possible and achieve the preloads reflected in the test data. 

Note that FEMs 1 and 2 are identical at the initial position, but FEM 3 is the pre-deformed ATD 

mesh and was visually lower and more rearward in the fixture.

Figure 3.2-16. Comparison between Three Preload Methods

All three FEM methods added additional run time to the beginning of the simulation to allow for 

settling the ATD due to gravity and/or the introduction of belt tension. In addition, all three 

methods improved the correlation by reducing belt slack, reducing the free belt length, and settling 

the ATD into a configuration similar to the test condition. For FEM 1, the approach did pull tension 

in the belts. However, it is not a robust method for tuning a defined preload. In addition, despite 

the belts being visibly tightened, little load was measured in the belt ends. The pre-tension and 

retractor allowed tight control over the belt preload and slack reduction. The rate of application of 

the pre-tension can introduce undesired dynamic effects at the start of the simulation. Slowly 

ramping the pre-tension requires more simulation time at the start to allow the ATD to fully settle. 

The ATD did not settle as much in FEM 1 as it did in FEMs 2 and 3.

It is possible that the sequence of timing of the tension between the lap and shoulders could affect 

the final position, but that scenario was not investigated.

For FEM 2, the desired deformed configuration can only be obtained after test or after knowledge 

of similar test conditions has been obtained. The introduction of the pelvis motion in the 62-msec 

period introduced a bouncing effect of the upper body and subsequent noise at the start of the 

simulation. As the shoulders bounced, some of the desired belt tightening was lost. This motion 

could be removed with approaches, such as additional run time and simultaneous enforced 

displacement of part of the upper body skeleton. Since the gravity load was introduced through 

fixture motion, the desired pelvis relative motion needs to take into consideration the fixture 

motion.
51

An initial simulation was required to understand the unmodified ATD behavior to judge how much 

motion to introduce.

For FEM 3, the desired deformed configuration can only be obtained after test or after knowledge 

of similar test conditions has been obtained. This method requires a pre-simulation to get the ATD 

into the desired deformed configuration. This deformed model must be updated every time there 

are system changes that affect the boundary conditions enough to affect the initial deformed shape 

(e.g., belt anchor locations, supporting structure modifications, occupant seating position, etc.). 

The configuration was obtained through trial and error by comparison to the available video. If 

setup information is available (e.g., test data of the pelvis location and other key points and angles), 

then the ATD could be moved into the position more reliably. As the mesh was pre-deformed, 

stresses that would normally be present were not, causing excessive further mesh deformation. 

This would have to be monitored to determine if the affected areas introduce negative effects to 

the overall simulation.

As the mesh is deformed, some elements become distorted, decreasing the solution time step, thus 

increasing the run time. The time step can be artificially increased through added mass techniques. 

This mass-scaling technique would have to be monitored to understand the potential effects to the 

solution. Run time is needed to tension the belts to remove excess slack. The 

80-msec was kept in this case for comparison to the other FEM methods. Less time could be needed 

as less slack needs to be taken up because less ATD settling occurs.

The belt tension/slack condition was influencing the ATD motions. Understanding the belt lengths, 

end conditions, and pre-loading will help set up the simulation and improve correlation. Utilization 

of pre-tensioners and retractors is recommended to achieve the proper pretest conditions.

The pre-deformed (FEM 3) approach allowed for good correlation with the available data. It can 

reduce some run time added at the beginning for settling. However, the extent to which the mesh 

had to be deformed to represent the test video added significant run time to the overall solution. In 

a development environment, the ATD may not be as tightly constrained thus requiring less initial 

deformation and consequently less mesh distortion. Mass scaling could be introduced to offset 

some of the added run time with minimal detrimental effects to the overall solution.

The use of lower-leg and arm restraints should be considered for future development. Constraining 

the leg motion can influence the pelvis motion, loading, and load timing. Similarly, the motion and 

loading due to the arm behavior can affect the chest, neck, and head responses.

The belt condition at the location of the five-point restraint buckle was undefined. In the 

simulation, the belts were constrained at that location resulting in no relative motion. The 

constraint condition could affect the belt slip/loading and subsequent ATD motions. 

The physical presence of the five-point hardware may be introducing ATD contact or restraint at 

the pelvis and torso joint. In the simulation, the torso skin was allowed to flow freely over the top 

of the pelvis. It is unclear if the physical hardware might affect the behavior.

The 1-D seatbelt material that was supplied to FTSS appeared to have an improperly defined mass. 

It should be cross-referenced to ensure that no unintended inertia effects are introduced.

Capturing more information of the physical test setup would allow for more accurately configuring 

the simulation and eliminate some analyses, which were performed to study possible setup 

conditions. A better understanding of the physical setup procedure may increase understanding on 

an approach that could facilitate a better simulation method.
52

The following might be considered in future testing to help with post-test correlation;

 Collect test setup measurements and photographic documentation.

 Additional camera with an orthogonal viewpoint would be beneficial in helping to 

understand the physical behavior especially if the development inputs become more 

complex.

 A grid board as background in the camera view to provide a reference for distances,

 Strobe lighting timed with the fixture trigger to provide a definitive indicator of time 

zero.

 Belt marking graduated to a known distance to help understand belt motion and relative 

motions between the belt and the ATD.

 Rolling, clipping, or treatment of excess belt material to improve clarity of the image.

 Painting fixtures in a bright color so that it provides a background and helps discern 

different parts of the image.

 Utilizing secondary measuring devices could validate if there were fixture motion.

Following the pilot study, it was decided to continue use of the FTSS ATD model for all WPAFB 

sled test comparisons using the FEM 3. All references to FTSS in succeeding sections relate to this 

ATD configuration.

3.2.2.1.4 MADYMO Modeling Using the Hybrid III 50th-percentile ATD

The 50th-percentile Hybrid III Ellipsoid model (Figure 3.2-17) version 6.4 as supplied by TASS 

was used without modification. It consists of 37 rigid bodies and 69 contact ellipsoids. Various 

kinds of joints were used to join the bodies. Properties of the joints are generally not known, as 

they are internal to the model and unavailable to the user. For example, the neck is represented by 

five kinematic joints located in the centers of the four rubber disks and on the rotation axis of the 

nodding joint. The rubber disks are represented by spherical joints and the nodding joint by a 

revolute joint [ref. 71].

Extensive component tests have been done by TASS to characterize the joint functions and 

segment contact characteristics. These include: neck, clavicles, shoulders, elbows, wrists, thorax, 

abdomen, lumbar spine, hip, knees, ankles, and feet. Details of the testing procedures and results 

can be obtained from the MADYMO Theory Manual [ref. 72]. The standard ATD calibration tests 

(shown for the LSTC model) were among the validation tests in addition to full ATD tests with 

belts on a rigid seat, vertical drop tests, and airbag tests. As with the LSTC model, the ATD was 

not validated for all impact orientations since the Hybrid III ATD was designed primarily for 

frontal impact use.
53

Figure 3.2-17. MADYMO Hybrid III 50th-percentile ATD – Version 6.4

The FE seat model used for the LS-DYNA® simulations was imported into MADYMO. Thus, the 

seat geometry was identical to that of the LS-DYNA® model. Flat surfaces (e.g., seat back, 

headrest, and lateral supports) were converted to plane surfaces. Three-inch-wide FE belts were 

generated using MADYMO-provided software and fitted as a five-point restraint system. Belt 

attachment points were in line with those used in the test, but more distant to accommodate 

retractors that were used to pre-tension the belts. At approximately 150 msec prior to the onset of 

the sled acceleration, the pre-tensioners were activated to apply approximately 100 N of pretension on each belt anchor point after the occupant “settled.” Gravitational acceleration was 

applied. It was found that it took approximately 150 msec for the occupant to stabilize to the 

reaction of gravity and belt pre-tension. Since the seat was assumed to be rigid (except where 

padding was used), the contact characteristics supplied with the MADYMO ATD were used. The 

specified sled test accelerations were applied to the seat as shown in Figure 3.2-18.
54

Figure 3.2-18. WPAFB Configuration

3.2.2.1.5 THUMS Model

The THUMS model was developed for use in automotive testing and has been extensively 

validated against PMHS data [ref. 73] primarily for automotive applications. The THUMS model 

results were compared to sled test results for the Hybrid III ATD to evaluate predictions from 

THUMS, the Hybrid III, and test data. Injury responses were extracted from the THUMS model, 

were examined, and are discussed in detail in Appendix E. 

The FEM used in this analysis was a modified version of the THUMS version 1.61c. This model 

had been previously modified for NASA simulations to stabilize the model response in 

Z-loading configurations. The original THUMS model could not successfully be run to completion 

without incurring numerical problems because of elements exceeding their failure limits. To 

circumvent this problem, modifications were implemented, including different response curves for 

the pelvis flesh and lung tissue. Additionally, all element failure criteria were removed to assess 

loading values in all the elements for the entire simulation. 

The seat configuration and material properties used were the same as the WPAFB sled test seat. 

This seat consisted of a rigid back with a cushioned headrest and seat bottom. A floorboard was 

included and the feet of the THUMS model were restrained to this component. The floorboard 

used with the THUMS model was translated in the X-direction to accommodate the original model 

knee bend. It was simpler to relocate the floorboard since the THUMS model cannot be easily 

articulated to arbitrary positions. It is expected that relocating the floorboard from its position 

would not considerably affect the results of the simulations. The WPAFB restraint system is a fivepoint harness configuration. Figure 3.2-19 illustrates the initial THUMS model before placement 

in the seat. Figure 3.2-20 and Figure 3.2-21 illustrate the WPAFB seat configurations. 
55

Figure 3.2-19. Isolated 

THUMS Model Prior to 

Integration with the Seat

Figure 3.2-20. Seat 

Configuration for Lateral 

Loading Simulations

Figure 3.2-21. Seat Configuration 

for Anterior/posterior and 

Inferior/superior Loading Pulses

There were four landing scenarios modeled in the simulations with corresponding tests conducted 

at WPAFB. The four tests selected were 8202, 8208, 8212, and 8245 (Table 3.2-1). These tests 

were selected to represent loading in the frontal (8202), spinal (8208), rear (8212), and lateral 

(8245) directions (see Figure 3.2-22). The input for the THUMS simulations, similar to that for 

the FTSS, LSTC Hybrid III FEM, and MADYMO simulations was the measured acceleration of 

the seat. This acceleration was applied directly to the modeled seat. 

Figure 3.2-22. Illustration of Seat Directions for Each Pulse

The metrics selected to compare to the test data were the head and chest acceleration and seatbelt 

forces. Peak head acceleration was measured by the nodal acceleration of a single node located at 

the center of the head (node 8890001) as shown in Figure 3.2-23. Accelerations at this node were 

compared to the accelerations obtained during the WPAFB sled tests with the Hybrid III. 

Additionally, the belt tension and the seat acceleration were evaluated for each simulation and 

compared to the test data. Visual inspection of the sled test videos was used to compare the overall 

kinematics of the THUMS and sled results. 
56

Figure 3.2-23. Location of the Node Used to Record Head Acceleration in the THUMS Model

Filtering Specifications

With the exception of the LSTC Hybrid III FEM ATD, all data were filtered in accordance with 

SAE J211/1 filtering specifications [ref. 69]. The one exception was the chest accelerations, 

which were filtered at channel frequency class (CFC) 1000 instead of CFC180 (see 
57

Table 3.2-2 through 

Parameter Filter Specification

Head Accelerations SAE J211/1 CFC1000

Head Velocity SAE J211/1 CFC1000

Head Displacement SAE J211/1 CFC1000

Chest Acceleration SAE J211/1 CFC1000

Sternal Deflection SAE J211/1 CFC600

Chest Displacement SAE J211/1 CFC1000

Neck Forces SAE J211/1 CFC1000

Neck Moments SAE J211/1 CFC600

Lumbar Forces SAE J211/1 CFC600

Lumbar Moments SAE J211/1 CFC600

Pelvis Acceleration SAE J211/1 CFC1000

Pelvis Displacement SAE J211/1 CFC1000

Femur Forces SAE J211/1 CFC600

Tibia Forces SAE J211/1 CFC600

Restraint Forces SAE J211/1 CFC60

Table 3.2-5). For the LSTC Hybrid III FEM ATD, CFC180 filters were applied to all but the leg 

axial and restraint forces, which were filtered using CFC60 filters per the developer’s 

recommendations [refs. 77 and 78] (see Table 3.2-6).
58

Table 3.2-2. Physical ATD Filter Specifications

Parameter Filter Specification – SAE J211/1

CFC60 CFC600 CFC1000

Head Accelerations X

Head Velocity X

Head Displacement X

Chest Acceleration X

Sternal Deflection X

Chest Displacement X

Neck Forces X

Neck Moments X

Lumbar Forces X

Lumbar Moments X

Pelvis Acceleration X

Pelvis Displacement X

Femur Forces X

Tibia Forces X

Restraint Forces X

Table 3.2-3. FTSS ATD Filter Specifications

Parameter Filter Specification

Head Accelerations SAE J211/1 CFC1000

Head Velocity SAE J211/1 CFC1000

Head Displacement SAE J211/1 CFC1000

Chest Acceleration SAE J211/1 CFC1000

Sternal Deflection SAE J211/1 CFC600

Chest Displacement SAE J211/1 CFC1000

Neck Forces SAE J211/1 CFC1000

Neck Moments SAE J211/1 CFC600

Lumbar Forces SAE J211/1 CFC600

Lumbar Moments SAE J211/1 CFC600

Pelvis Acceleration SAE J211/1 CFC1000

Pelvis Displacement SAE J211/1 CFC1000

Femur Forces SAE J211/1 CFC600

Tibia Forces SAE J211/1 CFC600

Restraint Forces SAE J211/1 CFC60
59

Table 3.2-4. MADYMO ATD Filter Specifications

Parameter Filter Specification

Head Accelerations SAE J211/1 CFC1000

Head Velocity SAE J211/1 CFC1000

Head Displacement SAE J211/1 CFC1000

Chest Acceleration SAE J211/1 CFC1000

Sternal Deflection SAE J211/1 CFC600

Chest Displacement SAE J211/1 CFC1000

Neck Forces SAE J211/1 CFC1000

Neck Moments SAE J211/1 CFC600

Lumbar Forces SAE J211/1 CFC600

Lumbar Moments SAE J211/1 CFC600

Pelvis Acceleration SAE J211/1 CFC1000

Pelvis Displacement SAE J211/1 CFC1000

Femur Forces SAE J211/1 CFC600

Tibia Forces SAE J211/1 CFC600

Restraint Forces SAE J211/1 CFC60

Table 3.2-5. THUMS Model Filter Specifications

Parameter Filter Specification

Head Accelerations SAE J211/1 CFC1000

Head Velocity SAE J211/1 CFC1000

Head Displacement SAE J211/1 CFC1000

Chest Acceleration SAE J211/1 CFC1000

Sternal Deflection SAE J211/1 CFC600

Chest Displacement SAE J211/1 CFC1000

Neck Forces SAE J211/1 CFC1000

Neck Moments SAE J211/1 CFC600

Lumbar Forces SAE J211/1 CFC600

Lumbar Moments SAE J211/1 CFC600

Pelvis Acceleration SAE J211/1 CFC1000

Pelvis Displacement SAE J211/1 CFC1000

Femur Forces SAE J211/1 CFC600

Tibia Forces SAE J211/1 CFC600

Restraint Forces SAE J211/1 CFC60
60

Table 3.2-6. LSTC Hybrid III ATD Filter Specifications

Parameter Filter Specification

Head Accelerations SAE J211/1 CFC180

Head Velocity SAE J211/1 CFC180

Head Displacement SAE J211/1 CFC180

Chest Acceleration SAE J211/1 CFC180

Sternal Deflection SAE J211/1 CFC180

Chest Displacement SAE J211/1 CFC180

Neck Forces SAE J211/1 CFC180

Neck Moments SAE J211/1 CFC180

Lumbar Forces SAE J211/1 CFC180

Lumbar Moments SAE J211/1 CFC180

Pelvis Acceleration SAE J211/1 CFC180

Pelvis Displacement SAE J211/1 CFC180

Femur Forces SAE J211/1 CFC60

Tibia Forces SAE J211/1 CFC60

Restraint Forces SAE J211/1 CFC60

Model Correlation

Error Measure for Comparing Transient Response Histories

To assess model acceptability, the Sprague-Geers (S&G) method was employed. The method 

assesses magnitude and phase errors and consists of three metrics. Magnitude Error compares the 

areas under the curves and peak measurements. This measure is insensitive to phase differences. 

Phase Errors are measured without being affected by magnitude effects. Total Error is the root 

sum of squares of the magnitude and phase errors. The method is biased toward the reference data. 

Equation 3.2-21 through Equation 3.2-26 show the application of the method. 

Equation 3.2-21. S&G Total Error

T = √M2 + P2 

Equation 3.2-22. S&G Magnitude Error

M = √

Igg

Iff

− 1 

Equation 3.2-23. S&G Phase Error

P =

1

π

∙ cos−1

Ifg

√Iff ∙ Igg

 

Equation 3.2-24. S&G Iff Definition

Iff =

1

t2 − t1

∫ f

2

(t)dt

t2

t1

 

Equation 3.2-25. S&G Igg Definition

Igg =

1

t2 − t1

∫ g

2

(t)dt

t2

t1

 
61

Equation 3.2-26. S&G Ifg Definition

Ifg =

1

t2 − t1

∫ f(t) ∙ g(t)dt

t2

t1

 

Where:

f(t) is the reference data (sled test results)

g(t) is the signal being evaluated (numerical prediction of the response)

M is the magnitude error (-1 ≤ M ≤ 1)

P is the phase error (-1 ≤ P ≤ 1)

T is the total shape error (0 ≤ T ≤ 1)

This method was studied by the FAA with other model validation methods and was found to be 

the best at describing variations [ref. 79]. In addition, the FAA conducted a study comparing 

39 test/simulation time histories. A panel of 16 subject matter experts (SMEs) evaluated the model 

performance and the results were compared to the S&G results. This panel found good correlation 

between the S&G and their results, with the recommended classifications shown in Table 3.2-7.

Table 3.2-7. S&G Classifications

SME Rating S&G Percent Error

Excellent 04

Good 410

Fair 1020

Poor 2030

Very Poor >30

Table 3.2-8 shows the classifications that were adopted for this study to quantify the model 

correlation to the sled test data. All S&G results are shown in Appendices B and C. Some responses 

are not valid in each impact orientation because the responses were small and any discrepancy 

between the model and test data shows a large S&G error, which may be erroneous. 

For the discussions in subsequent sections of this report, the column titled Qualitative Description 

of Correlation in Table 3.2-8 were used to judge the results. The column in the table titled SME 

Qualitative Description lists the correlation ratings used by the FAA to measure the quality of 

numerical ATD predictions for commercial aircraft qualification. These ratings differ from the 

rating used in the present study. However, the qualitative descriptions of the correlation between 

test and numerical predictions were arbitrary and were selected for the present study based on 

reasonable expectations for the modeling predictions.

Table 3.2-8. Relation between S&G Numerical Value and Qualitative Description of Test-predication 

Correlation

S&G Percent Error Qualitative Description 

of Correlation

SMEs Qualitative 

Description

04 Good Excellent

410 Good Good

1020 Good Fair

2030 Fair Poor

3040 Poor Poor
62

WPAFB Sled Tests

The objective of comparing the numerical models to test data was to assess the numerical model’s 

ability to model and simulate the Hybrid III ATD. This objective was not completely 

straightforward since the numerical response was affected by the ATD model accuracy, and by the 

ATD seat, harness, and the interaction with these components. To eliminate modeling 

discrepancies resulting from limitations in the seat model, responses that occur as a result of the 

ATD impacting the seat or the padding on the seat wereignored. Although the data traces show 

complete transient responses, not all of the responses were used for the S&G comparison 

calculations, which are used to quantify the correlation between test and numerical predictions. 

When the data in the traces were not used for the quantitative comparisons, they were indicated by 

the data trace line changing from a solid to a dashed line (shown in Appendix C). 

3.2.4.2.1 Front Impact

Front impact sled data were compared to the numerical models for a variety of front loading 

conditions. Front impact is unlike the condition expected for Orion landings since the crew is 

oriented to land primarily with a rearward impact. However, front impact is one of the primary 

conditions for automobile qualification and is the condition that the Hybrid III ATD was designed 

to replicate. Therefore, it was expected that the numerical models would perform best for front 

impact since both the models and ATD have undergone validation for this condition.

3.2.4.2.1.1 Test Cases 

Six sled tests were run for front impact using the Hybrid III automotive ATD with a curved spine 

and an additional three tests using the Hybrid III aerospace ATD with the straight spine. All tests 

were run near the 10-G level except for test 8205, which was run near 20 G. The pulse duration 

was 70 msec unless otherwise noted in the test matrix. Test numbers 8201, 8202, and 8204 were 

duplicate tests to assess the repeatability of the physical sled tests and obtain data to compare the 

numerical models to the test data. Maximum G level and delta velocity and the seat acceleration 

transient profiles (Appendix C-1.1) were found to be repeatable indicating the reliability of the 

sled and loading generation system. For the 10-G and 20-G sled tests, the loadings were relatively 

mild, indicated by the low BDR (Appendix C-1.2) even for the 20-G load case where the BDR 

injury risk criterion was 0.64 and the allowable limit was 1.0. Since all the loadings are in a 

relatively low range, it is expected the ATD will not experience extreme loading conditions and 

the ATD and the numerical models should produce good results.

3.2.4.2.1.2 Harness Loads

The first step toward producing good numerical results is for the numerical model to constrain the 

ATD in the seat by having an accurate representation of the harness and how it constrains the ATD 

during loading. If the harness belt forces do not match what is measured in the physical sled test, 

then there will be a mismatch in how the ATD is constrained, making it doubtful that the internal 

ATD measurements will match the numerical model. 

For the left shoulder belt forces (Appendix C-1.38), all of the models predict the belt force within 

an approximate 10-percent range, with the exceptions of the MADYMO model and the FTSS 

model for the 20-G case. For the 20-G case the FTSS model under-predicts the belt force by nearly 

30 percent. The MADYMO model consistently predicted a left belt force that differed by nearly 

30 percent. This probably indicates that the harness was not sufficiently modeled and as such, the 

prediction was not a deficiency in the MADYMO ATD representation. For the right shoulder belt 

force, all of the models reasonably predict the belt force. Considering the numerical models should 
63

be symmetric, it is surprising the predicted right belt forces correlate while the left shoulder belt 

forces do not. Small out-of-place sled seat accelerations combined with the exact manner in which 

the left and right shoulder belts are placed in the model may be leading to larger discrepancies in 

the left and right belt forces than are seen in the test data where the measured left and right belt 

forces are closer to each other in magnitude. The measured left and right shoulder belt forces are 

similar for each of the three repeat tests.

None of the numerical models predicted the right or left lap belt forces reliably (Appendix 

C-1.36), except for the THUMS model, which was within 5 percent of the test data for the single 

frontal impact case where THUMS was used to generated numerical predictions. The test data 

were repeatable for each of the three repeat tests, and the left and right lap belt forces were similar 

in magnitude. With additional study of the harness and ATD positioning, the numerical models 

may be made to better match the test data. However, these results are indicative of the challenge 

of using the models for pre-test predictions when many of the specific conditions are not well 

known. The correlation between test and models is worse for the crotch belt force (Appendix C1.40) where there is essentially no correlation between test and model predictions. The FTSS 

model does the best job of predicting the measured crotch belt force while some of the other models 

are off by an order of magnitude. These large discrepancies between test and predictions may be a 

result of the models transferring the shoulder belt loads directly into the crotch belt and transferring 

less of the load into the lap belts. Whereas in the tests, the crotch belt was carrying less load 

compared to the lap belts. The lap and shoulder belts, and the crotch belt feed into a common 

attachment point above the ATD crotch, creating a complex condition that was difficult to model. 

Major issues for Orion occupant protection are the type, configuration, and preload of the harness. 

Based on the present results it is questionable whether the ATD numerical models can be used to 

assess the differences and effectiveness of different harness configurations reliably. 

3.2.4.2.1.3 Head and Neck Results

Head and neck results are important since the BDR model does not consider this region of the 

body, and injury criteria, such as those discussed in Section 3.1.2 are important as a supplement to 

the BDR to quantify the potential of injury to this region. During the initial transient response, the 

head X acceleration (Appendix C-1.3) follows the magnitude and shape of the driving seat 

acceleration. The acceleration magnitude was 1030 G and within the allowable 100-G peak HSIR 

limit. Near 200 msec, the head impacts the chest leading to a sharp peak in the head acceleration. 

For the three repeat tests, the measured peak acceleration was around 80 G, while all of the models 

predicted considerably higher peak accelerations. For the 20-G sled test (8205), the peak measured 

acceleration appeared to be clipped and missing the peak while the predicted peak accelerations 

far exceeded the allowable 100-G HSIR limit. Note that the S&G calculations are performed for 

the response before rebound only, so they do not include the discrepancies in the large peaks. 

Considering the large difference between test and prediction, it might not be reliable to use the 

numerical models for predictors of the peak-acceleration-injury criteria. Using acceleration as 

injury criteria may make sense from an occupant protection perspective, but it is impossible for 

the models to accurately predict peak accelerations. If the desire is to use numerical models for 

predicting injury metrics, then it is unwise to use peak accelerations as injury predictors. For Zaxis direction acceleration (Appendix C-1.5) none of the models predict the peak acceleration or 

the acceleration profile leading to the peak accurately.

It is useful to compare the head translation X velocity and the acceleration. Since velocity is an 

integral part of acceleration, it is expected, and verified by a comparison of the results, that the 
64

models would better correlate with velocity than acceleration (Appendix C-1.7). Before and during 

the head to chest impact, there was good correlation between the measured and predicted head 

velocity for all of the models. It is interesting to note the discrepancy between the results for the 

peak head acceleration and velocity injury criteria. More thought should be given to consider 

which of these criteria would be more useful for Orion-type applications. 

The head rotational acceleration (Appendix C-1.11) and velocity (Appendix C-1.12), and the head 

translation X acceleration and velocity follow similar trends in terms of matching test and 

predictions. While it is difficult to match test and predictions for the rotational acceleration, the 

rotational velocity match is more consistent. It appears there is a sign difference between test and 

predictions, but the head is moving forward and rotating downward toward the chest during the 

frontal impact tests. The test data show a positive head rotation, except for test 8202, where there 

was an unexplained sign change in the test data. 

HIC was calculated from test and predictions for HIC15 (15 msec, Appendix C-1.15) and HIC36 

(36 msec, Appendix C-1.16). Although both HIC are compared, only HIC15 is used in the HSIR 

requirements. For HIC36, the LSTC Hybrid III and the THUMS models did a good job of 

predicting HIC before the head impact with the chest (only 8202 was run with THUMS). Beyond 

impact, all of the models predicated different HIC values from the test data. None of the models 

(expect for THUMS) accurately predicted HIC15, and THUMS did not predict HIC15 accurately 

during the impact region. While the peak X acceleration exceeded the HSIR injury criteria for most 

of the cases, HIC36 was within the allowable range for every case, including the 20-G sled test 

(8205). 

For front impact, upper neck X shear force (Appendix C-1.18) and axial force (Appendix 

C-1.20) were compared to the test data. All of the numerical models (except the LSTC Hybrid III 

FEM) do a good job of predicting the neck X shear force to the point of head impact with the chest. 

The LSTC Hybrid III FEM follows the general trend of the neck X shear force. However, it 

predicts an oscillation in the neck force that is not seen in the test data or the other models. Around 

the time of impact, none of the models accurately predicts the shear force. All of the neck shear 

forces were within the HSIR allowable, including the 20-G test case. All of the models produced 

wide variations in the neck axial force when compared to the test data, although none produced 

axial forces that exceeded the HSIR limit except for the 20-G case (where all the models predicted 

a force exceeding the limit). Similar to the neck shear force, the LSTC Hybrid III FEM predicted 

a large oscillation in the neck axial force. This oscillation may be due to how the ATD model was 

constrained in the seat rather than to a problem with the ATD model. With additional refinement 

to the pre-loading and placement of the ATD in the seat, it is expected that much of this oscillation 

could be eliminated. 

It is important to note that the models over-predicted the measured test axial neck force. This is 

important because neck axial forces obtained from the ATD in Orion landing conditions have 

exceeded the HSIR occupant injury limits for several landing conditions. The limits are more prone 

to be exceeded when a helmet is placed on the occupant’s head due to the added weight of the 

helmet. Since the limits are sometimes exceeded and the models appear to be overestimating the 

axial neck force, it is important to examine this value and verify the predicted results with test 

measurements.

For upper neck flexion/extension moment (Appendix C-1.22), the MADYMO model correlates 

with the test data while the other models do a poor job of matching the test data. There is less of a 
65

problem with test and model correlation during the time of impact of the head and chest since 

moments are dependent on rotation, which is limited once the head impacts the chest. None of the 

flexion/extension moments exceeds the HSIR requirement. It may be that this requirement is never 

or rarely exceeded since the head rotation is always limited by the chest in flexion and the seat 

back in extension. The neck Nij is a combination of neck axial force and flexion/extension moment. 

Correlation between test and predictions for this injury metric is useful since this metric has 

potential as a valid measure of neck injuries for Orion. For all test cases, the MADYMO was within 

20 percent of the test Nij before head impact with the chest while the other models generated larger 

differences. In the region of head impact, none of the models produced reliable results.

3.2.4.2.1.4 Chest Response

Chest acceleration and displacement, and sternal displacements were predicted using the models, 

although only chest acceleration (and for a limited number of cases, body movement) was 

measured during testing. It was expected that the models would do a good job of predicting chest 

acceleration since the ATD is constrained to the seat with the harness and the ATD should typically 

follow the prescribed seat acceleration. As expected, all of the models track the chest 

X acceleration to the time of maximum forward excursion (Appendix C-1.26). The MADYMO 

model predicted the largest deviation from the test results, but was within 20 percent of the test 

data. The deviation can likely be attributed to modeling of the harness and not to the ATD 

numerical model. As mentioned previously, utilizing peak accelerations for the numerical models 

is unreliable since the models are unable to match peak accelerations with the measured test 

accelerations. 

None of the models were able to predict the peak Z acceleration (Appendix C-1.28) that occurs 

during the time of maximum forward excursion, or the acceleration profile leading to the peak. As 

seen in previous results, the LSTC Hybrid III FEM predicts large oscillations about the test data, 

which are attributed to how the ATD is held with the harness rather than to the ATD numerical 

model. It could be expected that the numerical ATDs would produce a worse match for Z-axis

direction response than for X-axis direction since the Hybrid III ATD is designed primarily for Xaxis direction loadings.

3.2.4.2.1.5 Lumbar Response

The models did not accurately predict the lower lumbar shear X (Appendix C-1.33). The 

MADYMO model consistently under-predicted the shear force while the rest of the models overpredicted the force. For all of the test cases, including the 20-G case, the X shear force was below 

the HSIR limit. For the lower lumbar axial force, the results were somewhat scattered; sometimes 

the models correlated with the test data and other times they did not. In general, the lumbar axial 

force was below the HSIR limit as would be expected for frontal impact loadings. The lower 

lumbar flexion/extension moments (Appendix C-1.35) were under-predicted by the MADYMO 

model and over-predicted by the other models. Pelvis X-direction acceleration was not measured 

with the physical Hybrid III. However, all of the models predicted similar acceleration profiles for 

this parameter and none of the predictions exceeded the HSIR limit.

3.2.4.2.1.6 Body Movement

During the sled tests, video data were obtained for each test and targets were placed on the ATD 

making it possible to perform video analysis to obtain body movements, including head movement. 

Head movement was computed for some of the test cases and when available was compared to the 

predicted data. It was not practical to post-process all of the test data to the obtained body 
66

movements. While there was large disagreement in the various models’ predicted forward head 

movement (Appendix C-1.13), the MADYMO model consistently predicted the correct forward 

head movement. Chest movement in the X-direction (Appendix C-1.30) predicted by the models 

varied from approximately 30 to 90 mm for the FTSS model for the 

10-G sled tests. For the 20-G sled test, the variation in predicted chest movement was of a similar 

magnitude. The physical Hybrid III ATD is capable of measuring sternal deflection. However, the 

ATD used for the WPAFB tests was not instrumented to record this value. 

3.2.4.2.2 Rear Impact

Rear-impact sled data were compared to the numerical models for a variety of rear loading 

conditions. Rear impact is one of the primary impact directions expected for Orion landings. 

However, the Hybrid III is not normally used for rear impact testing. Since Orion loadings are 

combinations of loading directions and it is desirable to use a single ATD for numerical predictions 

and physical testing, it was desirable to assess the test versus model correlation for this direction 

of loading. While it is important that the numerical model replicate the physical ATD for rear 

impact, additional work will be necessary to determine the applicability of the injury criteria and 

their limits for this direction of loading. Since the ATD is pushed into the seat during rear impact, 

it was expected that accurate modeling of the harness would be less important and many of the 

device internal loads would be small since it is constrained by the seat. 

3.2.4.2.2.1 Test Cases

Seven sled tests were run for rear impact using the Hybrid III automotive ATD (Appendix 

C-3.1). All tests were run near the 10-G level except for tests 8211 and 8212, which were run near 

20 G. The pulse duration was 70 msec unless noted otherwise in the test matrix. No repeat tests 

were planned for rear impact. However, some of the tests had faulty data channels so the tests were 

repeated to obtain complete data sets. All tests were performed with the identical seat except for 

test 8197 where the seat headrest was moved from the Hybrid III head, leaving a 

1-inch gap between the headrest and the back of the ATD head. For all other rear impact tests, the 

ATD head was placed against the headrest. For the 10-G sled tests, the loadings were relatively 

mild, indicated by the low BDR. For the 20-G test cases, the sled loading was amplified in the 

BDR and the peak dynamic response was close to 35 G, which is just at the limit for a low BDR 

score. It is interesting to note that while the sled peak acceleration was 20 G, the BDR to this 20-

G input was close to 35 G.

3.2.4.2.2.2 Harness Loads

For the left and right shoulder belt forces, none of the models reliably predicts the belt forces 

(Appendix C-3.38 and C-3.39). During the initial segment of the ATD response, the device is 

pushed into the seat, the belts are relaxed, and the forces are relatively small. During this segment, 

a large percentage difference in belt force is less significant since the magnitudes of the belt forces 

are small. Once rebound of the ATD with the seat occurs and the device bounces back, the belt 

forces increase and the mismatch between tests and prediction is more pronounced. During this 

segment, the numerical models did not reliably predict the belt forces, and the test data were not 

symmetric. For rear impact, where the seat loadings are symmetric, it was expected that the belt 

forces would be symmetric. Furthermore, the numerical models were symmetric and their 

predicted belt forces were not as symmetric as expected. Comparing the left and shoulder belt 

forces for test case 8194, the measured peak left and right belt forces are 

120 and 225 lbs, respectively, which is impossible for the numerical model to predict since the 
67

models themselves are symmetric. Correlations for the lap and crotch belts are just as poor as they 

are for the shoulder belts. As mentioned in the results for the frontal impact cases, assessment of 

the harness is an important factor of occupant protection for Orion and the inability of the 

numerical models to predict harness lading accurately is an important deficiency of the numerical 

models.

3.2.4.2.2.3 Head and Neck Results

All of the models perform well for predicting the head X acceleration peaks and shape (Appendix 

C-3.3). It is not surprising that the models reliably predict head X acceleration for rear impact, 

since the ATD head is pushed into the seat headrest and the device mostly follows the seat motion, 

which is relatively smooth and defined by the driving sled acceleration. For the Z-direction 

acceleration (Appendix C-3.5), all of the models reasonably predict the acceleration. The FTSS 

model predicts the largest discrepancy between test data and analysis results. However, it appears 

that the FTSS predictions are higher frequencies and there may be justification for filtering the 

results at a lower frequency to provide a better match with the test data. The Z-direction 

acceleration, resulting from rear impact, is small for this direction so a lack of correlation is far 

less important than for some of the other parameters.

It is useful to examine the head translation X velocity (Appendix C-3.7) and acceleration. Since 

velocity is an integral part of acceleration, it is expected the models would be able to predict 

velocity better than acceleration. Similarly to frontal impact, all of the models were able to predict 

head X velocity accurately for rear impact loading. Head movement (Appendix C-3.13) was not 

recorded for all of the models, but in general, the models correctly predicted the head was pushed 

into the headrest and rebounded. As expected, there was more of a difference among the models’ 

predictions during the rebound segment.

In general, none of the numerical models did a good job of predicting the HIC15 (Appendix 

C-3.16). The models did a better job of predicting the HIC36 (Appendix C-3.17), which was 

expected since the HIC36 was integrated over a longer period of time. The FTSS model did a good 

job of predicting the HIC36 for all but one case. For HIC15 and HIC36, the peak values were 

below the HSIR limit, which was expected since the ATD head is prevented from moving and 

protected by the seat headrest. As mentioned in Section 3.2.4.2.1, frontal impact, from a modeling 

perspective, the HIC36 may be a more useful occupant injury criterion since the models are more 

capable of reliably predicting the HIC36 than peak acceleration.

For rear impact, upper neck X shear force and axial force are compared to the test data (Appendix 

C-3.18 and C-3.20). None of the models did a good job of predicting the neck shear or the axial 

forces. These predictions may be a result of the Hybrid III physical ATD and the numerical model 

not being designed for rear impact loading. The measured maximum axial neck tension load was 

near 200 N for every case except 8197, where the measured peak neck tension was near 1000 N. 

Test case 8197 was a repeat of 8196 so it was surprising that the neck force was so large in 8197. 

The tensions were expected to be almost identical since they were repeat tests. The predicted neck 

axial forces were more in the range of 600–1000 N. It should be noted that, similarly to frontal 

impact, the neck axial force was always over-predicted by the numerical models. Caution must be 

used since the neck axial force limit in the HSIR is a criterion that is sometimes exceeded during 

analysis of Orion landing conditions. Since the models appear to over-predict axial neck tension, 

occupant injuries may be predicted when they are not occurring. Upper neck flexion/extension 
68

moment and neck Nij were not reliably predicted by the models and in general, the numerical 

models over-predicted the parameters.

3.2.4.2.2.4 Chest Response

Chest acceleration and displacement, and sternal displacements were predicted using the models, 

although only chest acceleration was measured during testing. It was expected that the models 

would do a good job of predicting chest acceleration, since during rear impact the ATD is pushed 

back into the seat and the ATD should follow the prescribed seat acceleration. As expected, all of 

the models track the chest X acceleration (Appendix C-3.26). None of the models, except for the 

THUMS model, is able to predict the Z acceleration (Appendix C-3.28) accurately. However, the 

magnitude of the Z acceleration is small and less significant for rear impact when compared to the 

X-direction of loading. 

3.2.4.2.2.5 Lumbar Response

Similar to frontal impact, all of the models (except MADYMO) accurately predicted the lower 

lumbar shear X force (Appendix C-3.33). For all the test cases, including the 20-G case, the 

X shear force was below the HSIR limit. None of the models consistently predicted the lower 

lumbar axial force (Appendix C-3.34). The sign of the lumbar axial force is opposite between test 

and predicted. The test predicted tension while the models predicted compression. It is expected 

that during rear impact the lumbar load would be compressive. The models did not reliably predict 

the lower lumbar flexion/extension moments. Pelvis X-direction acceleration was not measured 

with the physical Hybrid III. However, all of the models predicted similar acceleration profiles for 

this parameter and none of the predictions exceeded the HSIR limit.

3.2.4.2.3 Spinal Impact

Spinal impact sled data were compared to the numerical models for a variety of rear loading 

conditions. Spinal and rear impacts (discussed in Section 3.2.4.2.2) are two of the primary impact 

directions expected for Orion landings. Similar to rear impact, the Hybrid III is not normally used 

for spinal impact testing. Since Orion landings are combinations of loading directions and it is 

desirable to use a single ATD for numerical predictions and physical testing, it was advantageous 

to assess the test versus model correlation for this direction of loading. While it is important that 

the numerical model replicate the physical ATD for spinal impact, additional work will be required 

to determine the applicability of the injury criteria and their limits for this direction of loading. The 

BDR has been validated for spinal impact direction loadings so any spinal direction occupant 

injury criteria based on the ATD behavior should be compared to BDR results.

3.2.4.2.3.1 Test Cases

Nine sled tests were run for spinal impact using the Hybrid III automotive ATD (Appendix 

C-4.1). Tests were run at the 10- and 20-G levels and an additional test was run at 15 G. Various 

pulse durations were used and are shown in Table 3.2-1. Tests 8207, 8208, and 8209 were 10-G 

repeat tests. For frontal and rear impact, the 10-G sled test loadings were relatively mild indicated 

by the low BDR. For spinal impact, the same level of loading (10 G) produced a BDR that 

exceeded the low level of injury (Appendix C-4.2). The reason that the BDR criterion was 

exceeded for spinal impact is that the model allowable for spinal direction loading was lower than 

the allowable for rear or front impact. This is due to the human tolerance being considerably lower 

in the spinal direction. For the 10-G spinal loadings, the BDR was lower for the 110-msec rise 

time pulse than for the 70-msec pulse since the latter pulse was closer to the BDR model resonance 

frequency. The 110- and 70-msec rise times correspond to 2.3- and 3.6-Hz driving frequencies, 
69

respectively. The BDR model responds most strongly near 10 Hz so it is expected the 70-msec rise 

time pulses would lead to a higher BDR. The 15- and 20-G pulses produced a BDR that exceeded 

the low BDR limit regardless of the rise time.

3.2.4.2.3.2 Harness Loads

For the repeat tests, the left and right shoulder belt forces were repeatable and symmetric 

(Appendix C-4.38, C-4.39). The minimum shoulder belt tension was 250 N and the maximum is 

500 N, so while the belt forces were symmetric and repeatable, there was a spread of 250 N. None 

of the models predicted the belt forces reliably and there was less correlation between test and 

models for the lap and crotch belt than for the shoulder belts.

3.2.4.2.3.3 Head and Neck Results

Unlike frontal and rear direction loadings, peak head X accelerations are less important for spinal 

loading. Overall, the models do an adequate job of predicting the head X accelerations and as 

expected the X acceleration was relatively small, since the primary loading was in the 

Z-direction (Appendix C-4.3). For Z-direction acceleration (Appendix C-4.5), all of the models 

reasonably predict the acceleration. The LSTC Hybrid III FEM did not meet the ‘good’ correlation 

for three of the nine cases and the THUMS model was ‘fair’ for case 8208. For all of the test cases, 

the measured and predicted HIC15 was small (Appendix C-4.16) and below the occupant injury 

criterion of 300. Since the HIC was so small, discrepancies between test and predictions are not 

important. For front impact, upper neck X shear force and axial force are compared to the test data 

(Appendix C-4.18, C-4.20). None of the models did a good job of predicting the neck X shear or 

axial forces. The one exception was the FTSS model, which did a good job of predicting the neck 

axial force for all but one of the test cases. All the models over-predicted the forces. 

3.2.4.2.3.4 Chest Response

Chest Z acceleration is of primary importance for spinal direction loading (Appendix C-4.28). All 

of the models did a good job of predicting the chest Z acceleration. The one exception was the 

LSTC Hybrid III FEM, which exhibited large oscillations about the measured Z acceleration. It 

was expected that the models would do a good job of predicting chest Z acceleration since during 

spinal impact the ATD is pushed back into the seat pan and the ATD should follow the prescribed 

seat acceleration. As noted previously, although the models received a ‘good’ grade from the S&G 

correlation metric, there are peaks in the predicted response that do not correlate with the test data. 

When only the peaks were selected for the occupant injury criterion, there were larger 

disagreements between test results and the numerical model predictions. 

3.2.4.2.3.5 Lumbar Response

None of the models consistently predicted the lumbar shear X force (Appendix C-4.33). The FTSS 

was good for 4 of the 9 cases, but not good for the remainder of the test cases. For all of the test 

cases, including the 20-G case, the test and predicted X shear forces were below the HSIR limit. 

While lumbar axial force is not currently an HSIR occupant protection criterion, numerous 

discussions have been made concerning this criterion and its use for occupant protection. The BDR

criteria is thought to be a superior criteria for spinal direction loading and though the criterion does 

not address head and neck injuries, it does capture the effect of lumbar loadings. An important 

consideration for including lumbar axial forces as injury criteria is the models’ ability to predict 

reliably and to measure this force physically during testing. It was unexplained during this 

assessment why the test results showed the ATD lumbar force to be positive (tension lumbar force) 
70

in 2 of the 9 test cases when the ATD lumbar load was expected to be compressive for all nine 

cases. The models always predicted a compressive lumbar force. If the sign difference between 

test and predictions was ignored, then the MADYMO model generated a ‘good’ correlation for all 

test cases and the LSTC Hybrid III and FTSS models predictions ranged from poor to good, 

depending on the test case. 

The models consistently over-predicted the magnitude of the lumbar force, which has been the 

case for most, if not all, of the model predictions for all of the parameters examined in this study. 

The argument could be made that since the models tend to over-predict the injury criteria the 

reductions are conservative and the occupant will be protected. This argument may be valid for 

occupant protection. However, if the over-predicting leads to the injury criterion not being met, 

then the occupant protection will need to be re-designed or the landing loads reduced. This 

situation could lead to unnecessary design changes, added vehicle weight, and compromises to 

other vehicle functions (e.g., egress or occupant mobility). The models did not reliably predict the 

lower lumbar flexion/extension moments (Appendix C-4.35). Pelvis X-direction acceleration was 

not measured with the physical Hybrid III. However, all of the models predicted similar 

acceleration profiles for this parameter and none of the predictions exceeded the HSIR limit.

3.2.4.2.3.6 Comparison of ATD Lumbar Loads to Brinkley Z-Axis Response

As part of this study, there was discussion to add the ATD lumbar axial load as injury criteria. 

Considerable discussion has taken place concerning the applicability of the lumbar axial load limit 

used for other applications outside of Orion and how this limit might be applied or modified for 

Orion and the loadings expected for spacecraft landings. One issue with using these criteria is that 

it appears that the limits used in such applications as automotive or military have different loading 

conditions and limits may not be easily transferrable to Orion crew loading conditions. The FAA 

currently uses a 1500-lb limit for commercial aircraft seat qualification. However, this limit is 

associated with a higher tolerance for injury than is acceptable for Orion. Data for lower injury 

risks are not currently available.

Another issue associated with using the ATD lumbar axial load as injury criteria is the relationship 

between the lumbar criteria and the results generated by the BDR in the spinal (Z-axis) loading 

direction. The two criteria should be consistent since they address spinal injuries. There is concern 

that the two criteria are not consistent. The BDR model considers the effect of the loading pulse 

amplitude and shape, and the dynamic response of the human in the spinal loading direction. The 

ATD lumbar load limits currently available were derived for specific loading conditions, and may 

not be applicable for other loading conditions. Resolving this issue is beyond the scope of this 

assessment. However, a comparison between the ATD lumbar axial force and the axial BDR is 

provided in the following section.

Figure 3.2-24 shows the BDR limits for the high, moderate, and low levels of injury risk for the 

spinal direction. These limits were derived from human testing [ref. 37] and include the effect of 

the human body frequency response to pulse shape. For short pulses, the human body does not 

have time to respond and the allowable level of acceleration is therefore high. For long-duration 

pulses, the tolerable acceleration level is less than for short pulses; for pulses just below 10 Hz, 

the tolerance level is lowest since the human body peak response is near this frequency. 
71

Figure 3.2-24. Limits for +Z-axis Acceleration as a Function of the Time-to-Peak Acceleration for 

Half-sine Acceleration Profile

Figure 3.2-25 shows a comparison between various spinal direction input pulses and resulting 

ATD lumbar forces. A combination of pulses that produced low (DRz = 13.0 G) and moderate 

(DRz = 15.4 G) BDR injury levels was obtained from Figure 7.2-33, and those pulses were used 

as input to load the Hybrid III numerical model. The lumbar force was extracted from the Hybrid 

III and the results were used to generate the comparison shown in Figure 7.2-34. As mentioned, 

the BDR model was sensitive to how the human body responds to pulse magnitude and duration 

while the Hybrid III ATD is thought to be less human-like in this direction and probably not a 

good predictor of spinal loading for a range of pulse durations. As shown in Figure 3.2-25, the 

peak ATD lumbar force varies depending on pulse magnitudes that generate identical BDRs. There 

were inconsistencies between the ATD lumbar load predictions and the BDR predicted injury 

levels. 

Figure 3.2-25. Comparison of +Z BDR Low and Moderate to ATD Lumbar Forces

high

moderate

low
72

3.2.4.2.4 Lateral Impact

Lateral impact sled data were compared to the numerical models for a variety of loading 

conditions. The Hybrid III is not normally used for lateral impact direction loadings so it is 

important to assess the numerical models’ abilities to predict the ATD response in this direction. 

Furthermore, if the models are able to predict Hybrid III lateral response, then the applicability of 

the injury criterion developed for this direction loading to the Hybrid III ATD will need to be 

assessed. This issue is discussed in detail in Section 3.2.4.2.5 where the response of the

Hybrid III is compared to the EuroSID ATD.

3.2.4.2.4.1 Test Cases

Fourteen sled tests were run for lateral impact using the Hybrid III ATD (Appendix C-2.1). For 

lateral testing, seat side supports were added to limit ATD lateral motion. Test 8239 was run 

without side supports, but there were concerns that the ATD would be damaged since it could not 

be constrained by the harness. A combination of 10- and 20-G load cases was run with a pulse 

length of 70 msec. A constant pulse length was used for the lateral testing to eliminate the time 

required to convert the sled-driving mechanism to the 110-msec pulse setup. Three-sided support 

configurations were used. A no gap configuration was used where the side supports the head, 

shoulder, and pelvis/leg region were placed directly against the ATD to provide full lateral support 

with no motion permitted between the ATD and the side supports. The second configuration was 

a 50-percent gap where the distance was 50 percent of the HSIR body movement requirements. 

The head, shoulder, and pelvis/leg side supports have difference gaps based on these body 

movement requirements. The final configuration tested had a gap equal to 100 percent of the HSIR 

body movement requirements. The ATD injury parameters derived from these tests did not provide 

direct insight into the effectiveness of the body movement requirements since the test loadings are 

different from spacecraft landing loads, but the results did provide a sense of the relationship 

between restricted body movement and resulting injury criterion.

Human tolerance to impact loading is far lower in the lateral direction than for the rear, frontal, 

and spinal directions. This lower tolerance level is seen in the BDR where the 20-G sled tests 

exceed the BDR low level of injury (Appendix C-2.2). A number of the 10-G impact tests exceed 

the low BDR limit, although the results were at the low limit. The remainders of the 

10-G test cases do not exceed the BDR low level. The BDR model was obtained for a specific seat 

design with side supports so the results obtained were not accurate since this seat differs from the 

seat used to develop the BDR model. The only way to remedy this complication is to redevelop 

the BDR model using the same seat that was used for the test, or in the case of a space vehicle, 

redevelop the BDR model using the spacecraft seat.

3.2.4.2.4.2 Harness Loads

Predicting harness loads is more challenging for lateral direction impacts than for the other 

directions. For the other directions, the loading is essentially symmetric and the loading in the belts 

is primarily tensile, providing less of a challenge for the models’ predictive capabilities. For side 

impact, the ATD belts are loaded in a complex manner and it is more difficult to predict the loads 

with accuracy. Additionally, for side impact, the seat side supports provide most of the lateral 

constraint and the harness is less important than for the other loading directions 

(i.e., where the harness provides the only constraint). 
73

3.2.4.2.4.3 Head and Neck Results

For lateral impact, the head and neck do not only move in the lateral direction. Instead, the head 

and neck motion is complex and the total loading is a combination of loadings in all three 

translational and rotational directions. As a result, it is important to consider other directions in 

addition to the lateral direction-oriented injury criteria. It is important to note that the injury criteria 

used in the HSIR were derived primarily for single-direction impacts. For combined loading, it is 

not clear how to combine injury criteria. At this time, the HSIR requirements for head and neck 

load limits treat each loading direction as though they act independent from each other. 

None of the models were able to predict the peak head resultant acceleration reliably (Appendix 

C-2.6). This result was not surprising since the ATD head impacts the side support, and for many 

of the test cases, it repeatedly impacts the side support causing multiple peaks in the acceleration 

profile. It is difficult to predict these peak-impact loads numerically, leading to the question of 

whether it is practical to utilize peak head resultant acceleration as an occupant injury criterion. 

Even though the HIC is computed from an integral of the acceleration and tends to smooth out 

discrepancies in peak accelerations, the HIC15 was not well predicted by the models since the 

discrepancies in the acceleration predictions were too large for the smoothing to have much effect 

on the correlations (Appendix C-2.16).

The correlation between test and predictions for the neck force in the X, Y, and Z directions ranged 

from “good” to “poor” for all of the models. The lateral Y force (Appendix C-2.19) was less 

reliable than the X-direction (Appendix C-2.18) predictions, which were expected since the ATD 

is not designed for loading in this direction, and the models were not expected to perform well in 

this direction as for frontal impact. The neck Nij and the neck lateral moment 

(Appendix C-2.22) were not well predicted by the models.

3.2.4.2.4.4 Chest Response

All the models did a “good” to “fair” job predicting chest Y acceleration (Appendix C-2.27). There 

was no trend between the level of load or the distance between the ATD and side supports. It was 

expected that the impact resulting from the gap between the seat side supports and ATD would 

challenge the models’ ability to predict the response accurately. However, the models generated 

reasonably correlated acceleration predictions. 

3.2.4.2.5 EuroSID Lateral Test Results

An issue related to the above approach is that the Hybrid III is not normally used or validated for 

lateral impact. While Orion landings generally produce primarily rear- and spinal-direction 

loading, there are landing conditions (particularly when vehicle roll is prevalent) where lateral 

loading occurs. Furthermore, the HSIR requirements include lateral direction-related criteria that 

must be satisfied. The EuroSID may be used for pure lateral loading. However, Orion loading is 

multi-axial and the EuroSID would not be amenable for use in the above approach. To date, NASA 

has no procedure for separating the multi-axial landing loads into individual directions so that 

direction-specific dummies (the EuroSID) may be used for assessing crew injuries during landings.

The objective of the test plan was to compare the Hybrid III lateral response to the EuroSID 

response with particular emphasis on comparing the HSIR lateral loading criteria between the 

Hybrid III and EuroSID for identical seats, constraints, and loading conditions. If it is determined 

that the Hybrid III produces similar results to the EuroSID for lateral loading, then the Hybrid III 

can be used for the approach involving multi-axial landing simulations and data extraction. If the 
74

Hybrid III does not produce similar results to the EuroSID, then the above planned approach may 

not work and an alternate approach will need to be developed.

During sled testing at WPAFB, additional testing was completed with the EuroSID2 for 

comparison to the Hybrid III lateral response (see Table 3.2-9). The EuroSID2re was tested in the 

same conditions as the Hybrid III with 10- and 20-G pulses, and with no gap and a 

50-percent gap between the ATD and the side supports (seeTable 3.2-10). The EuroSID2 was 

designed specifically for side impacts and was useful to compare how the Hybrid III compares to 

this ATD for side impacts. Comparisons were made for head, shoulder, pelvis, thigh, and knee 

lateral forces. Comparisons could not be made between internal ATD forces since the EuroSID 

was not instrumented in the same manner as the Hybrid III. Overall, the Hybrid III and EuroSID 

produced similar lateral forces. Since internal forces were not compared, and it is the internal 

forces that are used as injury criteria, it was not possible to draw a conclusion as to whether the 

Hybrid III can be used in place of the EuroSID for side impacts. 
75

Table 3.2-9. Lateral Impact Test Matrix
76

Table 3.2-10. EuroSID Testing Matrix

G Level Sled Y Gap (percent) DRy Case ATD

10 10.7 0 14.4 8285 EuroSID2re

10 10.3 0 14.3 8286 EuroSID2re

10 10.8 17 14.2 8287 EuroSID2re

10 10.5 17 14.2 8288 EuroSID2re

10 11.6 50 14.3 8289 EuroSID2re

10 12.1 50 14.3 8290 EuroSID2re

20 20.6 0 31.2 8291 EuroSID2re

20 21.3 0 31.8 8292 EuroSID2re

20 23.1 50 31.5 8293 EuroSID2re

10 10.8 0 14.8 8242 Hybrid III

10 10.5 0 14.2 8294 Hybrid III

10 11.5 50 14.9 8244 Hybrid III

10 11.7 50 14.8 8245 Hybrid III

10 11.2 50 14.3 8295 Hybrid III

10 13.5 100 15.0 8248 Hybrid III

20 22.0 0 32.6 8243 Hybrid III

20 21.9 0 31.9 8296 Hybrid III

20 22.9 50 32.6 8246 Hybrid III

20 23.7 50 32.3 8247 Hybrid III

20 27.0 100 32.2 8249 Hybrid III

3.2.4.2.5.1 Headrest Lateral Force

During testing, the ATD head impacts the headrest with different magnitudes. Figure 3.2-26 shows 

the relationship of head lateral forces between the EuroSID and Hybrid III dummies.

Figure 3.2-26. EuroSID/Hybrid III Headrest Lateral Force Comparison

y = 860.22e0.0004x

R2 = 0.9965

0

1000

2000

3000

4000

5000

6000

7000

8000

0 2000 4000 6000 8000

Hybrid-III Headrest Y Force [N]

EuroSID2re Headrest Y Force [N] 

10G, 0% Gap

10G, 50% Gap

20G, 0% Gap

20G, 50% Gap
77

3.2.4.2.5.2 Shoulder Lateral Force

Figure 3.2-27 shows the relationship of shoulder lateral forces between the EuroSID and 

Hybrid III dummies.

Figure 3.2-27. EuroSID/Hybrid III Shoulder Support Lateral Force Comparison

3.2.4.2.5.3 Pelvis Lateral Force

Figure 3.2-28 shows the relationship of pelvis lateral forces between the EuroSID and Hybrid III 

dummies.

Figure 3.2-28. EuroSID/Hybrid III Pelvis Support Lateral Force Comparison

y = 2975.7ln(x) - 20951

R2 = 0.9842

0

1000

2000

3000

4000

5000

6000

7000

0 2000 4000 6000 8000 10000

Hybrid-III Thigh Y Force [N]

EuroSID2re Thigh Y Force [N] 

10G, 0% Gap

10G, 50% Gap

20G, 0% Gap

20G, 50% Gap

y = 2975.7ln(x) - 20951

R2 = 0.9842

0

1000

2000

3000

4000

5000

6000

7000

0 2000 4000 6000 8000 10000

Hybrid-III Thigh Y Force [N]

EuroSID2re Thigh Y Force [N] 

10G, 0% Gap

10G, 50% Gap

20G, 0% Gap

20G, 50% Gap
78

3.2.4.2.5.4 Thigh Lateral Force

Figure 3.2-29 shows the relationship of thigh lateral forces between the EuroSID and Hybrid III 

dummies.

Figure 3.2-29. EuroSID/Hybrid III Thigh Support Lateral Force Comparison

3.2.4.2.5.5 Knee Lateral Force

Figure 3.2-30 shows the relationship of knee lateral forces between the EuroSID and Hybrid III 

dummies.

Figure 3.2-30. EuroSID/Hybrid III Knee Support Lateral Force Comparison

y = 2975.7ln(x) - 20951

R2 = 0.9842

0

1000

2000

3000

4000

5000

6000

7000

0 2000 4000 6000 8000 10000

Hybrid-III Thigh Y Force [N]

EuroSID2re Thigh Y Force [N] 

10G, 0% Gap

10G, 50% Gap

20G, 0% Gap

20G, 50% Gap

y = 2382.6ln(x) - 14215

R2 = 0.8715

0

2000

4000

6000

8000

10000

12000

14000

0 2000 4000 6000 8000 10000 12000 14000

Hybrid-III Knee Y Force [N]

EuroSID2re Knee Y Force [N] 

10G, 0% Gap

10G, 50% Gap

20G, 0% Gap

20G, 50% Gap
79

Conclusions and Recommendations

The sled test data collected in the AFRL HIA facility provides one of the most extensive databases 

for the Hybrid III ATD, including data for frontal, rearward, spinal, and lateral 

loading conditions. Detailed transient response data are provided in Appendix C. Appendix B 

provides a summary of the Sprague and Geers analysis obtained from the transient results, and 

Appendix D provides a comparison between the measured and predicted injury metrics. These data 

were used to assess four numerical models and their abilities to predict the ATD response for the 

four loading directions:

 LSTC Hybrid III

 FTSS Hybrid III

 MADYMO Hybrid III

 THUMS model

In general, none of the four numerical models generated results that correlated with the measured 

test data to within the desired 20-percent S&G magnitude criteria for the selected ATD response 

criteria, including the specific injury criteria specified in the HSIR document. This result is 

concerning since the assessment team would like to use numerical models to replace physical 

testing, especially for many of the vehicle-landing conditions where physical testing is not 

possible.

Overall, the FTSS Hybrid III model performed the best among the four numerical models, 

exhibiting the least deviation between numerical prediction and test results. This result was 

expected since the FTSS model has the longest development history and a long record of successful 

use for a wide variety of applications in the aerospace and automotive industries. The MADYMO 

model had less modeling capabilities than the FTSS and LSTC Hybrid III models since these 

models are FE-based, while the MADYMO ATD model consists of rigid links and the remainder 

of the structure was modeled with simple contact. The LSTC Hybrid III FEM is undergoing 

validation and improvements. The THUMS model is designed to replicate the human body and is 

not intended to directly replicate the Hybrid III ATD.

There are several possible causes of the lack of correlation between the models and the test data. 

First, there is test data variability. The repeat tests show this variability, and it is particularly 

evident for test cases where the results were expected to be symmetric and the test data were nonsymmetric. This non-symmetry is the result of operational test conditions, such as response’s 

sensitivity to initial conditions (e.g., ATD, belt positioning, and the nature of the test event). 

The second cause of limited correlation is in the numerical models. While all of the models have 

been validated for specific loading conditions, the validation tests tend to be specific to ATD body 

regions (e.g., isolated ATD pendulum head drop to validate the head model) and do not capture 

the full interactions of the entire ATD in a realistic seated environment. The present tests are more 

complicated than the ATD validation tests so more discrepancies are expected with the computed 

response. Furthermore, the ATD numerical model was placed in a seat with a harness, adding 

additional uncertainties and challenges to the modeling effort. 

Finally, there is the expertise of the individual modelers. For this study, independent modelers 

were used for each of the ATD numerical models. All of the modelers had extensive experience 

with ATD numerical modeling and come from industry and academic backgrounds. Even with this 
80

expertise, it was apparent a significant cause of the difference between test and predictions was 

from the approach and techniques used by the modelers. While each modeler used identical sleddriving accelerations obtained from the test data, each modeler made their own assumptions in 

terms of ATD positioning, harness and seat characterizations, and preloads. All of these factors 

have been found to affect the ATD numerical response and are a cause of discrepancy between 

test and predictions. 

For the present testing, the modelers had an opportunity to refine their models based on reviewing 

the test data and ATD response. Pre-test predictions before the models were refined and did not 

correlate with the test data. As a result, in situations where the models are to be used as a purely 

predictive tool, test data will not be available, which will lead to more uncertainty in the numerical 

predictions. The FAA requires numerical ATD models that will be used in place of testing to be 

validated for a physical environment (e.g., seat, constraints) and loading conditions similar to the 

test conditions that the model is to replace. 

Before a numerical model is used to perform a final assessment of occupant injury, it is 

recommended that it be validated through physical testing. This testing should be conducted in a 

similar physical environment, including similar or identical seat, harness, helmet, and suit, and 

additional physical hardware components that interact with the occupant. Testing should be 

performed with various size ATDs. In addition, loading conditions for these physical modeling 

validation tests should be similar to the loading conditions that will ultimately be used for the 

occupant protection assessments. 

3.3 BDR Model Refinements Summary

Introduction and Background

The primary objectives of this research effort were: 

1. To develop biomechanical models to simulate the dynamic inertial response of the human 

body to Y-axis impact. These models were intended for the design and analysis of the 

performance of the Orion crew seat landing impact attenuation system. The dynamic 

models describe the whole body response and the distribution of forces to the side body 

support structures of the crew seat, especially the shoulder and hip supports.

2. To improve a biomechanical model(s) to evaluate the safety of the Orion crew seat 

occupants with restraint harnesses and seats with side body supports. The model is 

intended to provide insights regarding the causes of injuries to the seat occupant’s 

shoulder that have been observed during high-speed racecar sideward impact. The model 

was intended to update the BDR model and exposure limits for Y-axis.

The first objective to simulate the dynamic inertial response of the human body to Y-axis impact 

was developed as a result of difficulties observed in simulations of the performance of the Orion 

impact attenuation system. The simulation included FEMs of the impact attenuators, seat support 

pallet, crew seats, and rigid masses representing the seat occupants. The CM impact caused the 

simulated assembly to resonate at high frequency with no evidence of significant damping. 

Furthermore, the impact attenuator stroke limit was exceeded at the extreme landing conditions.

The second objective was to develop a dynamic response model (DRy) based upon sideward tests 

with volunteer subjects, where rigid side panels supported the subjects. The original DRy natural 

frequency and damping coefficient were estimated on the basis of sideward impact tests where 

several frequencies were extracted using measurements of whole body force and impact velocity 

and an impedance computational method [ref. 21]. The natural frequency and damping coefficient 
81

ratio was determined from tests without rigid side supports. The study described was intended to 

improve the BDR in the lateral, Y-direction by using data from an existing experimental study. 

While not simulating the side supports provided by the Orion crew seat, the data could be used to 

more accurately define the whole body, upper body, and lower body dynamic responses. 

Based on prior experience with simulations of impact attenuator performance, a study was initiated 

to meet the first objective using data from impact tests with volunteers to develop simple lumpedparameter models of human impact inertial response to evaluate their influence on these 

simulations. 

To evaluate the feasibility of developing such a model, Abbott [ref. 80] selected data from impact 

tests of several mid-size subjects from AFRL’s Collaborative Biomechanics Data Network and 

evaluated the dynamic responses of the subjects using several linear analysis techniques, including 

a frequency domain transfer function analysis to identify the primary whole body natural frequency 

and associated damping coefficient ratio. From these analytical efforts (Figure 3.3-1), Abbott was 

able to develop simple models to represent the human body response for the +X-and +Z-axes. 

Figure 3.3-1 illustrates the results of Abbott’s transfer function analysis of forces and accelerations 

measured during a +Z-axis impact testing. The model Abbott developed to duplicate the 

impedance computed from the tests data consisted of two masses connected by a spring and 

damper. Abbott’s initial efforts to use the models demonstrated that the models eliminated the 

high-frequency ringing encountered in earlier simulations and demonstrated a potential to reduce 

impact attenuator stroke distance.

Figure 3.3-1. The Impedance Properties of a Volunteer Subject as a Function of Frequency 

Using a similar frequency domain method to compute the transmissibility between the 

measurement of acceleration at the seat and the forces that were exerted by the subject as a result 

of the impact, existing data from Y-axis impact tests were analyzed to address the identified 
82

objectives. Because of injuries to the shoulder complex observed as a result of NASCAR racecar 

sideward impacts, the effort was directed toward the establishment of a model that would be more 

representative of the forces acting through the test subject’s shoulder. The injuries incurred were 

clavicular fractures and dislocations and a fracture of a driver’s scapula. Although adverse 

cardiovascular responses had been identified as a potential injury mode in earlier experiments 

during the Apollo Program, these responses had been considered to be of no more than moderate 

risk and more recently have been proven to be minimized by preconditioning the crew prior to 

reentry [ref. 81]. Fracture or dislocation of the clavicle would prevent the crew from exiting the 

semi-supine crew seat and its restraint and disconnecting the pressure suit from the life support 

system without assistance. Medical attention would be required.

A significant number of concussions were noted during the analysis of the NASCAR racing 

accidents. Some of these accidents might only have resulted in stunning and a short period of 

disorientation, while other cases caused loss of consciousness (LOC). The injury predominately 

occurred during accidents where the highest impact occurred in the sideward direction. Only one 

concussion occurred from a rearward impact. This finding is consistent with clinical and 

experimental observations that sideward impact causes traumatic brain injury at lower impact 

levels than anterior-to-posterior or posterior-to-anterior impacts [ref. 82]. 

Current practice in the automotive industry is to use the HIC to evaluate the likelihood of head 

injury. Unfortunately, the HIC is intended to evaluate the skull fracture risk [ref. 83]. Although 

concussion may occur with skull fracture, the threshold for concussion is lower than the HIC would 

predict. Newman, Shewchenko, and Welbourne [ref. 84] used a laboratory-based reconstruction 

of football-helmeted head impacts using Hybrid III ATDs. The primary focus was to re-enact headto-head collisions analyzed from game videos. The data were divided into injured and non-injured 

players. The HIC15 function for 24 cases of mild traumatic brain injury cases was computed using 

the resultant acceleration of the ATD head where the injury set was dominated by sideward head 

impacts. The corresponding injury risk curve, computed using a logistic regression, is shown in 

Figure 3.3-2.

The consideration of two modes of injury, concussion and injury of the shoulder complex, presents 

an issue that can be addressed using the BDR criteria. However, it is unknown if the BDR 

acceleration injury risk curves for the two injury modes will overlap. This issue could not be 

adequately addressed due to the cancellation of CxP and the Task 1 activities.
83

Figure 3.3-2. Probability of Concussion Based on the HIC

Methods

The impact test data that were used in the analysis were collected using a horizontal impact facility 

located within AFRL (Shaffer, August 1976) [ref. 85]. The tests were conducted to meet the joint 

research objectives of the USAF and the NHTSA that included:

 Comparison of ATD impact responses to test subjects.

 Measurement of the linear and angular motions of the head and neck using a head 

mounted nine-accelerometer array.

 Measurement of the supporting panel sideward impact forces exerted by the test 

subjects.

The tests were conducted at 3, 4.5, and 6 G; impact velocities were 8.6, 11.7, and 14 ft/sec; and 

pulse durations were 133, 118, and 110 msec, respectively. An example 6-G acceleration-time 

history is shown in Figure 3.3-3.
84

Figure 3.3-3. The Acceleration-time History at the 6-G Impact Level

Volunteer subjects were military officers and airmen. The subjects included 13 males and 

one female. Each subject was given a physical that included a radiographic examination of the 

cervical spine. Informed consent of all volunteers was obtained in accordance with USAF 

Regulation 169-3 [ref. 86].

A special-purpose test apparatus was designed to support and restrain the subject’s body and 

extremities. Extra emphasis was given to ensuring that the frequency response of the test apparatus 

structure was above the anticipated measurements of the dummies and volunteer subjects. Impact 

tests were conducted to evaluate the structural adequacy of the test apparatus and to measure its 

frequency responses. The frequencies of the observed structural responses were 200 Hz or higher. 

The body support and restraint system of the apparatus is shown in Figure 3.3-4 and Figure 3.3-5. 

Two instrumented side support panels were used. One supported the subject’s upper body and 

extremities and the second supported the subject’s lower body and legs. The dividing point 

between the two supports was at the iliac crest of the subject’s pelvis to approximate the location 

of the test subject’s center of gravity. Three StrainsertTM FLU1-25G force cells supported each 

panel. The upper body support panel force cells were positioned to support the shoulder, elbow, 

and wrist. The lower body force cells were positioned to support the hip, knee, and ankle. It is 

important to note that there was no gap between the subject and either of the side supports. This 

feature is important since the resulting injury criteria were based on test data collected for a 

configuration without gaps between the subject and side supports. For configurations with gaps, 

the resulting injury criteria will most likely be different.
85

Figure 3.3-4. Sideward Impact Test Apparatus with Volunteer Subject

Figure 3.3-5. Side View of the Test Apparatus with a 95th-percentile ATD Subject

The force cells were located within a seat coordinate system. The origin (0) of this system 

was located at the intersection of the seat pan and seat back and centered at the seat back, 

sometimes referred to as the seat reference point. The force cell X-Z plane locations are defined in 
86

Figure 3.3-6. The upper body panel was located 13.63 inches from the origin and the lower body 

panel was located 10.63 inches from the origin on the Y-axis. 

The weight of the upper panel acting on the force cells was determined to be 6.6 lbs. The weight 

of the lower panel acting on the force cells was 16.8 lbs. During data processing, the force cell 

measurements were corrected for the tare weight of the panels. 

The force and acceleration measurements were filtered using SAE class 60 filters with a 120-Hz 

corner frequency.

Figure 3.3-6. Positions of the Centers of the Force Cells are Defined with Respect to the Coordinate 

System Origin

Multiple tests were conducted using rigid 50-lb weights mounted to the upper and lower panels to 

evaluate the adequacy of the force-measuring system. The variance between the measured force 

cell outputs under acceleration was less than 0.5 lbf or 1 percent.

Triaxial linear acceleration and velocity of the impact sled were measured. Triaxial accelerometers 

were chest mounted. In some cases, the chest accelerations were not found to be reliable due to 

inadequate mounting and these accelerations were not used for analysis. 

The first of the test program’s original objectives was accomplished using a nine-accelerometer 

array attached to the test subject’s and ATD’s heads. 

Data from 72 tests with volunteers were reprocessed from the original data files to use the tarecorrected upper body Y-axis forces, lower body Y-axis forces, and total body Y-axis forces.

Photometric reference targets were 1-inch-diameter quadrant-style targets mounted on the test 

subjects and the test apparatus. Four high-speed motion picture cameras mounted on the test 

apparatus and one on the side of the track recorded the movements of the subjects and the targets. 

The motion picture films were converted to video recordings. The application Visual FusionTM
87

was used to automatically track the positions of selected targets as a function of time. The positions 

of the targets were determined to be accurate to within 0.2 inches. In a few cases where the target 

contrast was not always adequate due to overexposure of the film or where the image of the target 

was momentarily lost, operator intervention was required. 

Since they were of specific interest to this study, the displacement of the targets mounted over the 

manubrium and xiphoid processes of each subject was tracked as a function of time. Due to 

overexposure of the photographic images during the majority of the tests at 3 G and 4.5 G, these 

data were inadequate for analysis.

The targets mounted over the manubrium and video targets for 44 6-G tests, 5 4.5-G tests, and 

1 3-G test were tracked to determine the maximum Y-axis manubrium displacement. The target 

over the manubrium was sometimes overexposed or in the shadow of the head and neck, which 

made it difficult to track it automatically in some cases. The anthropometric chest breadth 

measurement was used as the scaling factor for converting the video pixel to dimensional 

displacement. There was some variation is the computed scaling factors for each test so an average 

scaling factor was computed based on all of the 6-G tests. The maximum neck displacements from 

all of the tests were plotted as a function of the sled acceleration level. 

Preliminary analysis was done using data from tests numbers 1194 and 1195 to determine the 

method to define model parameters to simulate the force-time histories measured by the force cells 

mounted to the two sideward support panels. The time domain and frequency domain methods 

were evaluated.

The impedance magnitude, impedance phase, and transmissibility magnitude were computed using 

the upper, lower, and total body Y-axis forces and the velocity as a function of time 

[ref. 87]. The natural frequency and damping ratio of the physical responses were calculated using 

the impedance analysis results. 

The natural frequency and damping ratio of the upper body force, lower body force, total body 

forces, chest acceleration, and head acceleration were calculated for the two tests using a time 

domain approach. The LoadLsq computer program was used to conduct this analysis. The LoadLsq 

program uses a technique described by [ref. 88]. In this case, the program computed the natural 

frequency and damping coefficient ratio of a lumped parameter model that best fits the measured 

data by minimizing the least squared difference between impact parameters that are computed 

from the forces and the same parameters computed from the model. The four impact parameters 

used are the peak force level, impact start time, start time plus rise time, and end time. The program 

has the option to filter the data before finding the results. The natural frequency and damping 

coefficient ratio of the upper body, lower body, and whole body were computed for all 72 tests. A 

50-Hz low pass filter was used except in a few cases where the signal was too irregular to obtain 

a good fit. In those cases, a 30-Hz low pass filter was used. Scatter plots with trend lines were 

created for the hip and shoulder forces as a function of the impact acceleration and the subject’s 

weight, and for the hip and shoulder forces normalized by subject weight as a function of impact 

acceleration. 

After a number of tests had been processed, it was determined that the time domain method was 

analyzing shoulder impact responses that were a composite of two frequencies; one of about 

4 Hz and the other about 9.5 Hz. Analysis of the measured head accelerations indicated that the 

head and neck response was responsible for the lower frequency response and the shoulder 

complex was more likely to be the source of the higher frequency response. Additional 
88

confirmation was obtained by study of the measured acceleration of the subject’s chest. The 

analysis revealed that the response of the chest reflected a natural frequency of approximately 

7 Hz. Therefore, the frequency domain method was selected since it was capable of defining the 

defining the amplitude of the responses at each frequency, and most importantly, identifying the 

model parameters for the model of the shoulder response. 

Results

Figure 3.3-7 shows typical upper body forces and sled acceleration measured during the tests at 

the 6-G level. Note that the forces measured at the shoulder are dominant. The same relationship 

was observed in the measurements of the lower body forces. The forces measured at the subject’s 

hip were larger than the forces measured at the knee and foot.

Figure 3.3-7. Sled Acceleration and Forces Measured by the Upper Body Support Panel

Scatter plots with trend lines were created for the hip and shoulder force as a function of impact 

acceleration and the subject weight, and for the hip and shoulder forces normalized by subject 

weight as a function of the impact acceleration. These plots are provided in Figure 3.3-8 and Figure 

3.3-9.
89

Figure 3.3-8. Y-Axis Force Measured at Each Subject’s Hip as a Function of Sled Acceleration

Figure 3.3-9. Y-Axis Force at Each Subject’s Shoulder for Each Sled Acceleration Condition

Figure 3.3-10 shows the displacement of each subject’s manubrium in the Y-axis. Unfortunately, 

due to over exposure of the motion picture films at most of the lower impact levels, these data 

were not available. The trend line that has been calculated is misleading. Its slope should lead to 

zero Y-axis displacement. At the 6-G level, the mean displacement of the manubrium was 2.36 

inches (60.7 mm), standard deviation (std dev) = 0.37 inch (9.40 mm). 
90

Figure 3.3-10. Maximum Displacement of the Target over Each Subject’s Manubrium for Each Sled 

Acceleration Condition

The mean force measured at the force cell at the subjects’ shoulder was 520 lbs (236 kg), 

std dev = 64.3 lbs (29.2 kg). The mean force measured at the elbow was 112 lbs (51 kg), 

std dev = 46 lbs (21.0 kg). The mean force measured at the wrist was 58 lbs (26 kg) with a standard 

deviation of 15.3 lbs (6.94). The mean of the total upper body force measured was 

622 lbs (282 kg), std dev = 63.9 lbs (29.0 kg) n = 45.

The mean force measured at the hip was 520 lb (236 kg), std dev = 70.7 lbs (32.0 kg). The mean 

force measured at the knee was 354 lbs (161 kg), std dev = 69.3 (31.4 kg). The mean force 

measured at the subjects’ ankle was 153 lbs (63.4 kg), std dev = 27.9 lbs (12.7 kg). The mean 

lower body force was 880 lbs (399 kg), std dev = 108 lbs (49.0 kg) n = 45.

The mean of the total whole body force was 1,430 lbs (649 kg), std dev = 162 lbs (73.5 kg) 

n = 45.

The upper body dynamic response model mean natural frequency was 5.13 Hz, std dev = 1.12. 

The mean damping coefficient was 0.50, std dev = 0.16. The mean natural frequency of the lower 

body dynamic model was 10.9 Hz, std dev = 1.86. The mean damping coefficient was 0.07, std 

dev = 0.07. The mean natural frequency of the whole body dynamic response model was 8.0 Hz, 

std dev = 1.26. The mean damping coefficient ratio was 0.31, std dev = 0.10.

The dynamic response model parameters for the shoulder force were calculated for each impact 

level as shown in Table 3.3-1. 

Table 3.3-1. Dynamic Response Parameters for the Shoulder Impact Model

Impact Level (G) Natural Frequency (Hz) Damping Coefficient Ratio

3 9.8, std dev = 0.39 0.37, std dev = 0.08

4.5 11.5, std dev = 0.38 0.21, std dev = 0.06

6 12.9, std dev = 0.31 0.18, std dev = 0.05
91

Model stiffening is reflected in the increasing values of natural frequency as a function of 

G level. At these relatively low impact levels, this was not an unusual observation. The reduction 

in the damping coefficient is to be expected as the muscle tension of the subjects was overcome.

Discussion

Since tests with volunteers had not been planned as part of the technical efforts described within 

this report, the effort documented was limited to the use of data from an existing experimental 

study conducted by AFRL for a different purpose. However, the data were considered adequate to 

demonstrate the feasibility of meeting the two objectives of this technical effort. Furthermore, 

some of the findings may be used in the design of future crew seats and to analyze the interactions 

between an impact attenuation system, a crew seat, and the human body.

The notable differences between the test apparatus used in the experimental effort conducted by 

Killian and Boedecker [ref. 89] and potential designs of future crew seats and restraints include 

the use of full body support panels rather than sideward supports that may only support segment 

of the seat occupant’s body, such as the helmeted head, shoulder, knees, and feet. The restraint 

system used in the experimental apparatus is different from a restraint that might be used in a future 

spacecraft crew seat, and could influence the results of this assessment.

Sideward impact studies that have been conducted to define impact exposure limits for the 

shoulder have been performed using PMHS. The tests have been performed using impact sleds to 

produce whole body impacts or by pendulum or pneumatically actuated rams that strike the 

shoulder area. Data from these tests have been used to specify impact exposure limits and 

protective padding. For the most part, the experimental efforts have been accomplished to support 

the efforts to reduce the injuries and fatalities associated with automotive highway accidents. The 

data resulting from these experimental efforts are used to specify sideward force limits or allowable 

impact pressures. Measured impact responses of specially designed automotive SID may be used 

to limit the likelihood of injury from sideward collisions. 

The shoulder biodynamic response has not been extensively studied. The more recent studies 

conducted by Bolte, et.al, [ref. 90] and Koh, et.al [ref. 91] provide summaries of the previous 

research and their experimental efforts to define the response and injury thresholds of the shoulder 

complex to sideward and oblique impact. These recent studies have focused upon defining injury 

criteria in terms of shoulder impact force, deflection corridors, and the Y-axis motion of the first 

thoracic vertebra. The experimental efforts by Koh, et al. [ref. 91] supported a conclusion that a 

shoulder deflection of 40 mm for a 50th-percentile male and a Cmax of 

20 percent (chest compression ratio) would result in a 50-percent risk of an abbreviated injury 

scale (AIS) code-2 shoulder injury. A shoulder deflection of 40 mm was exceeded in all but one 

of the impact tests conducted with volunteers reported in Figure 3.3-10. This difference could be 

based upon differences in the experimental methods, analytical methods, test conditions, or the use 

of PMHS versus healthy young, military volunteers.

Neither the current BDR acceleration exposure limits nor the Hybrid III ATD limit values specified 

in the HSIR document relate to the conclusions reached by Koh, et al. [ref. 91].

The current BDR acceleration exposure limits are shown in Figure 3.3-11; it presents the amplitude 

of half-sine impact pulses and the time to peak acceleration. Estimated risk levels were based upon 

the threshold of potentially adverse cardiovascular events (low risk) recorded during test with 

volunteers [refs. 21, 23, 24, and 92]. The high-risk threshold was defined on the basis of the 
92

findings of Beeding and Mosley [ref. 92] and several accidents incurred during tests of escape 

systems (reports unavailable).

Figure 3.3-11. Current Dynamic Response Acceleration Exposure Limits for Sideward Half Sine 

Impacts

In Section 3.3.1, sideward impact injuries that were experienced during NASCAR racing accidents 

are reported. A logistic regression analysis of these injuries versus no injury cases revealed a low 

probability of injury (<0.5 percent) would correlate to a DRy value of 21.9 G versus the existing 

DRy low risk limit of 15 G. A moderate probability of injury threshold would correlate to a DRy

value of 31.3 G, whereas the existing moderate risk threshold is set at 20 G. A high risk probability 

of injury (>50 percent) would correlate to 48.9 G compared to 30 G using the current BDR model.

The model coefficients developed as a result of the tests of volunteers at an impact level of 6 G 

(see Table 3.3-1) were used to calculate the low, moderate, and high risk acceleration exposure 

limits for half-spine impact pulse as show in Figure 3.3-12. 

Figure 3.3-12. Dynamic Response Acceleration Exposure Limits Based on NASCAR Accidental 

Shoulder Complex Injuries
93

The most obvious reason for the difference is that the driver’s seats of the NASCAR vehicles 

provide better protection than the rigid generic seats used in the laboratory tests. For example, the 

NASCAR seats have padding that surrounds the headrest and shoulders of the drivers. 

Recent experiments conducted by the AFRL to study the responses of male and female sideward 

impacts with various helmets and headrest configurations have elicited strong complaints from the 

women. They have refused to participate in impact tests at levels above 6 G. Although no injuries 

of the cervical spine have been experienced, muscle pain and strains have been incurred. Lowering 

of the BDR acceleration exposure limits for the case where no side supports are provided should 

be considered after thorough analysis of these findings. Currently, the HSIR requirements 

implicitly require side supports as a result of the body movement requirements.

This section had two objectives; to develop biomechanical models to simulate the dynamic inertial 

response of the human body to Y-axis impact, and to update the BDR model and exposure limits 

for Y-axis. Progress towards meeting these objectives was made. However, the work is not 

complete, and no final recommendations were made.

3.4 THUMS Analyses

WPAFB Sled Test Analysis

The purpose of this study was to determine the similarity between the response of the THUMS 

model and the Hybrid III ATD given existing WPAFB sled tests. There were four tests selected 

for this comparison with frontal, spinal, rear, and lateral loading. The THUMS model was placed 

in a sled configuration that replicated the WPAFB configuration and the recorded seat acceleration 

for each test was applied to model seat. Once the modeling simulations were complete, they were 

compared to the WPAFB results using two methods. The first was a visual inspection of the sled 

test videos compared to the THUMS d3plot (3DPLOT2) files. This comparison resulted in an 

assessment of the overall kinematics of the two results. The other comparison was an assessment 

of the plotted data recorded for both tests. The metrics selected for comparison were seat 

acceleration, belt forces, and head and chest accelerations. These metrics were recorded in the 

WPAFB tests and were outputs of the THUMS model. Once the comparison of the THUMS to the 

WPAFB tests was complete, the THUMS model output was examined for possible injuries in these 

scenarios. These outputs included metrics for injury risk to the head, neck, thorax, lumbar spine, 

and lower extremities. The metrics to evaluate head response were peak head acceleration, HIC15, 

and HIC36. For the neck, Nij was calculated. The thorax response was evaluated with peak chest 

acceleration, the combined thoracic index, sternal deflection, chest deflection, and chest 

acceleration 3-msec clip. The lumbar spine response was evaluated with lumbar spine force. 

Finally, the lower extremity response was evaluated by femur and tibia force. The results of the 

simulation comparisons indicate the THUMS model had a similar response to the Hybrid III ATD 

given the same input. The primary difference seen between the two was a more flexible response 

of the THUMS compared to the Hybrid III. This flexibility was most pronounced in the neck 

flexion, and shoulder and chest deflections. Due to the flexibility of the THUMS, the resulting 

head and chest accelerations tended to lag the Hybrid III acceleration trace and have a lower peak 

value. The results of the injury metric comparison identified possible injury trends between 

simulations. Risk of head injury was highest for the lateral simulations. The risk of chest injury 

 

2 Oasys D3PLOT is a 3D visualization and interrogation package for post-processing the results of LS-DYNA®

analyses.
94

was highest for the rear impact. However, neck injury risk was approximately the same for all 

simulations. The injury metric value for lumbar spine force was highest for the spinal impact. The 

leg forces were highest for the rear and lateral impacts. The results of this comparison indicate the 

THUMS model performs in a similar manner as the Hybrid III ATD. The differences in the 

responses of model and the ATD are primarily due to the flexibility of the THUMS. This flexibility 

of the THUMS would be a more human like response. Based on the similarity between the two 

models, the THUMS should be used in testing to assess occupant injury risk.

Additional information regarding the study can be found in Appendix E.
95

4.0 Findings and NESC Recommendations

4.1 Findings

The following findings were identified:

F-1. The BDR model has the following limitations:

 Assumes basic seat geometry, restraint, and head protection and is therefore only an 

approximation for other seat designs and protection systems.

 Injury probability is not reduced with seat or restraint improvements. 

 Injury risk cannot be lowered by improving the seated environment. 

 Injury risk can only be lowered by reducing the driving loads into the seat.

 Injury risk is not reflected in gap distances between subject and seating support 

surface

 Estimates if an injury or adverse physiological response will occur, not the injury type 

or location.

 Valid for accelerations of less than 500 msec.

 Seat padding or cushions may not amplify transient linear accelerations transmitted to 

the occupant.

 +X-axis limits presume that the seat occupant’s head is protected by a flight helmet 

with a liner meeting the test requirements of ANSI Z-90 (latest edition) or equivalent. 

 Cannot predict injury caused by rigid suit elements (e.g., localized blunt trauma, 

localized point loading, and interference with restraints). 

F-2. The most extensive collection of body region specific biodynamic injury criteria was 

developed for the automotive industry and are being used to ensure automotive safety.

F-3. There is limited direct science to extrapolate most biodynamic injury criteria to spacecraft 

landing.

F-4. Biodynamic injury criteria used by the automotive industry and FAA are designed for 

higher allowable probability of injuries than are acceptable for NASA and manned 

vehicle landings.

F-5. Utilizing biodynamic injury criteria developed for applications other than manned 

spacecraft requires extrapolation, interpretation, and SME judgment.

F-6. The currently available biodynamic criteria can only be used to approximate the HSIR 

required 0.5-percent probability of injuries. 

F-7. Body movement requirements complement the biodynamic injury criteria by establishing 

requirements that limit relative body motions that could result in injury. 

F-8. The BDR is only applicable to seat configurations with equivalent or improved lateral 

support and body movement constraints to the seat design used in its development. 

F-9. HSIR occupant protection requirements can be quantified through numerical modeling 

and simulation, but requirements require physical testing for qualification. 
96

F-10. Occupant biodynamic injury criteria are specific to the physical and numerical model.

F-11. The MADYMO model provided less modeling capabilities than the FTSS and LSTC 

Hybrid III models since these models are FE-based while the MADYMO ATD model 

consists of rigid links and the seated environment is modeled with simple contact.

F-12. The MADYMO and LSTC Hybrid II models did not perform well for correlating with 

sled test data.

F-13. The FTSS Hybrid III model performed best among the three ATD numerical models 

exhibiting the lesser amount of deviation between numerical prediction and test results, 

although they did not consistently produce correlation results to within the desired 

20-percent of the test data.

F-14. The THUMS model is designed to replicate the human body and not intended to directly 

replicate the Hybrid III ATD. 

F-15. The possible sources for the lack of correlation between the numerical models and the 

test data include: test data variability; ATD model limitations; and the absence of a 

standard model development methodology. 

F-16. The Hybrid III ATD was primarily designed and validated for frontal impacts.

F-17. The THUMS model may provide an advantage over the Hybrid III numerical models for 

situations where it is desired to investigate specific regions or features of the human 

body.

F-18. It is more practical to utilize physical and numerical ATDs than the THUMS model to 

assess occupant injury. 

F-19. HSIR requirements are qualification through analysis, which differs from the NHSTA 

and the FAA approach, which requires physical testing and analysis. 

F-20. Crew exposure to microgravity has not been studied in the context of landing impact 

injury. 

F-21. Limited information exists related to female tolerance of impact loads.

F-22. The influence of pressure suits and their design features cannot be evaluated until specific 

developmental designs are defined.

F-23. The FAA and NHTSA utilize standardized testing for commercial aircraft and 

automobile qualification for occupant safety. NASA does not provide for standardized 

testing and instead requires that occupant protection be assessed for all possible load 

conditions. Since all possible load conditions cannot be physically tested, qualification 

can only be performed through analysis.

4.2 NESC Recommendations

The following NESC recommendations were identified and directed toward CxP unless 

otherwise identified:

R-1. It is recommended that NASA perform actual human testing specific to the environment 

and loadings of the planned manned vehicles to anchor the occupant protection injury 

criteria with actual relevant test data. In the past, NASA deemed this type of testing 
97

necessary and performed it for Apollo and similarly, the Russians performed this testing 

for Soyuz. Currently, there is no plan to obtain these data for Orion or for any other 

future planned manned vehicle. (F-3, F-5, F-6)

R-2. Experts from the FAA and NHTSA have offered to review the current HSIR occupant 

protection requirements and to make suggestions for improvements. It is recommended 

that this offer be accepted. (F-3, F-4, F-5)

R-3. Recent sled and VDT testing of the Hybrid III ATD in the Orion seat design has 

disclosed the difficulty with measuring all of the body movement requirements as 

specified in the HSIR. It is recommended these requirements be reviewed and possibly 

simplified. (F-7)

R-4. The FAA requires that for numerical ATD models to be used in place of testing, the 

numerical model must be validated for a physical environment (e.g., seat, constraints, 

etc.) and loading conditions similar to the test conditions that the model is intended to 

replace. It is recommended that before any numerical model is used to perform a final 

assessment of occupant injury that the model be validated through physical testing. This 

testing should be performed in a similar physical environment, including similar or 

identical seat, harness, helmet, suit, and additional physical hardware components that 

interact with the occupant, and for different size crew representing both male and female 

crew. In addition, loading conditions for these physical modeling validation tests should 

be similar to the loading conditions that will ultimately be used for the occupant 

protection assessments. (F-12, F-13)

R-5. It is recommended that additional human data be mined and analyzed to refine the injury 

criteria. (F-2 through F-6)

R-6. Additional research is required to address landing impact injury tolerance and gender 

differences. (F-21)

R-7. Additional research should be conducted to investigate the unique challenges of landing 

impact, such as crew deconditioning and suit-related trauma. Emphasis should be given 

to areas of the body that carry the highest impact loads during landing impact, such as 

hip, shoulder, and spinal column. (F-22)

R-8. Consider the influence of specific design features of pressure suits, such as placement and 

design of suit bearings, on potential crew injuries due to blunt force trauma. (F-22)

R-9. It is recommended that the qualification process be replaced or simplified through the use 

of sled or vertical drop tower (VDT) testing. (F-23)

5.0 Other Deliverables

No unique hardware, software, or data packages, outside those contained in this report, were 

disseminated to other parties outside this assessment.
98

6.0 Lessons Learned

This assessment was supported by a combination of contractors and civil servants at JSC and Glenn 

Research Center. In addition, several consultants were used to provide expertise that NASA did 

not have in-house in the area of occupant injury criteria and occupant protection. Utilization of 

these experts was critical for the success of this effort by providing the needed knowledge and for 

building a stronger in-house capability to work in the area of occupant protection in the future.

While there were many advantages to utilizing expert consultants, it was critical that the 

consultants provide written documentation on a regular basis so that items discussed verbally 

during meetings and teleconferences were captured in an efficient and timely manner. Waiting 

until the assessment is complete to write the final report is too late to capture all of the valuable 

discussions and insights. It is important that a regular reporting process be in place from the effort’s 

inception so that the final report progresses instead of a last item to complete.

7.0 Definition of Terms

Finding A conclusion based on facts established by the investigating authority. 

Lessons Learned Knowledge or understanding gained by experience. The experience may be 

positive, as in a successful test or mission, or negative, as in a mishap or 

failure. A lesson must be significant in that it has real or assumed impact on 

operations; valid in that it is factually and technically correct; and applicable 

in that it identifies a specific design, process, or decision that reduces or 

limits the potential for failures and mishaps, or reinforces a positive result. 

Observation A factor, event, or circumstance identified during the assessment that did not 

contribute to the problem, but if left uncorrected has the potential to cause a 

mishap, injury, or increase the severity should a mishap occur. Alternatively, 

an observation could be a positive acknowledgement of a 

Center/Program/Project/Organization’s operational structure, tools, and/or 

support provided.

Problem The subject of the independent technical assessment.

Recommendation An action identified by the NESC to correct a root cause or deficiency 

identified during the investigation. The recommendations may be used by the 

responsible Center/Program/Project/Organization in the preparation of a 

corrective action plan.

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Crew Exploration Vehicle (CEV) (Orion) Occupant Protection

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Currie-Gregg, Nancy J.; Gernhardt, Michael L.; Lawrence, Charles; 

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NASA Langley Research Center

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