Abstract A finite element model of a human chest, modified THUMS model, was validated by means of mechanical post mortem human subject table top and indenter tests. The tissue conditions for the table top tests were eviscerated, denuded and intact. The loading configurations were diagonal belt, hub, distributed and criss‐cross belt. The validation was carried out by comparing predicted chest deflections and posterior reaction force by measured deflections and forces. The indenter tests were carried out on the denuded tissue condition. For the indenter tests the validation was carried out by comparing predicted chest deflections at various locations with measured post mortem human subject deflections. In the table top validation the chest effective stiffness predicted with the model for the various tissue and loading conditions were all within the 95% confidence interval of the mechanical test results. For the model evaluation by means of the indenter tests generally the correct displacement trend was predicted. The thorax model was used to carry out a force and energy analysis. For the whole thorax the greatest amount of energy for the diagonal belt, distributed and criss‐cross loading conditions was absorbed by the skin and fat. For the hub loading condition the greatest amount of energy was absorbed by the internal tissues of the chest. For the ribs the greatest amount of energy was for the intact thorax loaded by the diagonal belt and the primary energy‐absorbing structures were ribs 8 – 10 and the left clavicle. For the distributed loading condition, the greatest amount of energy was absorbed in left and right ribs 8. For the hub loading condition, the greatest amount of energy was absorbed by left and right ribs 2 – 4. For the criss‐cross loading condition, the greatest amount of energy was absorbed by left and right ribs 8 – 10. Keywords chest injuries, human body model, PMHS, energy, maximum principal strain I. INTRODUCTION In automobile crashes thoracic injuries are very common. It was found that the thorax was the most frequently seriously (AIS 3+) injured body part in automobile frontal accidents [1]. For vehicle occupants age 45+ the parts of the thorax that were most frequently injured were lungs and ribs. Rib fractures were found to be sustained by almost 40% of the elderly passengers [2]. Rib fractures and flail chest are the most frequent types of thoracic injuries for both drivers and passengers, followed by pulmonary, liver and arterial injuries. The prevention of these injuries remains a major concern. Thoracic injuries are sustained when the chest is dynamically loaded at high energy levels. Typically the chest of a belted occupant in a 56 km/h full front rigid wall crash is compressed 30 – 60 mm in 0.1 seconds at a force level of 4 – 6 kN [3]. Therefore a need exists to study human thoracic response to dynamic loading under various impact conditions. Such conditions involve loading through restraint systems, e.g., blunt loading from airbags and/or steering system, seatbelt alone or in combination with airbags. Thoracic response under various loading conditions has been evaluated using post mortem human subjects (PMHS) [4‐5]. These studies have been time consuming and the vast majority of them have been conducted with PMHS representing the older population. Thoracic response evaluations can be done more efficiently with finite element (FE) models of the human body. The mathematical human body models can improve our understanding of injury mechanisms and predict injury risks. The models can enable evaluation of physical variables mechanically related to injury, e.g. energy and Bengt. Pipkorn Autoliv Research, Vårgårda Sweden (corresponding author +46 (0)322 626341, [email protected]). Richard Kent is professor in Mechanical and Aerospace Engineering at University of Virginia. Validation of a Human Body Thorax Model and its Use for Force, Energy and Strain Analysis in Various Loading Conditions Bengt Pipkorn 1 , Richard Kent 2 IRC-11-56 Ircobi Conference 2011 - 210 -
Transcript
Abstract A finite element model of a human chest, modified THUMS model, was validated by means of
mechanical post mortem human subject table top and indenter tests. The tissue conditions for the table top tests were eviscerated, denuded and intact. The loading configurations were diagonal belt, hub, distributed and criss‐cross belt. The validation was carried out by comparing predicted chest deflections and posterior reaction force by measured deflections and forces. The indenter tests were carried out on the denuded tissue condition. For the indenter tests the validation was carried out by comparing predicted chest deflections at various locations with measured post mortem human subject deflections.
In the table top validation the chest effective stiffness predicted with the model for the various tissue and
loading conditions were all within the 95% confidence interval of the mechanical test results. For the model evaluation by means of the indenter tests generally the correct displacement trend was predicted.
The thorax model was used to carry out a force and energy analysis. For the whole thorax the greatest
amount of energy for the diagonal belt, distributed and criss‐cross loading conditions was absorbed by the skin and fat. For the hub loading condition the greatest amount of energy was absorbed by the internal tissues of the chest. For the ribs the greatest amount of energy was for the intact thorax loaded by the diagonal belt and the primary energy‐absorbing structures were ribs 8 – 10 and the left clavicle. For the distributed loading condition, the greatest amount of energy was absorbed in left and right ribs 8. For the hub loading condition, the greatest amount of energy was absorbed by left and right ribs 2 – 4. For the criss‐cross loading condition, the greatest amount of energy was absorbed by left and right ribs 8 – 10. Keywords chest injuries, human body model, PMHS, energy, maximum principal strain
I. INTRODUCTION In automobile crashes thoracic injuries are very common. It was found that the thorax was the most
frequently seriously (AIS 3+) injured body part in automobile frontal accidents [1]. For vehicle occupants age 45+ the parts of the thorax that were most frequently injured were lungs and ribs. Rib fractures were found to be sustained by almost 40% of the elderly passengers [2]. Rib fractures and flail chest are the most frequent types of thoracic injuries for both drivers and passengers, followed by pulmonary, liver and arterial injuries. The prevention of these injuries remains a major concern.
Thoracic injuries are sustained when the chest is dynamically loaded at high energy levels. Typically the chest
of a belted occupant in a 56 km/h full front rigid wall crash is compressed 30 – 60 mm in 0.1 seconds at a force level of 4 – 6 kN [3]. Therefore a need exists to study human thoracic response to dynamic loading under various impact conditions. Such conditions involve loading through restraint systems, e.g., blunt loading from airbags and/or steering system, seatbelt alone or in combination with airbags. Thoracic response under various loading conditions has been evaluated using post mortem human subjects (PMHS) [4‐5]. These studies have been time consuming and the vast majority of them have been conducted with PMHS representing the older population. Thoracic response evaluations can be done more efficiently with finite element (FE) models of the human body. The mathematical human body models can improve our understanding of injury mechanisms and predict injury risks. The models can enable evaluation of physical variables mechanically related to injury, e.g. energy and
Bengt. Pipkorn Autoliv Research, Vårgårda Sweden (corresponding author +46 (0)322 626341, [email protected]). Richard Kent is professor in Mechanical and Aerospace Engineering at University of Virginia.
Validation of a Human Body Thorax Model and its Use for Force, Energy and Strain Analysis in Various Loading Conditions
Bengt Pipkorn1, Richard Kent2
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strain [6]. Also, the models can be used to evaluate complex chest deformation patterns that can not be assessed using mechanical anthropometric test devices (ATDs) with a single‐point slider at the sternum as is used in the Hybrid‐III dummy. It was found in testing with post mortem human subjects (PMHS) loaded by a single diagonal belt that maximum chest deflection was observed at points other than mid sternum [7].
There is a considerable number of mathematical human body models developed. The models are both
isolated chest models [8] and whole body models [9‐11]. The models are developed to understand thoracic structural responses and injuries under various impact loads. Gross motions and stress or strain distributions of the body at impact can be predicted. The models reproduce anatomical geometry and biomechanical properties of the human body such as the stiffness of the bone and the flexibility of the skin. Each joint is modeled anatomically with the major ligaments and tendons and contact interfaces between the proximal and distal ends of the joint. Ribs have typically been modeled as homogeneous isotropic, linear elastic‐plastic materials, sometimes with different types of element for trabecular and cortical bone. The costal cartilage and the ligaments have been modeled as homogeneous, isotropic linear elastic materials; the thoracic vertebrae as rigid bodies; the muscles and the diaphragm as linear elastic membranes; and the entire interior thoracic volume as an incompressible viscoelastic material (or each organ as a hyperelastic material). These all represent simplifications of reality. Even if human models can be built and biomaterials can be applied, one of the greatest challenges remains; their proper validation. Experimental tests on PMHS and on PMHS body segments are the data that are available. In addition there is a need to take human variability into account. In a table top sensitivity study using a thorax model it was found that the thorax orientation had the largest effect on the force‐deflection response [12]. In addition it was found that the costal cartilage model played an important role in the response to diagonal belt loading.
Human body models have been validated by means of sled, impactor and belt compression tests over range
of impact energies [13‐15]. The aim of this study is to stepwise validate a thorax model by comparing predictions from the model with results from PMHS tests for various tissue‐ and loading conditions. An extended aim is to use the model to evaluate the force flow through‐ and the energy absorbed by various structures or the thorax for various loading conditions.
II. METHODS
The upper body part of THUMS developed by Toyota Central R&D Labs., Inc. and Toyota Motor Corporation, was used as the thorax model [16]. The original model was updated with a number of in‐house modifications. The complete chest was remeshed, rhomboideus muscle elements were added and material data were modified. The ribs were modeled using a combination of solid and shell elements. The cortical bone was modeled with shell elements and the trabecular (spongy) bone was modeled with solid elements. In The original THUMS the trabecular bone was modeled with 1 solid element through the thickness of the ribs while in the modified THUMS the trabecular bone was modeled with 3 solid elements. The skin of the original THUMS was modeled with one element through the thickness while in the modified THUMS the skin was represented by 2 solid elements. For the complete original THUMS the number of nodes and elements were 66617 and 87282 respectively while for the complete modified THUMS the number of nodes and solid elements were 110 100 and 156 522 respectively. The ribs of the original THUMS consisted of 1022 shell elements and 3128 solid element. For the ribs of the modified THUMS the number of shell elements was 18512 and the number of solid elements was 15276. In the original THUMS there was a failure criterion defined for the cortical bone which was excluded in the modified THUMS. The rohmboideus muscle was added due to the fact that the muscle was missing in the original model and the muscle is an important connection between the scapula and the rest of the body.
The material model used for the modified THUMS ribs was an elasto‐plastic model. For the cortical bone the
Young’s modulus was 10.2 GPa, yield stress was 65 MPa. For the trabecular bone the Young’s modulus was 40 MPa, and the yield stress was 1.8 MPa. No strain‐rate dependency was defined. The column vertebrae were modeled with rigid bodies connected by spring elements. The intervertebral disks were modeled with an elastic material model. The ligaments were represented by spring elements. The sternum was modeled with a similar material model as the ribs and the cartilage was modeled using an elasto‐plastic material. The spino‐costal joint was defined with spring elements and the sterno‐costal joint was defined with solid elements. The internal
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organs were modeled with viscoelastic solid elements but were simplified as a large balloon‐like structure inside the rib cage. The geometry and detailed description of the table top compression tests were presented by Kent et al. [4‐5]. 67 tests on 15 PMHS were used to develop the chest effective stiffness for the various loading conditions and tissue states. The chest deflection of the PMHS was generated at approximately 900 mm/ms which was a rate similar to that experienced by restrained PMHS in 48 km/h frontal sled tests [4]. Each mechanical compression test was replicated with the mathematical model.
Three tissue states were replicated in the thorax model. The states were intact, denuded and eviscerated. In
the denuded tissue state all superficial solids and shells from the intact model were deleted. In the eviscerated tissue state all intra‐ribcage solids and shells were deleted.
Four types of loading conditions were replicated with the thorax model. The loading conditions were hub,
diagonal belt, distributed and criss‐cross belt. The hub was modeled as a cylindrical rigid body with a diameter of 152 mm. The distributed belt loading was represented with 203 mm width and 2 mm thick shells, the loaded area approximately between the 2:nd rib and 7:th rib. Single diagonal belt and double diagonal belt were represented by 50 mm width and 2 mm thick shell to pass over the shoulder and cross the anterior thorax approximately 30 degrees tilted to the saggital plane.
The single belt model was similar in size and orientation to an automotive seatbelt. Penalty contact was
defined between the loader and thorax with 0.7 in friction coefficient, due to the fact that a sheet of sandpaper was attached to the anterior thorax side of load in order to prevent the loader from sliding along the surface of the thorax [4‐5].
The thorax was positioned on the table by initially positioning the thorax a small distance above the table and
letting the thorax move to and impact the table under the influence of gravity. The positioning procedure was repeated for each tissue state. Once the initial position was defined for each tissue state the diagonal belt, hub, distributed, and criss‐cross belt was added. In addition to the diagonal belt, hub, distributed and criss‐cross belt validation of the thorax model an
indenter evaluation of the thorax model was carried out (Figure 1). The responses of 5 male cadaver torsos were evaluated by dynamic anterior loading by rigid rectangular indentors designed to approximate a section of a shoulder belt [17]. Indentor load and three‐dimensional deflection measurements were recorded in order to quantify regional force‐deflection and mechanical coupling of the anterior ribcage. Triaxial deflection measurements were taken at nine measurement locations (measurement sites). However, all measurements were not taken in all tests. The average deflection for the 5 cadaver torsos was computed. The tests were replicated with the modified THUMS thorax model validated by means of the table top tests. The indentor displacements varied somewhat between the tests. Therefore the deflection measurements were normalized to 1 based on the indentor displacement. Equation 1 was used for the calculation of the normalized values illustrated in the presentation of load site deflection:
Calculated value = (measurement site deflection)/(indentor site deflection) (1)
The thorax model was used to predict cross section resulting force, maximum principal strain and energy for each rib (entire rib) and clavicles for the various tissue states and loading configurations. The predicted strain energy for coastal cartilage, sternum, internal organs and skin and fat was also evaluated.
III. RESULTS
Results Table Top Validation
The PMHS chest stiffness was scaled to correspond to a 50th percentile male human. The predicted effective chest stiffness was within the experimental range for all tissue states and all loading configurations (Figure 2). In addition the same trend was shown for the various tissue states for all loading conditions. The intact conditions showed the greatest chest stiffness while the eviscerated condition showed the lowest chest stiffness. For the distributed loading conditions the predicted chest stiffness for all tissue conditions was lower than the average PMHS stiffness.
Summary of Stiffness Trends (Mean and Range Shown)
0100200300400500600700800900
10001100120013001400
Int Den Evis Int Den Evis Int Den Evis Int Den EvisTissue Condition
Effe
ctiv
e St
iffne
ss (N
/cm
)
15.2 cm
20.3 cm
5 cm Ynotch
PMHS
Modified THUMS
Figure 2. Chest effective stiffness for a 50th percentile male. The triangles are test results from PMHS data and circles are modified THUMS predictions. 15 PMHS were tested
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Results Indenter Evaluation
The force‐deflection values were normalized to the standard anthropometry of the 50th percentile male weighing 75 kg [18]. The displacements in the indenter test results were normalized to 1 by equation 1. Full displacement of the indenter was 0 and no displacement was 1.
In the upper left indenter evaluation the Indenter displacement was 36 mm. For the upper left indenter all
rib displacement predictions but 2 were within the range of the PMHS displacements (Figure 3). The two displacement predictions that were not within the range of the PMHS displacements were right rib 6 and sternum at rib 1. Greatest predicted and measured displacement was for sternum at rib 5. Smallest predicted and measured displacement was for right rib 7. Greatest difference between the predicted deflection and measured deflection was 6 mm. It was for the left rib 5.
Upper Left Indenter
0,0
0,2
0,4
0,6
0,8
1,0
Right R
ib 3
Right R
ib 5
Right R
ib 6
Right R
ib 7
Sternu
m at R
ib 1
Sternu
m at R
ib 5
Left R
ib 3
Left R
ib 5
Left R
ib 6
Left R
ib 7
PMHSModified THUMS
INDENTER
Figure 3. Upper left indenter normalized average displacement measurements for 5 cadaver torsos for right ribs 3, 5, 6, 7 and sternum at ribs 1 and 5 and left ribs 3, 5, 6 and 7. The range of the measurements is shown
For the sternum indenter evaluation the indenter displacement was 34 mm. All predicted rib displacements
but one were within the range of the of the PMHS displacements. The prediction that was not within the range of the PMHS displacement was for right rib 6. The PMHS displacements were similar for the left and right side (Figure 4). For the PMHS there was a maximum at left and right rib 6. For the modified THUMS predictions the displacements were also similar for the left and right side. However the predicted maximum was for left and right rib 5. Greatest difference between the predicted and measured deflection was 7 mm. It was for the left and right rib 6.
Sternum Indenter
0,0
0,2
0,4
0,6
0,8
1,0
Right R
ib 3
Right R
ib 5
Right R
ib 6
Right R
ib 7
Sternu
m at R
ib 1
Sternu
m at R
ib 5
Left R
ib 3
Left R
ib 5
Left R
ib 6
Left R
ib 7
PMHSModified THUMS
INDENTER
Figure 4. Sternum indenter normalized averaged displacement measurements for 5 cadaver torsos for right ribs 3, 5, 6, 7 and sternum at ribs 1 and 5 and left ribs 3, 5, 6 and 7. The range of the measurements is shown
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For the lower right indenter evaluation the indenter displacement was 36 mm. For the model, 3 predicted rib displacements were within the range of the PMHS displacements (Figure 5). The rib displacement predictions that were not within the range of the PMHS displacements were for right rib 3, 5, 7, sternum at rib 1, left rib 6 and 7. Greatest difference between predicted and measured displacement was 9 mm. It was for right rib 3 and left rib 6.
Lower Right Indenter
0,0
0,2
0,4
0,6
0,8
1,0
Right R
ib 3
Right R
ib 5
Right R
ib 6
Right R
ib 7
Sternu
m at R
ib 1
Sternu
m at R
ib 5
Left R
ib 3
Left R
ib 5
Left R
ib 6
Left R
ib 7
PMHSModified THUMS
INDENTER
Figure 5. Lower right indenter normalized average displacement measurements for 5 cadaver torsos for
right ribs 3, 5, 6, 7 and sternum at ribs 1 and 5 and left ribs 3, 5, 6 and 7. The range of the measurements is shown
Results force, energy and strain analysis for various loading conditions
The predicted cross section resulting force, maximum principal strain and energy for each rib and clavicles were evaluated. For the evaluation cross sections for each rib were defined at approximately the most lateral point of each rib. For the clavicle the cross sections were defined at approximately the midpoint of the bone. The predicted energy used in the evaluation was the strain energy for each rib and clavicles.
The cross section force, energy and maximum principal strain were evaluated at 40 mm of chest deflection
for diagonal belt, hub, distributed and criss‐cross loading condition. The evaluation was carried out for the intact tissue condition.
For the diagonal belt loading configuration the greatest cross section forces were for the left clavicle, right rib 1 and right rib 8 ‐11 (Figure 6). For the hub loading configuration generally the cross section forces were small. The greatest cross section force was for left and right rib 1. For the distributed loading configuration the greatest force was predicted for left rib 1, 4, 8 and right rib 1 and 8. For the criss‐cross loading configuration the greatest forces were predicted for left and right rib 1, 8, 10 and 11 and right rib 1, 4, 8, 10 and 11. The forces for the major part of the ribs were between 30 ‐ 100 N.
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Cross Section Force Left Rib Including Clavicle
0
50
100
150
200
1 2 3 4 5 6 7 8 9 10 11 12 13
Forc
e (N
)
Diagonal BeltHubDistributedCriss-Cross
Intact
Rib Clavicle
Cross Section Force Right Rib Including Clavicle
0
50
100
150
200
1 2 3 4 5 6 7 8 9 10 11 12 13
Forc
e (N
)
Diagonal BeltHubDistributedCriss-Cross
Intact
RibClavicle
Figure 6. Cross section force for rib 1 – 12 and clavicle for intact chest for diagonal belt, hub, distributed and criss‐cross belt loading
For the diagonal belt in the intact tissue configuration the greatest strain energy (entire rib) was for right rib 8 – 10 and for the left clavicle (Figure 7). Low force was obtained for all ribs on the left hand side. For the hub configuration the greatest energy absorbed was for rib 3 ‐ 5 for both left and right side. For the distributed configuration the energy absorbed for the ribs was generally low and distributed evenly between the left and right ribs. For the criss‐cross configuration greatest energy was absorbed for rib 8 ‐10 for both left and right side. In addition for the criss‐cross loading configuration a significant amount of energy was absorbed by both clavicles.
Energy Left Rib Including Clavicle
0,00,10,20,30,40,50,60,70,8
1 2 3 4 5 6 7 8 9 10 11 12 13
Ene
rgy
(J)
Diagonal BeltHubDistributedCriss-Cross
Intact
RibClavicle
Energy Right Rib Including Clavicle
0,00,10,20,30,40,50,60,70,8
1 2 3 4 5 6 7 8 9 10 11 12 13
Ener
gy (J
)
Diagonal BeltHubDistributedCriss-Cross
Intact
RibClavicle
Figure 7. Energy absorbed in rib 1 – 12 and clavicle for intact chest for diagonal belt, hub, distributed and criss‐cross belt loading
For the diagonal belt in the intact tissue condition the greatest maximum principal strain was for left rib 1 and 2 and right rib 1, 8, 9 and 10 and for the left clavicle (Figure 8). For the hub loading condition the greatest maximum principal strain was for ribs 3 ‐ 5 for both left and right side. For the distributed and hub loading conditions the maximum principal strain was generally more evenly distributed between the ribs than for the other loading conditions. For the criss‐cross belt the greatest maximum principal strain was for rib 8 ‐ 10 for both left and right side. In addition there was significant strain for both the clavicles and rib 1 and 2.
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Maximum Principal Strain Left Ribs Including Clavicle
0,000
0,002
0,004
0,006
0,008
0,010
0,012
0,014
1 2 3 4 5 6 7 8 9 10 11 12 13
Stra
in
Diagonal BeltHubDistributedCriss-Cross
Intact
RibClavicle
Maximum Principal Strain Right Ribs Including Clavicle
0,000
0,002
0,004
0,006
0,008
0,010
0,012
0,014
1 2 3 4 5 6 7 8 9 10 11 12 13
Stra
in
Diagonal BeltHubDistributedCriss-Cross
Intact
RibClavicle
Figure 8. Maximum Principal Strain in rib 1 – 12 and clavicle for intact chest for diagonal belt, hub, distributed and criss‐cross belt loading
For the diagonal belt in the intact tissue condition a significant amount of the energy (strain energy) was absorbed by the skin and fat and internal organs (Figure 9). For the hub loading condition greatest amount of energy was absorbed by the internal organs followed by the ribs. For the distributed and criss‐cross loading conditions greatest amount of the energy was absorbed by the skin and fat and the organs.
Energy All Ribs, Internal Organs, Skin and Fat
0,0
5,0
10,0
15,0
20,0
25,0
30,0
35,0
Ene
rgy
(J)
Diagonal BeltHubDistributedCriss-Cross
Intact
All Ribs CostalCartilage
Sternum Internal Organs
Skin & Fat
Figure 9. Energy absorbed by ribs, costal cartilage, sternum, internal organs and skin and fat for intact chest for diagonal belt, hub, distributed and criss‐cross belt loading
IV. DISCUSSION For the table top analysis generally good agreement between model predictions and mechanical test results
was obtained (Figure 2). The highest stiffness for the intact thorax condition was obtained for the distributed loading condition. The distributed loading condition loaded most ribs and also the clavicles (Figure 6). The stiffness of the chest in the criss‐cross loading configuration was greater than the stiffness of the chest in the diagonal belt loading configuration. It was due to the fact that the criss‐cross belt loaded the lower ribs 8 – 11 on both left and right side in addition to loading the left and right clavicles and rib 1. The diagonal belt loaded the left clavicle, left and right rib 1 and right rib 8 – 11.
In the table top validation the predicted effective chest stiffness in the distributed loading configuration was
consistently lower than the average PMHS stiffness. A reason may be that there was a discrepancy between the mathematical distributed loading model and the mechanical counterpart (Figure 2). In the indenter evaluation it was observed that for the upper left indenter and sternum indenter there was a weaker coupling between the left and right side of the thorax for the mathematical model than for the PMHS. When the upper left side and sternum was loaded with the indenter there was generally smaller deflection on the right side of the thorax (right ribs 3 – 7) for the mathematical model than for the PMHS. For the lower indenter, however, there was a stronger coupling between the left and right side for the mathematical model than was observed for the PMHS. Which means, when the lower right side was loaded with the indenter there was generally greater deflection on the left side of the thorax (left ribs 3 – 7) for the mathematical model than for the PMHS.
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The modified THUMS chest model predicted the PMHS chest stiffness in the table top tests. In the indenter
validation there was some small discrepancies between the model predictions and the PMHS test results. In particular the discrepancies were for the lower indenter. However the trends predicted in the indenter simulations were similar to those observed in the PMHS indenter tests. The indenter tests were local deformation of the chest that is generally not the case in frontal car crashes in which the load generally is distributed by a belt and an airbag. However, the indenter evaluation contributed to the validation of the model.
Limitation of the table top test configuration include the use of a constrained back condition which may
result in different response than a thorax loaded only by its inertia, as it is in most frontal car crashes. One possible effect of the posterior boundary is an increase in stiffness due to constraint of the costovertebral and costotransverse articulations. However, it is likely to be small compared to interspecimen variability.
The results from the indenter test evaluation indicate that there is a need to continue the evaluation of
sternum and costal cartilage stiffness. However, there is an uncertainty in the PMHS test results due to the fact that there were few PMHS specimens evaluated.
The in‐house modifications to the model were carried to improve the stability of the chest model and also to
reduce the stiffness of the thorax model in the hub loading and diagonal belt loading configurations. The rohmboideus muscle was added due to the fact that the muscle was missing in the original model and the muscle is an important connection between the scapula and the chest. In the model strain rate effects were excluded. Adding such effects will increase the thorax stiffness when loaded at high energy levels.
The forces for the major part of the ribs were between 30 – 100 N (Figure 6).The failure force for ribs was in
an experimental study found to vary from 48 to 147N [19]. Therefore the forces for some of the ribs were in a force range in which a fracture can occur. However, the forces applied in the experimental study were not necessarily the same as the cross‐section forces predicted by the model.
In the diagonal belt configuration there were high levels of maximum principal strain on right rib 1 in the
mathematical model. The reason was that there was a strong coupling between rib 1 and the sternum. Rib 1 was pulled by the sternum resulting in high levels of maximum principal strain (Figure 10). In the indenter evaluation it was observed that there was a stronger coupling between left and right side of the thorax in the model than in the PMHS. In another study it was found that the properties of costal cartilage play an important role in the thorax response to diagonal belt loading [12]. Therefore the strain and energy absorption by right rib 1 predicted by the model can be somewhat higher than what was sustained by the PMHS in the mechanical tests.
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Figure 10. chest deformation pattern for intact chest for diagonal belt loading at 0% and at 35% chest deflection
One study concluded that the potential for seat belt related rib fractures declined with increasing body mass
index [20]. The simulations results suggest a potential mechanism for such a decline: the fat and skin of the occupant have a significant energy‐absorbing effect (Figure 9) that may off‐load the ribs. More research is needed to fully understand the phenomenon.
For the hub loading condition the energy absorbed by the skin and fat was small (Figure 9). In the hub
configuration the hub was loading a small area at the middle of the sternum. The fat layer that was compressed by the hub was relatively thin therefore the energy absorbed was small.
In the table top PMHS tests a number of tests for all loading conditions were run beyond the elastic limit of
the chest and numerous rib fractures were sustained by the PMHS [4]. Those tests were replicated with the mathematical model and the predicted maximum principal strain and energy absorbed pattern were compared to the fracture pattern in the PMHS tests. Maximum chest deflection was between 35 – 42%. The predicted maximum principal strain and energy absorbed by each rib was extracted for the same chest deflection for each loading configuration.
In the PMHS test with the diagonal belt loading condition at 35% chest deflection, the greatest number of rib
fractures was obtained for left rib 3 – 6 (Figure 11). For the diagonal belt condition high maximum principal strain was predicted for left and right rib 1, for right rib 2 and for left rib 9 and 10. Significant energy absorption was predicted for right rib 2 and for left rib 9 and 10. While the general asymmetry of the loading is reflected in both the fracture patterns and in the principal strain distribution, the concentration of rib fractures around left rib 3 – 6 was not predicted by the maximum principal strain or by the energy absorbed. This may be a consequence of the use of constant cortical shell thickness in the ribs of the modified THUMS model. Cortical shell thickness has been shown to vary both longitudinally and circumferentially in the ribs and this variability does influence the model response and rib fracture pattern [21‐24].
In the PMHS test with the hub loading condition at 43% chest deflection, the greatest number of rib fractures
was obtained for left and right rib 2 – 5 (Figure 12). Maximum principal strain was predicted for left and right rib 1 and left and right rib 3 – 5. Greatest predicted energy absorbed was for left and right rib 3 – 6. The concentration of rib fractures around rib 2 – 5 in the PMHS test was not predicted by the maximum principal
Figure 11. Rib fracture locations in PMHS tests for diagonal belt loading at 35% chest deflection and modified THUMS predicted maximum principal strain and energy for ribs 1 – 10. S50 is the distance around the rib cage from the mid‐sternal line to the fracture location
Figure 12. Rib fracture locations in PMHS tests for hub loading at 43% chest deflection and modified THUMS predicted maximum principal strain and energy for ribs 1 – 10. S50 is the distance around the rib cage from the mid‐sternal line to the fracture location
In the PMHS test with the distributed loading condition at 35% chest deflection the greatest number of rib
fractures were obtained for left rib 3 – 8 (Figure 13). Highest predicted maximum principal strain was for left and right rib 1 and 5 ‐10. Highest predicted energy absorbed by each rib was for left and right rib 4 – 8. The concentration of rib fractures around rib 3 – 8 was not predicted by maximum principal strain or energy absorbed.
In the PMHS test with the criss‐cross belt loading condition at 29% chest deflection the greatest number of
rib fractures was obtained for left and right rib 2 – 7 (Figure 14). Highest predicted maximum principal strain was for left and right rib 1 and 7 ‐10. Highest predicted energy absorbed by each rib was for left and right rib 7 – 10. The concentration of rib fractures around rib 2 – 7 was not predicted with the maximum principal neither
Figure 13. Rib fracture locations in PMHS tests for distributed loading at 35% chest deflection and modified THUMS predicted maximum principal strain and energy for ribs 1 – 10. S50 is the distance around the rib cage from the mid‐sternal line to the fracture location
Figure 14. Rib fracture locations in PMHS test for Criss‐Cross loading at 29% deflection and modified THUMS predicted maximum principal strain and energy for ribs 1 – 10. S50 is the distance around the rib cage from the mid‐sternal line to the fracture location
For the diagonal belt and criss‐cross belt configurations, greatest predicted energy absorbed and greatest
maximum principal strain was for rib 8 – 10. In the PMHS tests greatest number of rib fractures was around rib 3 ‐ 7. One explanation for the discrepancy between model predictions and mechanical PMHS test results may be that the belt in the model was located somewhat lower on the chest than in the PMHS tests resulting in greater energy and strains in rib 8 – 10. An alternative explanation may be that there was significant difference in thorax geometry between the PMHS and THUMS resulting in a difference in the belt load path between THUMS and PMHS. Another parameter that can influence the stiffness of the thorax is the rib angle [2]. The rib angle of the PMHS tested was not known. Therefore a rib angle comparison between THUMS and the PMHS cannot be carried out. Rib curvature, skew, and taper have been proven to be of minor importance [25].
Generally there was high predicted maximum principal strain for rib 1 and 2. In the model the same cortical
bone thickness was used for rib 1 and 2 as for the other ribs. However, the cortical bone thickness of a human can be greater for rib 1 and 2. Therefore increasing the cortical bone thickness for rib 1 and 2 in the model may
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reduce maximum principal strain of rib 1 and 2 and make the predicted peak maximum principal strain distribution more similar to the distribution of fractures of the PMHS in the table top tests. In future research analysis a sensitivity analysis will be carried out to evaluate the influence of rib 1 and 2 cortical bone thickness and belt routing over the chest on both chest stiffness prediction and peak energy and strain predictions.
The table top rib fracture prediction evaluation with modified THUMS indicates that a combination of energy
absorbed and maximum principal strain may be feasible to prediction injury in the distributed and criss‐cross loading configuration. It may not be feasible for the diagonal belt and hub loading configuration. However, the evaluation will be carried out for other PMHS data sets. Additional PMHS rib fracture data is needed before these issues can be fully elucidated.
V. CONCLUSIONS
The validated mathematical thorax model predicted PMHS force deflection characteristics for diagonal belt, hub, distributed and criss‐cross loading conditions and eviscerated, denuded and intact tissue conditions. For the intact thorax condition loaded with the diagonal belt the highest cross sectional force was for the left clavicle, right rib 1 and right rib 8‐11. The greatest amount of energy was absorbed by rib 8 – 10 and the left clavicle. Peak maximum principal strain was predicted for left and right rib 1, left rib 2, left clavicle and right rib 8‐ 10. For the distributed loading condition, the highest cross sectional force was for left and right rib 4 and 8. Greatest amount of energy was absorbed in left and right rib 8. Predicted maximum principal strain was evenly distributed between the ribs. For the hub loading condition, highest predicted cross sectional forces were for left and right rib 1. Greatest amount of energy absorbed and peak strain levels were predicted for left and right rib 2 – 4. For the criss‐cross loading condition highest predicted cross sectional forces were for rib 8 – 11. Greatest amount of energy absorbed and peak strain levels were for left and right rib 8 – 10. Greatest amount of energy for the diagonal belt, distributed and criss‐cross loading condition was absorbed by the skin and fat. For the hub loading condition greatest amount of energy was absorbed by the internal tissues of the chest.
VI. References 1. Carroll, J., et al. Overview of Serious Thorax Injuries in European Frontal Car Crash Accidents and
Implications for Crash Test Dummy Development. in International Research Council on the Biomechanics of Injury (IRCOBI). 2010. Hannover, Germany.
2. Kent, R., Structural Material Changes in the Aging Thorax and Their Role in Crash Protection for Older Occupants. Stapp Car Crash Journal, 2005.
3. Kent, R., et al. Age‐related Changes in the Effective Stiffness of the Human Thorax Using Four Loading Conditions. in International Research Council On Biomechanics Of Injury (IRCOBI). 2003.
4. Kent, R., L. David, and S. Chris, Thoracic Response to Dynamic, Non‐lmpact Loading from a Hub, Distributed Belt, Diagonal Belt, and Double Diagonal Belts. Stapp Car Crash Journal, 2004.
5. Kent, R., D. Murakami, and S. Kobayashi. Frontal Thoracic Response to Dvnamic Loading: the Role of Superficial Tissues, Viscera, and the Rib Cage. in International Research Council On Biomechanics Of Injury (IRCOBI). 2005.
6. Rouhana, S.W., P.G. Bedewi, and P. Prasad, Biomechanics of 4‐Point Seat Belt Systems in Frontal Impact. Stapp Car Crash Journal, 2003.
7. Ceasari, D. and R. Bouquet. Comparison of Hybrid III and Human Cadaver Thorax Deformations Loaded by a Thoracic Belt. in 38th Stapp Car Crash Conference. 1994: SAE 942209.
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8. Kimpara, H., Development of a Three‐Dimensional F E Chest Model for the 5th Percentile Female. Stapp Car Crash Journal, 2005.
9. Baudrit, P., et al., Comparative Studies of Dummy and Human Body Models Behavior in Frontal and Lateral Impact Conditions. Stapp Car Crash Journal, 1999.
10. Iwamoto, M., et al. Development of a Finite Element Model of the Total Human Model for Safety (THUMS) and Application to Injury Reconstructlon. in International Research Council On Biomechanics Of Injury (IRCOBI). 2002.
11. Robins, S. Human Model for Safety ‐ A Joint Effort Towards the Development of Refined Human‐Like Car Occupatn Models. in 17th Inernational Technical Conference on the Enhanced Safety Vehicle. 2001.
12. Murakami, D., et al., Finite Element Analysis of Hard and Soft Tissue Contributions to Thoracic Response: Sensitivity Analysis of Fluctuations in Boundary Conditions. Stapp Car Crash Journal, 2006.
13. Oshita, F. and M. Kiuoshi. Development of Finite Element Model of the Human Body. in 7th International LS‐Dyna Users Conference. 2002.
14. Iwamoto, M. Ankel Skeletal Injury Predictions Using Anisotropic Inelastic Constitutive Model of Cortical Bone Taking into Account Damage Evolution. in Stapp Car Crash Journal. 2005.
15. Chawla, A., et al. Validation of the Cervical Spine Model In THUMS. in 19:th International Technical Conference on Enhanced Safety of Vehicles. 2005. Washington D.C.
16. Labs, T.C.R.D. THUMS (Total HUman Model for Safety) Occupant Model: Version 2.21‐040407. 2005. 17. Shaw, G., et al. Quasi‐Static and dynamic thoracic loading tests: cadaveric torsos. in International
Research Council On Biomechanics Of Injury (IRCOBI). 2007. 18. Eppinger, R.H., R.M. Morgan, and J.H. Marcus. Side Impact Data Analysis. in 9th Internationa Technical
Conference on Experimental Safety Vehicles. 1984. 19. Charpail, E., T. Xavier, and P. Philippe, Characterization of PMHS Ribs: A New Test Methodology. Stapp
Car Crash Journal, 2005. 20. 0'Brian‐Mitchell, B., et al. Analysis of Injury Trends in Frontal University of Michigan Ciren Cases in the
Context of Crash Tests. in 21st International Tehcnical Conference on the Enhanced Safety of Vehicles. 2009. Stuttgart, Germany.
21. Choi, H.Y. and I. Lee. FE model for older population in Japan Society of mehcnaical Engineering Conference. 2009. Fuoka, Japan.
22. Kent, R., et al. Progress on the development of the GHBMC thorac finite element model. in Proceedings 38:th international Workshop on Human Subject for Biomechanical Research. 2010. National Highway Traffic Safety Administration, U.S.D.O.T.
23. Stitzel, J., et al., Defining regional variation in the material properties of human rib cortical bone and its effect on fracture prediction. Stapp Car Crash Journal, , 2003.
24. Mohr, M., et al., Geometry of human ribs pertinent to orthopedic chest‐wall reconstruction. J. Biomech 40:1310‐1317, 2007.
25. Granik, G. and I.D. Stein, Human ribs: Static Testing as a Promising Medical Application. J. Biomech, 1973. 6: p. 237‐240.
