Should a Sternum Fracturebe considered as

Life-Saving or Life-Threatening?A numerical feasibility study

about blunt trauma aorta rupture

BMTE 07.31

MSc Thesis of Erik van WalbeekSeptember, 2007

Exam date: 12 September, 2007

Committee: Prof. dr. P.R.G. BrinkProf. dr. ir. F.N. van de VosseProf. dr. ir. J.S.H.M. Wismansir. L. van Rooij

Eindhoven University of TechnologyDepartment of Biomedical EngineeringDivision of Materials Technology

Maastricht UniversityDepartment of General SurgeryDivision of Traumatology

TNO Automotive Helmond

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Abstract: In medical literature, thoracic great vessel injury is associated with palpable fracture of the sternum. A statistical analysis showed that in the last 5 years in the Netherlands no frontal Blunt Trauma Aorta Rupture (BTAR) occurs if the sternum of survivors is fractured by heavy car crash. Remarkably, survivable BTARs are found within the same patient group if the sternum is not fractured. Due to this paradox, the following hypothesis is formulated. A sternum fracture (as a sign of ventral compression) reduces significantly the change of simultaneous occurrence of BTAR. This hypothesis is controversial, because sternum fractures (16 %) are found by deadly BTAR, which could be an indication that, a crushing mechanism for BTAR is probable. Therefore, this hypothesis only could make sense if the shearing mechanism due to deceleration loads at the isthmus is recognized for playing a key role. This study questions whether a computer model (HUMOS 2) is capable of biomechanically explaining this hypothesis. The method tested to realistically simulate the mass inertia phenomenon of the heart and ascending aorta was to divide the model into two systems. The advantage of applying deceleration to only one system is that deceleration of the heart and ascending aorta model was calculated by the model itself in stead of being applied and therefore mass inertia could lead to a difference in deceleration between ascending and descending aorta (stretching of the isthmus). This resulted in values of von Misses stress for the aorta model which are comparable to another numerical study, which has positively pointed out the feasibility of a finite element model predicting the occurrence of BTAR. These values of von Misses stress could be an indication that a biofidelic simulation has been performed, however, higher von Misses stresses were found at the ascending aorta in stead of the isthmus region, where BTARs are mostly found. Therefore, computer simulations in this report could not reject or explain the compression during deceleration hypothesis. To eventually achieve this goal, the limitations of the model and possible future work have been described. If a biomechanical explanation is available, then this could have impact to the automotive design, which would expose the body to a controlled compression which could prevent BTAR. Because of the difference in healing between BTAR (85%) and other possible thoracic injuries associated to non BTAR (maximum 60%), which could be caused by compression, this would benefit trauma patients.

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Contents: 1. Introduction……………………..5 2. Literature Survey………………..6

2.1 Introduction………………………………... 6 2.2 General background………………………...6 2.3 Road accident data studies………………….8 2.4 Experimental studies………………………10 2.5 Numerical simulations studies…………….13 2.6 Discussion…………………………………15

3. Methods………………………..16 3.1 Introduction………………………………. 16 3.2 Global model structure…………………….17 3.3 Research goals and approach……………...18 3.4 Input of the simulations…………………....19 3.5 Output of the simulations………………….21 3.6 HUMOS 2 model modification……………21 3.7 FE model elements in detail……………….22 3.8 Discussion…………………………………24

4. Results…………………………25 4.1 Introduction………………….……………25 4.2 Instability as a numerical problem………..26 4.3 The airbag model…………………………27 4.4 The hub model……………………………27 4.5 The safety belt model……………………..28

5. Conclusions and discussions…..28 6. Future work……………………29 REFERENCES…………………………………...30

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1. Introduction Heavy car impacts cause severe injuries, which could lead to dead. Better understanding of the injury mechanisms will lead to higher crashworthiness of cars and, therefore, to more survivors of heavy car crashes. Last decades, research in the area of thorax injury biomechanics was focused on the external part of the body and lung injuries. Airbags and other safety equipment were designed to prevent fractures of the bones of the thorax. In medical literature, bone disruption often predicts internal injuries, therefore, research towards internal organs was found to be of less importance. However, recent developments show that research towards internal injuries also is of interest for the benefit of trauma patients. It is even considered that in some cases external injuries could prevent internal injuries to occur or to be not survivable. For example, Blunt Trauma Aorta Rupture (BTAR) as a result of frontal collisions occurs with an extremely high mortality (85%). Deflection (compression of the thorax) was seen as the main cause of BTAR, however, only minor surrounding soft tissue damage was noticed by BTAR survivors [1]. Survival of BTAR can be possible because the inner layer of the aorta ruptures first in the transverse direction [2]. A statistical analysis of a patient database shows that for trauma patients with a sternum fracture no aortic injuries have been reported (negative coincidence) [3]. Remarkably, when the sternum is not fractured BTARs are found within the same patient group. In this study the following hypothesis is developed which tries to explain the statistical analysis: A sternum fracture (as a sign of ventral compression) reduces significantly the change of simultaneous occurrence of BTAR. If this hypothesis can not be rejected, this will benefit the automotive industry, because a controlled ventral compression during ventral deceleration then may lead to less aortic injuries in frontal impacts. Due to this, more sternum fractures will be observed. Nevertheless, due to the considerably better healing of sternum fractures (no mortality) than recovery from BTAR (80-90% mortality), this strategy will be very useful regarding to automotive safety design. An optimal value of deflection, causing no sternum fractures and no BTAR simultaneously could even be possible. Biomechanical research towards the mechanism of BTAR could exclude or confirm the ventral compression during ventral deceleration hypothesis. In literature, several mechanisms have been proposed for the occurrence of BTAR in frontal impacts and a multivariate hypothesis probably is necessary. The hypothesis mentioned above only could make sense if the shearing mechanism at the isthmus plays a key role. Transverse ruptures are mostly found at the isthmus location of the aorta; therefore, the shearing-due-to deceleration-loads mechanism at the isthmus seems to be a crucial mechanism in frontal impacts. A recently performed experimental cadaveric study by Baqué et al. (2006) has significantly demonstrated the shearing-due-to-deceleration-load mechanism with accelerometers, which should lead to a computerized model of BTAR [4]. A computer study of Richens et al. [5] has pointed out the feasibility of a FEM, simulating the probability of occurrence of BTAR with a high-detail model with boundary conditions from a low-detail model. This procedure for applying boundary conditions is also used in this project to obtain a quick and stabile simulation. HUMOS 2 is a recently developed model, which contains finite element methods, multi-body dynamics interaction. The purpose of creating this HUMOS 2 model by several European universities, institutes and companies was to model a whole body during an impact with enough level of detail to investigate the effect on organs. This complicated model consists of more than 50.000 elements. The first research goal in this project is to find which components

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of the HUMOS 2 model are required to find prove for the stretching mechanism at the isthmus location, due to deceleration loads. If high values of von Misses stress are found at the isthmus location of the aorta model, then this model could indicate the occurrence of BTAR. The next research goal is to observe the feasibility of a computer model excluding or confirming the hypothesis mentioned in the second paragraph of this chapter. If more compression leads to less value of von Misses stress in the simulation, then the same components of the HUMOS 2 model that could indicate the shearing-due-to-deceleration-loads mechanism could biomechanically explain the hypothesis that tries to reason the negative coincidence between BTAR and sternum fracture. In this report first in section 2 a literature survey is given as a background for the simulations performed in this study. Then, in section 3 the model that has been used is descriped together with the application of crash specific boundary conditions. The simulation approach is to investigate three realistic cases, in where a human body model crashes with a hub, airbag and safety belts. Boundary conditions for these three cases are created from a vehicle model with low detail human model and applied to certain parts of the high level of detail model. Computed values of von Misses stress in the aorta model will be presented in section 3. In the next section (section 4) the results and limitations of the model used are discussed. If computer simulations could be able to indicate that a controlled compression during deceleration could reduce the risk of BTAR, than further experimental and numerical work is recommended in the last section of this report.

2. Literature Survey

Introduction Publications in the field of injury biomechanics concerning BTAR occurring at traffic accidents with a frontal impact are briefly reviewed in this chapter. First the most common injury location and three possible injury mechanisms are discussed as a general background. Next, studies were divided into three subcategories and regarded uninfluenced by these three injury mechanisms and therefore were seen as objective tools for studying the likeliness of these three mechanisms. The subcategories were defined in order to organize the studies around the kind of study that has been performed. The three kinds of studies were road accident data, experimental and numerical simulation studies. At the end of this chapter it is discussed how this brief literature review is used as motivation of certain choices in the approach that is followed in this study. General background Three mechanisms (one hemodynamical and two mechanical) of the occurring of aortic injuries in traffic accidents are proposed in the literature [6]. However, no actual prove is found and a multivariate hypothesis in which several mechanisms play together would be the most likely. Although the injury process seems to be multivariate, all articles show that aortic injuries occur mostly at one particular location, the isthmus region (See Figure 1.1) of the aorta. Therefore, focusing on this specific region of the aorta seems to be useful for research on the occurring of BTAR. The hemodynamical mechanism given by Oppenheim et al. [7] and Klotz et al. [8] is based on a sudden and dramatic rise in arterial pressure. A first mechanical mechanism is based on

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a sudden stretching of the isthmus location of the aorta (shearing-due-to-deceleration loads) and described by Sevitt et al. [6]. In Figure 1.1 an anatomical visualization shows that the risk at the isthmus location could be reasoned by a difference in deceleration of the heart and the aorta. The difference in deceleration could be caused by a difference of mass and high deceleration applied to the total body. The other mechanical mechanism is the osseous pinch theory, in which aortic injuries occur due to pinching of the aorta between the osseous bodies (bony structures) of the chest. This mechanism was proposed by Crass et al. [9] (See Figure 1.2).

Figure 2.1 and Table 2.1: Anatomical overview. The isthmus region of the aorta is the most

common injury location.

Fig. 2.2, Osseous pinch theory, the osseous bodies pinch the aorta at isthmus region. [9]

A=Aorta ascendens B=Arcus aortae C=Isthmus aortae D=Aorta descendes

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Road accident data studies. In Fig. 2.3 till Fig. 2.8 a straight forward case of Blunt Trauma Aorta Rupture in heavy car crash survivors is illustrated [1]. From these kinds of clinical experiences it has become known that occurrence of BTAR and little surrounding soft tissue damage is possible.

Figure 2.3; 2.4; 2.5; 2.6; 2.7; 2.8, A series of images demonstrating a contained aortic arch rupture (or traumatic pseudo aneurysm) in a 36 year old male driver of an automobile struck head-on at high speeds. He wore a lap belt and shoulder restraint but his vehicle was not equipped with an airbag.

He presented with normal vital signs, complaining only of left posterior chest wall pain. [15]

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To find prove that survivable frontal BTAR occurs with minimal surrounding soft tissue damage, the coincidence of BTAR and sternum fracture has been investigated statistically and showed results that were in contrast with common medical knowledge. In medical literature, a sternum fracture is a predictor for BTAR in frontal impacts. However, recent incidence data from Prismant about heavy car crash survivors [3] show that no coincidence between a sternum fracture and BTAR has been found in five years (See Table 2.2). These data are from a large control group (30.496 patients) who were ‘at risk’ for BTAR injury exposure, because comparable thorax injuries were found in these patients. This appearance seems to be controversial with medical theory.

BTAR + BTAR - Total: SF + 0 2113 2113 SF - 75 28.308 28.383

Total: 75 30.421 30.496 Table 2.2, Prismant Incidence data of BTAR and sternum fracture in 5 years. [3]

From these data it could be interpreted that a sternum fracture is life-saving. On the other hand for heavy car crash deceased’s sometimes a positive coincidence between BTAR and sternum fractures was documented [3] (See Table 2.3) and therefore a sternum fracture could also been seen as live-threatening.

Author; year BTAR + SF + Mortality (%) Swan et al; 2001 36 0 42 Williams et al; 1994 90 25 100 Arajarvi et al; 1989 140 34 100 Arajarvi et al; 1989 98 24 100 Kram et al; 1989 12 0 0 Stark et al; 1987 49 0 ? Marsh et al; 1976 5 0 0 Zeldenrust et al; 1962 88 1 100 Total: 518 84 (16.2 %)

Table 2.3, incidence BTAR, sternum fracture and mortality. [3]

At the New Jersey medical school, the UMDNJ Crash Injury Research & Engineering Network Center [10] a positive correlation has been shown between no BTAR and sternum fracture. This leads sometimes to survival, 8 against 20 % (See Table 2.4) [10]. Furthermore, for all major thoracic non aorta injury thoracic trauma patients the mortality from other thoracic and head injuries does not get higher than 60 % (See table 2.4) [10], which are significantly lower than the 85 %, which is the averaged mortality of BTAR. From these data it can not be distinguished whether a sternum fracture is life-saving or life-threatening.

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Table 2.4 [10],

Head and thoracic injuries associated with non aorta injuries by outcome. Knobloch et al. [11] showed that sternum fractures occur most often in accidents with old cars to seat belted drivers without any airbag often with concomitant spinal injuries. (See Table 2.6 and 2.7). The large control group (42.055 patients archived from 1985 till 2004) shows that sternum fracture occurs mostly in car passengers as well as absolutely (251 cases) as relatively (0.81 %). A significant difference between airbag and seat belted drivers (91 % against 13 %, with 18 % airbag malfunction) demonstrates that an airbag decreases sternum fracture for heavy car crash victims. From these data, it could be questioned whether an airbag could be dangerous or save.

Table 2.5, Sternum fracture against different vehicle types. [11]

Vehicle type. Sternum Fractures (%). Car. 251 of 31,183 patients (0.81 %). Motorbike. 5 of 2,633 patients (0.19%). Truck. 4 of 3,258 patients (0.11%). Bicycle. 6 of 4,971 patients (0.12 %). Total. 267 of 42,055 patients (0.64%).

Table 2.6, Relative from total sternum fracture (%) against car characteristic. [11]

Experimental studies. In a cadaver research performed by Hardy et al. [2] it is concluded that the aorta is characterized by a nonlinear stress-strain response. Furthermore, the aorta fails in the transverse direction and the intima (inner layer of the aorta) fails before the media or adventia. The transverse BTAR obtained by controlled experiments is an affirmative of the shearing due to deceleration loads mechanism by Sevitt et al. [6]. Ruptures at the most wide part of the aorta, which have not been found, would have predicted the hemodynamic mechanism, because aorta explosion, due to high arterial pressure would occur at the most widely part. They also conclude that thoracic deformation is required for traumatic aortic injuries, but whole body acceleration is not and that loading of the aorta via the ligamentum arteriosum is not required, but may contribute to traumatic aortic injuries [2]. This last conclusion could be an affirmative for the osseous pinch theory by Crass et al. [9] instead of

Car characteristic. Relative from total sternum fracture (%). Seat belted drivers. 91 %. Airbag. 13 %, 18 % airbag malfunction.

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the shearing-due-to-deceleration-loads theory by Sevitt et al. [6]. A general conclusion from Hardy et al. is that cadaver testing is beginning to shed light on the potential mechanisms of traumatic aortic ruptures, but further investigation is required. Bass et al. [12] performed experiments with the goal of making a new element for finite element modeling that is characteristic for the aortic wall during an impact. (See Figure 2.9) They criticize the material model used in a computer study by Richens et al. [5] as incapable of simulating aortic injuries and uses more complex visco-elastic material models. The age dependency of aortic tissue has also been researched and young people have relatively more risk of having aortic injuries than old people and therefore a ‘young’ population would be more interesting than an ‘old’ population regarding to material properties research. For further experimental work young aorta tissue has to be used as a worst case scenario, because this tissue is more likely to rupture. Pressurization experiments in-vitro (extruded from the cadaver) and in-situ (still in the whole cadaver) (See Figure 2.9) showed no huge differences in number of obtained aortic injuries, concluding that the effect of tissues among the aorta seems to be non significant for this application and therefore implementation of external boundary conditions is not necessary to simulate the effect of tissues among the aorta. If the shearing mechanism at the isthmus is the most important mechanism, these pressurization tests could be considered as meaningless if high pressure is not necessary for the shearing mechanism.

Figure 2.9, bi–axial experiments and pressurization tests as an important background for numerical simulations. [12]

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A recent (September 2006) experimental cadaveric study by Baqué et al. [4] showed that a significant relative motion of the isthmus of the aorta and the heart has been found during sudden deceleration of the thorax (See Figure 2.10). Accelerometers have been used to prove this and a difference of 17 % in acceleration (Δa) was found within a range of 5-25%. This corresponds with a difference in displacement during time difference (Δt) of approximately: ½ΔaΔt2 = 0.5•50•0.025•0.025 m =0.016 m The free fall of the total body during brutal deceleration is approximately: ½ΔaΔt2 = 0.5•200•0.08•0.08 m =0.64 m The relative difference in displacement and total body free fall is 0.016/0.64*100=2.5 %. Baqué et al. notes that the osseous pinch theory is improbable because of the anatomic protection of the isthmus in the vertebral column in the modal anatomy of the mediastinum. A significant difference in deceleration between spine and sternum, which could be an indication for the osseous pinch theory, was observed only in 6 of 19 cases. This study provides for the first time physical demonstration and quantification of the stretching of the isthmus, leading to a computerized model of BTAR. Miniaturization of accelerometers together with hemodynamic measurements could lead to a better understanding of chest trauma. Note that the accelerometers used measure absolute values of acceleration vectors and that therefore the possibility of the heart moving side wards due to frontal deceleration can not be excluded. This possibility will not be investigated in this project.

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Figure 2.10, Study Baqué Exp. Set –up, cadaver accelerometer position

and a graph to show one result. [4] Numerical simulation studies: Two numerical simulation studies have been found, which both study BTAR by compressing the model, because compression was seen as the main cause of BTAR. Richens et al. [5] researched the feasibility of a finite element model predicting major vessel injury. The material models being used are linear elastic and isotropic which, as has been mentioned in the previous section, is characterized by Bass et al. [12] as incapable of simulating aortic injuries. Richens et al. [5] claims that the relative importance of deceleration, acceleration and crushing injuries in the etiology of BTAR remains uncertain. Two different models were used in this study. One model has relatively high detail describing heart, aorta, spine, blood, ligamentum arteriosum, pulmonary vessels and bronchi. The other model consists of a low detail impact thorax model which is used to provide boundary conditions for the high detail

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model. (See Figure 2.11) This procedure will also be useful regarding to the study described in this report to obtain a quick and stabile simulation. Another interesting thing about this study of Richens et al. is that it speculates about the influence of the position within the cardiac and pulmonary cycle during the crash towards the occurring of aortic injuries and the futuristic idea of controlling the safety applications dependent of the cardiac en pulmonary cycle of the individual passenger. The article’s main conclusion is that analysis of the response of the finite element heart – aorta model during blunt thoracic trauma demonstrates its potential for predicting major vessel injury.

Figure 2.11, Low detail Thorax model (left) and results from high detail heart aorta model (Right). [5]

In another computer model study by Shah et al. [13], the material properties were assumed linear elastic and isotropic as well as in the study by Richens et al. In this study the heart, lungs, rib cage, sternum, spine, diaphragm, major blood vessels and intercostal muscles, ligamentum arteriosum, blood, et cetera and highly detailed aorta model (See figure 2.12) have been modeled to predict aortic rupture due to impact loading. The computer model output indicated that the aortic isthmus was the most likely site of aortic rupture regardless of impact direction. In medical practice, the same kind of rupture is found uninfluenced by the impact direction. For frontal impact, the sternum pushed the ascending aorta posterior, causing direct compression of ascending aorta. The heart moves sideward, due to the frontal compression. (See figure 2.12) In this study compression was seen as the cause of BTAR. Stretching of the aorta would also be possible if the heart moves to the side and also torsion forces could lead to high values of von Misses stress. The model used is more detailed than the HUMOS 2 model. This model could only be validated on a global scene due to lack of experimental data. For further research validation of this model at a lower level of detail is recommended by the authors for obtaining a realistic simulation.

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Figure 2.12, Cross-sectional kinematics of the thorax simulation in frontal direction. Anatomical overview (above right), compressed simulation (above left) and highly detailed

aorta model (down). [13]

Discussion. Brink et al. [3] has formulated a hypothesis to reason the controversial negative coincidence between BTAR and sternum fracture (See Table 2.2) by survivors which is stated below: A sternum fracture (as a sign of ventral compression) significantly reduces the change of simultaneous occurrence of BTAR. This hypothesis is controversial because statistical data can not distinguish whether a sternum fracture is life-saving or life-threatening. Because inertial, elastic and viscous forces play a key role in thoracic injury biomechanics (See Figure 2.13) biomechanical research could confirm or exclude this hypothesis. This hypothesis only makes sense if the shearing due to deceleration loads mechanism at the istmus plays a key role in occurring BTAR. Baque et al. [4] provides for the first time physical demonstration and quantification of the stretching of the isthmus, leading to a computerized model of BTAR

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(See Figure 2.10). A recent cadaver experiment study by Hardy et al. [2] obtained transverse BTAR, which could demonstrate the shearing mechanism by Sevitt et al. [6]. Nevertheless, the same cadaver experiment study by Hardy et al. concluded that thoracic deformation is required for traumatic aortic injuries, but whole body acceleration is not, which is opposite due to the mechanism by Sevitt et al. [6], where deceleration loads cause BTAR. Richens et al. [5] concludes that a FE model is able to predict major vessel injuries with linear isotropic material model used for the aortic wall, which is the same kind of material model as used in the aortic wall of the HUMOS 2 model. Although Bass et al. [12] criticizes this isotropic, linear elastic material model as incapable of simulating BTAR and the non linear stress strain response for aortic tissue found by Hardy et al. [2], this Master project gives no attention to material- models and parameters. The isotropic linear elastic elements in the HUMOS 2 model could be a first step towards a more realistic numerical model with traumatic visco-elastic aortic wall elements from Bass et al. [12]. Because the appearance of a sternum fracture seems to be associated with the absence of an airbag, the approach chosen is to simulate three cases. These are the hub, airbag and safety belt application. Less von Misses stress in the safety belt application could be an indication that the compression during deceleration hypothesis could make sense.

Figure 2.13, Relative importance of elastic, viscous and inertial forces play a key role in thoracic injury biomechanics.

3. Methods

Introduction: This feasibility study questions whether a Finite Element Model (FEM) is able to research the ventral compression during deceleration hypothesis (See section 1 and 2). One example of a model that could be used to achieve this goal could be a part of the HUMOS 2 model (See Figure 3.1), which, because of its design, should be able to simulate the whole body during a car crash with enough level of detail to analyze the effect on the organs [14].

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In this chapter first the global model structure is discussed. Then the research goals and approach are given. Subsequently the application of crash specific boundary conditions as an input for the simulation are demonstrated for three types of crashing (airbag, safety belt and hub) and the use of von Misses as output. A method is considered for modifying this HUMOS 2 model to analyze the shearing-due-to-deceleration-mechanism at the isthmus. Finally, in a subsection, details of the FE model, the material models used, and the contact definitions are described.

Fig. 3.1, Global Overview of the HUMOS 2 model.

Global model structure: The HUMOS 2 model [14] runs in the program MADYMO [15]. This model combines multi body dynamics and finite element methods. The hard skeleton is modeled by rigid bodies and soft tissues such as heart, aorta and connective tissue are described by deformable finite elements. The figure below shows the interaction procedure.

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MULTIBODY MODULE (Bones of the HUMOS 2 model).

FINITE ELEMENT MODULE (Cartilage, Compact bone, Spongy bone and surrounding soft tissue)

Fig 3.2, The Finite element module, multi-body module interaction. [15] In the multi body dynamics the Newton Euler equation (See Eq.3.1) that describes the movement of the rigid bodies is taken as a point of departure. In equation 3.1, m represents mass, r the displacement and F the applied force for each body i. These rigid bodies are connected to each other by joints. The human joint connects the total model to the reference space. Each vertebra is modeled by a rigid body and a closed chain between the human joint and the sternum joint is necessary in the model. Eq. 3.1 The finite element modeling describes the deformable structures by means of an iterative loop of three differential equations (See Figure 3.3), which solution can not analytically be obtained. Therefore, a numerical procedure by a computer has to be performed, which estimates the exact solution. The finite element method is introduced in 1960 by Clough et al., which today is a commonly used tool for model based predictions of engineering problems.

Fig.3.3, the Finite element equations are looped and solved by a numerical procedure. Research goals and approach: To know which components of the HUMOS 2 model are essential and to obtain a quick and stable simulation, the approach chosen was to begin with a minimal amount of components and to increase the number of components until biofidelic results were obtained. The ventral compression during deceleration hypothesis could only make sense if the shearing-due-to-

Forces and torques Supports and contacts

iii Frm =&&

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deceleration-loads mechanism at the isthmus plays a key role. Therefore, high values of von Misses stresses at the isthmus location of the aorta model could be a sign that a biofidelic simulation has been computed. A factor that complicated this procedure was the implementation of crash specific boundary conditions, which let the model simulate a crash simulation. Because it was uncertain if unsatisfactory results were caused by wrongly chosen components of the HUMOS 2 model and/or unrealistically application of boundary conditions, more procedures were tried. In this report, however, only one procedure for applying boundary conditions is described. A model with lower level of detail was used to create these boundary conditions. The same components of the HUMOS 2 model, which indicates the shearing-due–to-deceleration mechanism, should be used to investigate the predictive value of a FEM that can exclude or confirm the ventral hypothesis. If lower values of the von Misses stress are found for a higher value of compression, this model could be feasible to test the compression during deceleration hypothesis. Input of the simulations: To let the model simulate a crash, appropriate boundary conditions are necessary. The effect of deceleration and compression is essential for conforming or excluding the compression during deceleration hypothesis. Therefore, at least two variables had to be used to prescribe the motion of the model as an input for the simulation. The body acceleration at spine T3 in the x direction and the deflection (compression of the thorax) between spine T3 and sternum were taken because the ascending aorta is located between spine and sternum in the model. Because deceleration is a relative clause, negative acceleration of the vehicle could be simulated by applying positive acceleration to the Human Body Model (HBM). (See figure 3.4) Vehicle deceleration as an input gives body deceleration as an output for the low detail model, which is an input for the simulations with the high detail model.

Figure 3.4, Definition acceleration field (left) parallel to the x axis and vehicle deceleration

applied to the HBM in all three simulations (right). In the creation of boundary conditions, it has been tried to obtain the difference of airbag, safety belt and hub compression and deceleration with a model having a lower level of detail than the HUMOS 2 model, the MADYMO facet human model. If the hub, airbag and safety belt conditions would be applied to the original HUMOS 2 model, numerical problems would be more likely to occur. (See Figures 3.5, 3.6 and 3.7)

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Fig. 3.5, The begin (left) and end (right) animation of the airbag simulation with the

MADYMO human facet model.

Fig. 3.6, The begin (left) and end (right) animation of the hub simulation with the MADYMO

human facet model.

Fig. 3.7, The begin (left) and end (right) animation of the safety belt simulation with the

MADYMO human facet model.

Fig. 3.8, Body Deceleration at spine T3 against Time for the airbag, hub and safety belt

model.

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It has to be mentioned that the acceleration signals were processed such that they were relative to the reference space by deducting the vehicle motion (See Figure 3.4) from the simulation output. The safety belt model (See Figure 3.8) firstly decelerates, following by the airbag model and finally the hub model. This could be realistic, because it should take more time to get to the hub, then to get to the exploded airbag or to get into the safety belts. The higher derivatives of the hub and airbag model signals seem to indicate the times where the human body model contacts the airbag or hub. If the airbag and safety belt model signals are examined then it could be noticed that the human body contacts the airbag faster than the hub, which is realistic. The airbag model deceleration seems to be the most flat during the crash.

The Safety belt model deflection has the highest maximum, followed by the hub, which maximum deflection is at the end of the simulation (See Figure 3.9). It could be that a higher value of deflection would be obtained if a longer simulation time had been taken. The airbag model has the lowest minimal deflection, which should be obvious, because this safety equipment has been designed to obtain lower deflection during the crash.

Fig. 3.9, Deflection between spine T3 and sternum for the airbag, hub and safety belt

application. Output of the simulations: To simulate the shearing-due-to-deceleration-loads mechanism at the isthmus, the model used should generate shear stress as output. A problem for this shear stress mechanism is that it exists in the 3D domain, while the aortic wall is modeled by 2D shell elements. Von Misses stress is computed by calculating main stresses in the direction where high values of stress occur in the 2 D strain field (principle stress). Therefore, high values of von Misses stress could indicate high values of shear stress. A reason for using von Misses stress in stead of any other stress is that von Misses stress is no vector, but an absolute scalar value. This makes it easier to find out if the level of detail of the model used and boundary conditions were appropriate to research BTAR. Using von Misses stress could also be interpreted as a limitation, because it represents an average of stress in many directions

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HUMOS 2 model modification: To make a biofidelic simulation of the shearing-due-to-deceleration-load mechanism at the isthmus, a model modification is implemented. The heart has the ability to change in volume (contraction and relaxation) and therefore, the heart and the ascending aorta are not attached to the rest of the thorax. Because a difference in deceleration between sternum and the descending aorta was seen in experiments (See previous section), this phenomenon has to be present in the model. In the HUMOS 2 model, the aorta is modeled as a whole part (See figure 3.10). Because the descending aorta is attached to the whole body and the ascending aorta not, it was necessary to divide the aorta into two parts (ascending and descending) of which the elements still had the same parameters as the whole aorta. Between both parts a gap of 0.001 m was created. A contact definition connects both ascending and descending aorta. The descending aorta is a master of the ascending aorta and therefore, stretching at the isthmus could be possible. By prescribing deceleration to the rest of the thorax and abdomen, the deceleration of the heart and aorta is calculated by the model itself. (See Figure 3.11) The isthmus region was defined at the part of the descending aorta, which bordered the ascending aorta.

Fig. 3.10, original HUMOS 2 model, left whole and right inside of the model where heart and

aorta are visible.

Fig. 3.11, Rest of the thorax and abdomen (left) and heart and aorta (right) as two separated

systems. FE model elements in detail: The whole thorax and abdomen part of the HUMOS2 model was used to analyze the difference between linear elastic 2D shell elements or (visco-)elastic 3D volume elements. (See Table 3.1) To let the separated elements crash to each other contact definitions were defined in the HUMOS 2 model. Because structures are deformable, it has to be decided which structure is the master and which the slave. The slave surface adjusts itself to the master surface as shown in Figure 3.12. In the HUMOS 2 model, for the same thoracic soft tissues, the heart has other contact definitions than the aorta. Separating the aorta into descending and ascending aorta meant that contact definitions had to be arranged and it has

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been chosen that the ascending aorta has the same contact definitions as the heart and the descending aorta as the aorta. (See Table 3.4)

Linear Visco elastic: Isotropic linear elastic: Isotropic elasto plastic:2D Shell Elements

Ascending / descending Aorta, Thorax ribcage muscle, Lungs.

Sternum.

3D Volume Elements

Heart, Thoracic connective tissue, Superficial muscle.

Table 3.1, Material models of certain components of the model used.

Structure: Thick: E: ρ: ν: K: GDL: G∞: σy: β:

Ascending and descending Aorta

1.6 10-3 2.5 106 5000 0.4 - - - -

Heart - - 1000 - 6.6 104 3.3 103 1.4 104 - 1000 Thoracic superficial muscle

- - 1000 - 2.5 104 6.9 104 1.55 105 - 1000

Thoracic connective tissue

- - 1000 - 2.5 105 6.9 104 1.55 105 - 1000

Sternum compact bone.

0.82 10-3 1.9 1010 6000 0.3 - - - 7.3 107 -

Table 3.2, Material Parameters of certain components of the model used.

Parameter symbol: Definition: Entity: Thick Thickness of the element m E Elasticity modulus Pa ρ Density Kg/m3 ν Poisson’ s ratio - K Yield stress Pa GDL Additional dynamic shear change mode I Pa G∞ Quasi-static shear modulus Pa σy Yield stress Pa β Decay constant stress relaxation s-1

Table 3.3, Definition of material parameters.

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Master: Slave:

Heart and ascending aorta. Thoracic connective tissue.

Heart and ascending aorta. Ribcage surface.

Heart and ascending aorta. Lungs.

Heart and ascending aorta. Thorax Ribcage muscle inner.

Diaphragm. Heart and ascending aorta.

Diaphragm. Descending aorta.

Thorax Ribcage muscle inner. Descending aorta.

Lungs. Descending aorta.

Thoracic connective tissue. Descending aorta.

Descending aorta. Ascending aorta.

Table 3.4, Difference in contact definitions for the two systems.

Fig. 3.12. Master Slave interaction of finite elements[15].

Discussion: The airbag, safety belt and hub model have created boundary conditions, which let the model simulate a crash situation and seem to be appropriate to be used for conforming or excluding the compression during deceleration hypothesis. They are considered to be appropriate because the deflection and deceleration are different from each other like expected (See input of the simulation). The safety belt model contained no hub and the airbag model contained no safety belt. These two facts, which could be interpreted as making the boundary conditions less realistic, were assumed to be of no significance. The abdomen and thorax part of the HUMOS 2 model are isolated from the rest of the HUMOS 2 model. The aorta model has been divided into two separated parts to simulate the shearing-due-to-deceleration-loads mechanism and certain contact definitions with these two aorta parts have been implemented without model validation.

Slave surface

Master surface

Motion of boxes:

F

Slave surface

Master surface

F

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4. Results. Introduction: The ventral compression during deceleration hypothesis only makes sense if the shearing-due-to-deceleration-loads mechanism at the isthmus plays a key role (See section 1 and 2). The approach chosen was to vary deceleration and compression as an input and to use Von Misses stress as an output and to let the model calculate the heart and ascending aorta deceleration by themselves (See section 3). As a reference, the image at t=0.000 s is captured, which is for all kinds of boundary conditions the same and is plotted below (See figure 4.1). First instability, as a numerical problem, is given in this section. Then, Von Misses stress results are shown for each case (airbag, safety belt and hub).

Fig. 4.1, Reference for simulations with all kinds of boundary conditions at t=0.000 s.

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Fig. 4.2, The last image before the simulation turned instable with boundary conditions from

the hub model.

Fig. 4.3, The last image before the simulation stopped due to instability.

With boundary conditions from the safety belt model.

Instability as a numerical problem: Instability as a numerical problem was found in two of the three simulations. The model with boundary conditions from the airbag application was able to calculate the entire simulation time without having instability problems, while the hub application proceeded until 68 % of the simulation time. The safety belt application had even only a stable simulation until 32 %

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of the simulation time. For the hub simulation the last image before stability problems made the simulation unable to run further is given and the position of some finite elements of the heart appear to be outside the ribcage (See figure 4.2). This appearance is not seen in the last image of the safety belt application simulation (See figure 4.3). The airbag model: To show that the heart and ascending aorta are moving away from the descending aorta in the airbag simulation, one simulation image at t=0.112 s is visualized, which difference in position of the heart relative to the descending aorta was the greatest obtained during the simulation. This difference in position leads to realistic values of von Misses stress (See figure 4.4). The maximum value of von Misses stress was 7.44 105. Von Misses stress higher than 1.0 104 was found at the isthmus location at t=0.112 s. The Highest values of von Misses stresses were found at the ascending aorta.

Fig. 4.4, simulation at t=0.112 s with boundary conditions from the airbag model.

The hub model: For the hub model, the last image is given before the positions of finite elements of the heart go into the ribcage surface, which is not realistic. Although this unrealistic appearance, realistic values of von Misses stress are found at the isthmus region of the aorta model, before the model gets unstable. The highest values of von Misses stress are found at the ascending aorta. The heart moves to the sternum as is visible in figure 4.5.

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Fig. 4.5, Simulation with boundary conditions from the hub model at t = 0.132 s.

The safety belt model: Also for the safety belt application, the heart moves to the sternum and realistic values of Von Misses are being found near the isthmus region (See figure 4.6). Again, the highest values of von Misses stress are at the ascending aorta. A lower value is found at the isthmus region compared to the two other boundary conditions, however, due to instability problems the simulation stopped at t=0.070 s, and therefore nothing can be discussed about this fact.

Fig. 4.6, Simulation with boundary conditions from the safety belt model at t=0.070 s.

The results from the simulation with all kinds of boundary are comparable to the results computed by Richens et al [5], where maximum von Misses stresses of approximately 1*106 were found at the isthmus region. Therefore, the feasibility of this finite element model with

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the proposed method of applying boundary conditions to predict the occurrence of BTAR has been indicated.

5. Conclusions and discussions.

For the same model and method of applying boundary conditions, instability was found for two of the three simulations. The deceleration of the airbag model is more flat, which could be a reason why this is the only simulation performed entirely stabile. For the simulation with boundary conditions from the hub model positions of heart finite elements came outside the ribcage, which could indicate that the severity of the hub boundary conditions maybe is the reason for the instability. Realistic values of Von Misses stress are found at the isthmus for all three simulations of the aorta model and the simulation started with zero von Misses stress (See section 4). Therefore, this method of applying boundary conditions could be useful for further research. Because the highest values of von Misses stresses are at the ascending aorta and not at the isthmus, where BTARs are found most often, the current model with the current method of applying boundary conditions, could not give the feasibility of a finite element model simulating the shearing-due-to-deceleration-loads mechanism. The human airbag hypothesis, where compression significantly reduces the risk of simultaneous occurrence of BTAR by survivors only could make sense if the shearing-due-to-deceleration-loads mechanism at the isthmus is recognized for playing a key role. Therefore the simulations in this report were not able to confirm or exclude this hypothesis. For further research, stabile simulations with the right level of detail to simulate the injury mechanism of BTAR have to be performed.

6. Future work.

The current aorta model of the HUMOS 2 model and method of applying boundary conditions has many limitations. For further numerical work these limitations have to be solved:

- First, the material model of the aorta model in HUMOS 2 is linear elastic and isotropic, which is characterized as incapable of simulating BTAR (See section 2) and for further research another material model could be implemented.

- The aorta model in the HUMOS 2 model has a low level of detail. For a similar BTAR study, an aorta model with higher level of detail is used. (See section 2) At this level of detail model validation must be possible to obtain a more biofidelic simulation.

- Numerical problems have to be solved. The hub model creates boundary conditions with high order derivatives, which could be the reason for the instability phenomena. Further research could exclude whether a stable simulation with these hub boundary conditions is possible.

- The shearing mechanism, which has to be obtained in the model to explain or reject the human airbag hypothesis exists in a three dimensional domain and the aorta model used consists of 2D shell elements (See section 3). To simulate the shearing mechanism, the separated layers should be modeled with 2 D shell

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elements. Two different aorta wall layers (intima and adventia) should be mechanically characterized by Bi- axial experiments to obtain a validated 3 D model by 2 D shell elements meshed parallel to each other. The third layer (media) consist only of endothelial cells and could probably mechanically be neglected.

- The current method of applying boundary conditions by a model with a lower level of detail to obtain a stabile simulation has the disadvantage that the inflection of the spinal column is not present in the simulation with the high level of detail model. For further research, boundary conditions should be applied such that the spinal column flexes.

- Kroell [16] and Bouquet [17] tests could be used in stead of the airbag, hub and safety belt models. These injury tolerance models have been validated with human cadaver material and sternum fractures were observed after these tests. Deceleration and compression should both be obtained from the same model, therefore these Kroell and Bouquet tests should be implemented in the HUMOS 2 model and chair model, which decelerate.

- Dividing the current model into two systems as a ‘proper’ method to investigate the injury mechanism of BTAR has to be validated. Contact definitions and material parameters could be chosen differently for the two aorta models (ascending and descending). Fall experiments with human cadaver material could be compared to simulated falls with the modified HUMOS 2 model. Fitting material parameters and correcting contact definitions could lead to higher biofidelity (match between simulation and experiment).

- During the cardiac cycle the mass of the heart and aorta change periodically. This phenomenon is not present in the HUMOS 2 model. Simulations with maximum and minimum heart and aorta mass could be implemented.

- The current aorta model is hollow and therefore, blood and wall mass is modeled in the 2D shell elements. An average car crash takes about 220 ms. and during this crash blood is able to flow and computational fluid dynamics could be a necessary to obtain a simulation with the right level of detail to research the occurrence of BTAR.

REFERENCES: [1]: Alameda County Medical Centre/ Highland General Hospital. [2]: Warren N. Hardy, Chirag S. Shah, James M. Kopacz, and King H. Yang, Chris A. van Ee, Richard Morgan and Kennerly Digges, Study of Potential Mechanisms of Traumatic Rupture of the Aorta Using In Situ Experiments, , Stapp Car Crash Journal, Vol. 50 (November 2006). [3]: Prof. Dr. P. R. G. Brink, Drs. G.Raven, Drs. J.Scholten, Sternum fracture; the human airbag? International Proceedings of 7th European Trauma Congress Ljubljana, 2006:53-7 [4]: Baqué, Patrick MD, PhD; Serre, Thierry PhD; Cheynel, Nicolas MD; Arnoux, Pierre-Jean PhD; Thollon, Lionel PhD; Behr, Michel PhD; Masson, Catherine PhD; Delotte Jérôme MD; Berdah, Stéphane-Victor MD; Brunet, Christian MD, An experimental cadaveric study for a better understanding of blunts traumatic aortic rupture, The journal of Trauma, volume 61(3), September 2006, pp 586-591. [5]: David Richens; Mark Field; Shahrul Hashim; Micheal Neale; Charles Oakley A finite element model of blunt traumatic aortic rupture:, European journal of cardio-thoracic surgery, January 2004, pp 1039-1047. [6]: Sevitt S., The mechanisms of traumatic rupture of the thoracic aorta., Br. J, Surg., 1977.

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[7]: Oppenheim F., Muench, Gibt es seine spontanruptur der gesunden aorta und wie kommet es zustande?, Med. Woschenschr. 1918;65:1234-1237. [8]: Klotz, O., Simpson W., Spontaneous rupture of the aorta., Am. J. Med Sci. 1932;184:455-473. [9]: Jeffrey R. Crass, MD, Alan M. Cohen, MD, Antonino O. Motta, MD, Joseph F. Tomashesfski, Jr, MD, Srnest J, Wiesen, BSEE, A proposed New Mechanism of Traumatic Aortic rupture: the Osseous Pinch,. [10]: Jonh H. Siegel, MD, FACS, FCCM; Joyce A. Smith, MS; Nadegda Tenenbaum, MD; Laurie McCammon; Shabana Q. Siddiqi, MD New Jersey Medical School: UMDNJ Crash Injury Research & Engineering Network Center, Fatal Versus Potentially Survivable Motor Vehicle Crash (MVC) Aortic Injuries (AI): The ratio of deceleration Energy to change in Velocity on Impact and the presence of Associated Injuries as Determinants of Outcome.. [11]: Kasrten Knobloch, MD, Carl Haasper, MD, Christian Probst, MD, Christian Krettek, MD, PhD, Dietmar Otte, PhD, and Martinus Richter, MD, PhD.Sebastian Wagner, MS, Sternal Fractures Occur Most Often in Old Cars to Seat-Belted Drivers without Any Airbag Often with Concomitant Spinal Injuries: Clinical findings and Technical Collision Variables Among 42,055 Crash Victims. [12]: C. R. Bass, K. Darvish, B. Bush, J. R. Crandall, S. C. M. Srinivasan, C. Tribble, S. Fiser, L. Tournet, J. C. Evans, J. Patrie, and C. Wang Material properties for modeling traumatic aortic rupture:, Stapp Car Crash Journal, Vol. 45 November 2001. [13]: Chirag S. Shah, Kinh, H. Yang, Warren N., Hardy, H. Kevin Wang and Albert I. King Development of a Computer Model to Predict Aortic Rupture Due to Impact Loading.. November 2001, pp. 161-182. [14]: MADYMO Human Models Manual Version 6.3, December 2005. [15]: MADYMO Theory Manuel Version 6.3, December 2005. [16]: Kroell, C. K. , Schneider, D. C., Nahum, M., Impact Tolerance and Response of the Human Thorax II. 1974 [17]: Bouquet, R., Ramet, M., Bermond, F., Cesari, D, “Thorax and Pelvis Human Response to Impact”, Proceedings of the 14th International Technical Conference on the Enhanced Safety of Vehicles. pp. 100-09, 1994.