Eindhoven University of Technology

MASTER

Modelling of the human shoulder complex in impact conditions

van Hassel, E.

Award date:1998

Link to publication

DisclaimerThis document contains a student thesis (bachelor's or master's), as authored by a student at Eindhoven University of Technology. Studenttheses are made available in the TU/e repository upon obtaining the required degree. The grade received is not published on the documentas presented in the repository. The required complexity or quality of research of student theses may vary by program, and the requiredminimum study period may vary in duration.

General rightsCopyright and moral rights for the publications made accessible in the public portal are retained by the authors and/or other copyright ownersand it is a condition of accessing publications that users recognise and abide by the legal requirements associated with these rights.

• Users may download and print one copy of any publication from the public portal for the purpose of private study or research. • You may not further distribute the material or use it for any profit-making activity or commercial gain

Modelling of the Human Shoulder Complex in Impact Conditions

Edwin van Hassel

WFW-report 98.024

Master's thesis, June 1998

Study of Biomechanical Engineering Department of Mechanical Engineering Eindnoven University of Technoiogy (TÜE) Eindhoven, The Netherlands

Performed at: Dutch Organisation for Applied Scientific Research TNO Road-Vehicles Research Institute Crash-Safety Research Centre Delft, The Netherlands

S upervi sors : Prof. Dr. Ir. J.S.H.M. Wismans (TNORUE) Prof. Dr. Ir. D.H. van Campen (TUE) Dr. Ir. R. Happee (TNOíTUE) Dr. Ir. P.H.M. Bovendeerd (TUE)

2

Contents

Summary . . . . . . . . . . . . . . . . . . . . . 1 Introduction . . . . . . . . . . . .

1.1 Introduction . . . . . . . . . . . 1.1.1 Crash Safety Research . . . . . . 1.1.2 MathematicalModelling . . . . .

1.2 Shoulder and Upper Extremity Research . . 1.3 Available Models . . . . . . . . . .

P . 3.1 Crash Dummy Shoulders . . . . . 1.3.2 Mathematical Models . . . . . .

1.4 Goal of this Study . . . . . . . . . 1.5 Model Requirements and modelling methods 1.6 Preview . . . . . . . . . . . . .

. . . . . . . . . 7

. . . . . . . . . 7

. . . . . . . . . 7

. . . . . . . . . 8

. . . . . . . . . 8

. . . . . . . . . 9

. . . . . . . . . 9

. . . . . . . . . 10

. . . . . . . . . 12

. . . . . . . . . 12

. . . . . . . . . 13

2 Anatomy of the Human Shoulder Complex and Thorax . . . . 15

3 The Shoulder Complex and Torso Model e . o

3.1 Spine . . . . . . . . . . . . . . . . . 3.2 Thorax . o a o o o o o o o o

3.3 Shoulder . . . . . . . . . . . . . . . 3.3.1 Scapulothoracic Connection . . . . . . . 3.3.2 Clavicular Connection. AC-. SC- and GH-joints 3.3.3Arm . . . . . . . . . . . . . . . 3.3.4 Ellipsoids . . . . . . . . . . . . . 3.3.5 Muscles . . . . . . . . . . . . . .

3.4 TotalModel . . . . . . . . . . . . . .

. . . . . . 17

. . . . . . 17

. . . . . . 18

. . . . . . 20

. . . . . . 20

. . . . . . 20

. . . . . . 22

. . . . . . 23

. . . . . . 24

. . . . . . 25

4 Validation O f the Model . . . . . . . . . . . . . . . . . 27

5 Sensitivity Analysis . . . . . . . . . . . . . . . . . . 36

6 Interaction with Side.Airbags . . . . . . . . . . . . . . . 35

7 Conclusions . . . . . . . . . . . . . . . . . . . . . 37

8 Recommendations . . . . . . . . . . . . . . . . . . . 39

References . . . . . . . . . . . . . . . . . . . . . . . 41

Appendices . . . . . . . . . . . . . . . . . . . . . . . 45 A Images of available models . . . . . . . . . . . . . . . . 46 B Images of the model . . . . . . . . . . . . . . . . . . . C Scapulothoracic contact . . . . . . . . . . . . . . . . . 51 D Passive Muscle Characteristics . . . . . . . . . . . . . . . 52 E ModelValidation . . . . . . . . . . . . . . . . . . . . 53 F Sensitivity Analysis . . . . . . . . . . . . . . . . . . . 57 G Side Impact Application . . . . . . . . . . . . . . . . . 77

49

3

4

Summary

Although safety measures like airbags reduce the risk of fatality in crashes, the risk of non-fatal injuries, often to the upper extremity and shoulder complex, can be increased. To investigate these injuries, more insight in the biomechanics of the shoulder is required.

This thesis describes the development of a multi-body model of the human shoulder compiex, suitabie Îor impact conditions. Tnis mociei is based on a shouidermociel by D e 3 University, and an existing human body model by TNO.

Lateral bending of the spine had to be enabled to simulate side impact. Torsion and frontal and lateral translation between vertebrae have been suppressed using very high stiffnesses, to which critical damping had to be added to prevent the occurrence of high-frequent vibrations. To be able to model thorax deformation, a more detailed thorax model has been implemented.

two points of the scapula, the angulus inferior and the trigonum spinae. A point restraint (a 3D- spring) pulls the scapula against the thorax, to maintain contact, which in real life is done by the muscles. For the three shoulder joints (sternoclavicular, acromioclavicular and glenohumeral) maximum rotation angles have been specified. The scapulothoracic gliding plane and the joints form a closed chain, giving two different load paths from the scapula to the thorax. Simulations of lateral cadaver sled tests, showed that the distance between the acromions of the shouldermodel did not decrease enough. To correct this, translation was allowed in the sternoclavicular and acromioclavicular joints, which resulted in a model response much closer to the cadaver tests.

The geometry of the clavicles, scapulae, arms and hands has been modelled with several ellipsoids. The range of motion of the glenohumeral joint is limited by the contact between the ellipsoids modelling the acromion and the humerus.

For each shoulder, 95 Hill-type muscle models have been attached to the correct bones, except those that in real life are attached to the ribs. Only passive muscle behaviour has been modelled, because active muscle behaviour may have little influence in short impact times.

The shoulder model has been validated for side impact, using lateral cadaver pendulum impact tests on the shoulder and thorax and lateral sled tests and for frontal impact using frontal cadaver pendulum impacts. These validation tests have not only been used for validation, but also for tuning of model parameters. It was not possible to fulfil all requirements at the same time, so the parameters have been chosen to give the best overall fit.

The validated model has been used for a sensitivity analysis. This analysis showed that some parameters of which no measurement values are available have a large influence on the model response. Parameters tnat need more measurement are the effective clavicle and scapula mass, translational stiffnesses of the AC- and SC-joints, thorax stiffness, and contact stiffnesses. For other test conditions then used in this thesis, also individual joint range of motion and stiffness data may be required.

The human shoulder- and torsomodel has been used for a simulation in which the human body model is seated in a car that is struck from the side by a barrier. This car has a side- airbag. This simulation also has been performed with an EuroSID-I crash-dummy model. The results of this simulation for both models showed much resemblance. But the shoulder of the human body model was pushed upwards by the inflating airbag. In the EuroSID, this movement is not possible.

FEM may be required, but this seems to be less urgent. To have enough data for the development of an improved model, more measurements of the mechanical properties of the shoulder and the thorax are required. For validation of an improved shoulder model, cadaver tests with more emphasis on the shoulder are necessary. The current model shows the global kinematics of the shoulder and the distinction between two load paths rather well.

The scapulothoracic contact has been modelled with contact between a thorax-ellipsoid and

Further model improvement may need FEM techniques for the thorax. Also for the shoulder

5

6

1 Introduction

1.1.1 Crash Safety Research

During the last decades the number of road vehicles has increased significantly. In 1956 there were 0.3 million road vehicles in The Netherlands. In the following decades this increased from 3.4 million in 1976 to 5.7 million in 1996 (CBS,1997). Till 1972 this increase in traffic lead to an increase in traffic deaths. After 1972 (3264 deaths), the number of traffic deaths decreased. In 1956 1628 people died as direct result of traffic accidents in The Netherlands, 2432 in 1976 and 1180 in 1996 (CBS, 1997). This decrease of deaths since 1972, in spite of an increase in traffic, is caused by successful injury reduction strategies, resulting in improved vehicle safety.

Vehicle safety can be divided in two categories: active and passive safety. Active safety comprises measures, to prevent an accident from happening (e.g. ABS). Passive safety or crash safety comprises the measures to reduce the injury if an accident occurs (e.g. seat belt).

Crash safety research aims to predict the injury caused in a certain accident. With this knowledge it is possible to control the accident, e.g. by changing the construction of the vehicle or by applying seat belts or airbags, thus reducing injury.

To predict the human biomechanical response to certain impacts, a model is needed. Five different types of models are being used: Human volunteers, human cadavers, living and dead animals, mechanical models and mathematical models (Wismans, 1994). Each of these models has its own (dis-)advantages.

Human volunteers are the most realistic model, but can only be used in low severity tests, and mounting of the instrumentation is difficult. Human cadavers can be used in higher severity tests and easily be instrumented, but the soft tissue is too stiff compared to living humans, there is no muscle tonus and most cadavers are old. Animals (living or dead) can be used in higher severity tests as humans, but "scaling" to humans is difficult. Mechanical models or crash dummies can be used in high severity tests and tbie momting of instamentation is easy, but the nechzfiica! behzviour of hnms nust be known to develop a dummy. Mathematical models are relatively cheap and can be used in all kinds of impacts, but knowledge of the human mechanical behaviour is required for developing the model.

In crash safety research usually volunteer, cadaver and animal tests are being used to gain insight in the human mechanical behaviour. With this knowledge mechanical and mathematical models are being developed. These models then are being used for impact tests. In these tests several parameters are measured (like accelerations and displacements). These measurements are being used to calculate injury criteria, that are supposed to be related to injuries occurred in humans. In this way it should be possible to predict injury from a certain impact.

7

1.1.2 Mathematical Modelling

A very important advantage of mathematical modelling is the low cost. After a mathematical human body model has been created, it can be used for different kinds of tests. Also mathematical models can be useful in the design process of (parts of) road vehicles, enabling to test the new product before it is completed. A problem in mathematical modelling is getting information about what geometric, kinematic (degrees of freedom) and dynamic (masses and stiffnesses) parameters should be used.

After a mathematical model kas been completed, vaiidaiion (cornparing io response requirements) is necessary, to show the biofidelity of the model. A good model is a model that shows good biofidelity in validation, using realistic parameters. Only a successful validation does not indicate the quality of the model, because each model can be fitted to have a response that satisfies the requirements, just by changing a few important Parameters, which is known as tuning.

Three types of mathematical modelling can be distinguished; lumped mass, multi- body and finite element (FEM) models. Lumped mass models are one-dimensional models, consisting of masses, connected by massless springs and dampers, and are in fact simple multi-body models. Multi-body models are two- or three-dimensional models, consisting of several rigid bodies, connected by different types of joints. Between two bodies elastic interactions and contacts can be specified. Also flexible bodies can be defined. A finite element model is a model of a continuum, divided in many small elements. Also contacts between two continua can be specified. FEM models are much more complicated and require much more computational time, but are also more detailed than multi-body models. Several computer programs have been developed for multi-body (DADS, CAL3D), for FEM (MARC, DYNA3D, PAMCRASH, RADIOSS), or both (MADYMO).

1.2 Shoulder and Upper Extremity Research

As described in 5 1.1.1 ~ during the last two decades the number of traffic deaths has decreased in spite of an increase in traffic, due to improvements to automobile design and the introduction of improved safety systems, such as front and side airbags. Although safety measures like airbags reduce the risk of fatality in crashes, the risk of non-fatal injuries can be increased (Bass,1997). In 540 crashes with steering wheel airbag deployment, 38% of the drivers sustained upper extremity injury (34% AIS- 1, and 4% AIS-2 or -3 (fracture)), caused by the airbag deployment (HueZke,1997). Also side airbags can cause AIS-2 or -3 injuries to the shoulder and upper extremity (KaZZieris,1997). Shoulder injuries further can be caused by the seat belt (Frumpton,1997). Although such injuries to the upper extremity are not fatal, they can cause significant long term morbidity and loss of employment.

The biomechanics of the shouldergirdle and upper extremity in impact are poorly understood. Most mechanical and mathematical models model the shoulder as just one, ore sometimes two or three joints, while it consists of three joints and a gliding plane (51.3, Chapter 2). To study shoulder injuries in more detail, more detailed modelling of the shoulder complex in impact is required to improve the understanding of its biomechanics in impact.

8

1.3 Available Models

1.3.1 Crash Dummy Shoulders

To get an idea how the human shouldergirdle is modelled in mechanical human body models (dummies), the shoulders of three types of dummies have been examined. These three dummies are the EuroSID-I, the Hybrid 111 and the Q3-child dummy. Chapter 2 gives a description of the human shoulder anatomy and degrees of freedom.

The EuroSD has been developed especially for side impact. This dummy has no lower arms or hands. In the EuroSID the upper arm (humerus)

is connected to a ‘clavicle’, which is connected to the spine. The clavicle-spine joint has only one degree of freedom; mainly posterior / anterior

rotation of the clavicle, with a vertical rotation axis, with some coupled translation. A spring (k = 10 Nm / rad) pulls the clavicle (length = 20 cm) backwards against a stop.

acromioclavicular joints) is a revolute joint. This joint has a horizontal rotation axis (lateral-medial) that enables flexion f extension of the humerus (forward / backward movement of the arm).

Due to some elastic deformation of the shoulder joint and the arm, a small range (< 10’) of abduction / adduction of the arm is possible (sideward movement of the arm). A larger range of motion might be unnecessary for side impact in which the arm is pushed against the torso.

There is no elevation / depression of the clavicle possible (lifting of the shoulders). Rotation of the humerus around its own axis is not possible, but because there are no lower arms, this does not seem to be a limitation.

specifications for the shoulder have been used

The humerus-clavicle joint (representing the human glenohumeral and

In the development phase of this dummy, lateral compression biofidelity

The Hybrid III is a dummy, merely for frontal impact. It is the most widely used dummy.

two rotations of the humerus are possible: flexion / extension (forward / backward movement of the arm) and abduction / adduction (sideward movement of the arm). There is no motion (or very little due to limited stiffness) possible of the glenoid.

Rotation of the humerus around its own axis is not possible in the shoulder joint, but it is possible just above the elbow joint, which results in the same degrees of freedom.

The Hybrid II dummy is able to lift its shoulders (elevation i depression of the clavicle). In the Hybrid Ill this is limited.

No biofidelity specifications for the shoulder have been used for this dummy.

The shoulder of the Hybrid III consists only of the glenohumeral joint. In this joint

9

Q3-child dummy

The Q3 dummy is a rather new dummy. This dummy represents a three year old child. The shoulder of the 43 dummy appears to be much more sophisticated as the other

dummies' shoulders e The humerus is connected to the shoulder with a ball and socket joint (glenohumeral

joint). Remarkable is that the ball is connected to the shoulder, and the socket to the humerus, while this is the other way in real life. In this joint two rotations are possible: flexion i extension (forward / backward movement of the arm) and abduction i adduction (sideward movement of the arm). Rotation of the humerus around its own axis is possible in a joint just below the glenohumeral joint.

The shoulder itself (with a shoulderblade connected to it) is connected to the spine with a beam and to the front of the thorax (sternum) with a clavicle. The beam to the spine is flexible (but still rather stiff, resultant stiffness in glenoid = lo4 Nm), just as the clavicle. The clavicle is mounted elasticly to the thorax. Thereby, some motion between the two sternoclavicular joints is possible.

EuroSID and Hybrid I11 shoulders.

shoulder have been used

Although this shouldergirdle model is not perfect yet, it seems to be better than both

In the development phase of this dummy, lateral biofidelity specifications for the

1.3.2 Mathematical Models

Currently, several mathematical (complete human body) models are available, that might be of interest. Images of some of these models can be found in appendix A.

The ergonomics software RAMSIS (1998) uses a complete human body model, with a shoulder consisting of two joints: the sternoclavicular and glenohumeral joints. The rotation angles of these (and other) joints can be entered, to put the human in a certain position.

The GEBOD (GEnerator of Body Data) is a human body multi-body model creating program, which gives main anthropometric information after some characteristic measures (e.g. weight or standing height) have been entered. This output can be given in several formats, like ATB (Ma, I995), MADYMO ( I 996) or CAL3D (Irwin, 1994). In this model the shoulder consists of only one joint, between the thorax and the upper arm.

The model by Ma et al. (1995) is the GEBOD-ATB 50* percentile male model. In this model, joint characteristics have been implemented. These characteristics are joint stiffnesses and joint sinus cones. Joint sinus cones represent the maximum flexion angle as a function of the direction in which this flexion takes place.

The model by Wayne State University (Huang, 1994a), is a MADYMO multi-body model. This model has been based on side-impact cadaver tests, performed by Wayne State University (Irwin,1993). In this model the shoulder girdle is modelled with a ball and socket joint, the humerus is directly connected to the thorax, the scapula and the clavicle are fixed. At first glance, tRis model looks three dimensional, but in fact it is

10

just one dimensional (only for side impact). This is caused by the way the contact between the body and the impact wall has been defined. To the lateral side of the body five one-dimensional mini-models have been defined. Those of the shoulder, thorax and abdomen consist of two bodies, connected with each other and with the human body model with three Kelvin elements. The mini-models of the pelvis and the left upper leg consist only of one body and one Kelvin element. The impact wall impacts the most distal body of each mini-model. Wayne State University also has developed a FEM model of a complete human body in Side Impact ( H u n g , 1994b).

The model by Irwin (1994) has been validated using the same tests as the model by Wayne State University (previous paragraph). This CAL3D model is based on the 50th percentile male GEBOD model. In this model the shoulder has been modelled using separate bodies for the clavicle and scapula. There is no direct connection between the scapula and the thorax (scapulothoracic gliding plane). The response seems to be rather well fitting the corridors, but the number of validation sets is limited and Joint characteristics for the three shoulder Joints have been tuned to give good results and are not based on direct measurements.

This model seems to be the most detailed shoulder model available for impact studies.

Yang and kövsund have developed a pedestrian model, based on a 58" percentile male GEBOD model, for car-pedestrian collisions (Yang, 1997). In this model the knee has been replaced by a more detailed knee model. Also ranges of motion and stiffnesses of the joints have been implemented. Just as in the GEBOD model, the shoulder is represented by only one joint.

Another complete human body model has been developed at TNO by Van den Kroonenberg (1997). This model is based on a Hybrid III model. In this Hybrid III model, the neck model by De Jager (1996) and a detailed spine model have been implemented. In this model the shoulder is modelled with only a clavicle, connecting the humerus and T4. This model has a detailed spine model, in which only forward bending is possible. The lateral bending stiffness is very high. This model has only been validated for rear impact.

The last available model has bcen developed at TUE by Willems (6997) arid is based on the human body model by Van den Kroonenberg (1997) and the shoulder model by Van der Helm (1991). From the model by Van den Kroonenberg all joint stiffnesses have been removed and the detailed shoulder model by Van der Helm has been implemented. In this shoulder model, the shouldergirdle has been modelled with separate bodies for thorax, clavicle, scapula and humerus. Muscle models have been inserted between these bodies. Some muscles (linear geometry, one dimensional behaviour) are modelled with only one active element, while other muscles, with large attachment sites, are modelled using several active elements to be able to model the multi-dimensional behaviour of these muscles. In this combined model the clavicle is connected to a 'thorax' body, to which also the gliding plane is connected. This 'thorax' body is fixed to T1 (bracket joint).

11

1.4 Goal ofthis Study

The goal of this study is, to develop and validate a mathematical model of the human shouldergirdle (including the arm) for impact conditions, as a part of a complete human body model.

The model should be suitable to simulate side-impact loading and preferably also frontal and rearward or even vertical loading. The model should be able to predict the globai response of arm and shoulder girdle in impact conditions (movement of the humerus, scapula, clavicle and sternum). The model should potentially also be able to predict the transfer of forces from arm to thorax. Here it is particularly relevant to discriminate different load paths like through the clavicle and the sternum and from the scapula directly to the ribs.

Further, attention will have to be paid to the connection between the shouldergirdle and the rest of the body, formed by the thorax.

Finally the resulting forces, accelerations, velocities and displacements of the several bony structures could be translated to injury criteria and injury mechanisms, to be able to predict the injuries resulting from a particular impact condition. With the final model simulations can be performed, like side airbag interaction with the shoulder.

The model is meant for impact, so parameters like masses, moments of inertia, joint and other stiffnesses, and the functional geometry should be entered correctly into the model. Due to the possible large rotations in joints during impact, for each joint not only joint stiffnesses, but also stop angles (maximum bending angles), as well as the stop-stiffnesses, should be defined. Active time-dependent muscle behaviour is supposed to be of lesser interest for impact, because the impact times (400 ms) are rather small, so that limited active muscle forces can be exerted during this period. However, passive muscle and ligament stiffness is important, just as possible prestress in muscles, present before impact, e.g. if the victim sees there will happen an accident and strains his muscles.

1.5 Modelling Methods

Two techniques are available for three dimensional mathematical modelling, Finite Elements (FE) and Multi-Body (MB) Modelling. FE requires a well-known geometry and material properties of each body, while MB treats a body as rigid. Due to this, MB has much shorter calculation times than FE. An advantage of FE is that stress and strain inside a body can be calculated.

The deformation of the shoulder occurs mainly in the joints and less in the bones, because the joint stiffnesses (inside the free range of motion) is much lower than the bone bending stiffness. If the deformation of the bones is neglected, MB seems to be the most appropriate for modelling the shoulder. The goal of the shouldermodel is to determine shoulder kinematics, not stress and strain in bones, which can be modelled using M B . As not enough data are available about the geometry and material properties of the bones and the surrounding tissue, developing a FE shoulder model,

12

also modelling bone deformation, may be very hard. As a first approach, the shoulder will be modelled using multi-body techniques. FE might be used in a later stage.

The thorax (ribcage) behaves quite different than the shoulder. In impact, there is a relative large deformation of the ribs and the costal cartilage. This indicates that modelling the thorax using only multi-body techniques may be not realistic enough, so that Finite Elements Modelling may be required. For sake of simplicity and because the main subject of this study is the shoulder and not the thorax, the thorax will be modelled using multi-body techniques. The use of FE for thorax modelling might be necessary in a fìïtheï stage. T h s , the complete humm body ;;rodel with detailed shoulder will be a multi-body model.

1.6 Preview

After the introduction (Chapter 1) first the anatomy of the human shoulder complex and thorax will be explained in Chapter 2.

discussed. A detailed description will be given of the methods used, the model geometry and the parameters used to create the final model. This description will be split into separate discussions for the spine, the thorax and the shoulder.

The model, created in chapter 3, will be validated in Chapter 4. This validation is based on side-impact cadaver test data, both sled and pendulum tests.

The sensitivity of the model response to several model parameters will be shown in Chapter 5. In this chapter, a sensitivity analysis will be performed, in order to obtain information about the necessity of measurement of several parameters.

arm with a deploying side airbag. This simulation and comparison of the results with EuroSID simulation data will be discussed in Chapter 6.

Finally the conclusions and recommendations will be given in Chapters 7 and 8.

In Chapter 3 the building of the shoulder complex and torso model will be

The validated model will be used for computer simulations of the interaction of the

13

14

2 Anatomy of the human shoulder complex and thorax

The base for the shoulder complex is formed by the ribcage. This ribcage consists of twelve pairs of ribs. At the posterior side of the body, all these ribs are connected to the thoracic spine, one pair of r h for each thoracic vertebra. From the spine, the ribs go to the lateral side of the body and on to the anterior side, forming thethoracic skeleton.

At the anterior side of the body, the upper seven pairs of ribs (ribs 1 to 7), called the true ribs, are connected directly to the sternum (breastbone) by the costal cartilage. The next three pairs of ribs (ribs 8 to 10) are not connected directly to the sternum. At the anterior side the costal cartilage of these ribs is connected to the costal cartilage of rib 7. So these ribs are connected indirectly to the sternum and are called false ribs. The last two pairs of ribs (ribs 11 and 12) are not connected to the sternum at all. These are the floating ribs.

_. Cervical vertebra VI rm ; - - lion

d process

cavity

Figure 2.I:Frontal view of the human rib cage (SoboîtaJ977)

15

The first bone of the shoulder complex is the clavicle. This long thin bone is directed medial-lateral. At the medial side the clavicle is connected to the upper side of the sternum by the sternoclavicular (SC) joint. The lateral end of the clavicle is connected to the acromion, a protrusion of the scapula just above the humeral head, by the acromioclavicular (AC) joint. The system consisting of the scapula and the clavicle has two joints with each three degrees of freedom, resulting in six degrees of freedom.

Another connection between the ribcage and the scapula is made by the scapulothoracic gliding plane. The medial border of the scapula glides over the posterior side of the ribcage, thus constraining two degrees of freeciorri of the scapula.

Of the remaining four degrees of freedom of the scapula and the clavicle the rotation of the clavicle around its own axis is not relevant and will only occur when a torque is applied direct on the clavicle. The three major degrees of freedom of the scapula with respect to the thorax are: - Lifting the shoulders - Movement of the shoulders to the front back - Rotation of the scapula, sliding across the thorax, the angulus inferior (lowest point

of the scapula) moving medially.

At the lateral end of the scapula there is the glenoid cavity. This is a rather small socket, which forms the glenohumeral (GH) ball and socket joint together with the humeral head, the proximal end of the humerus (upper arm).

Acromioclavicular Joint Coracoid Process of Scapula Acromium Process -L-kZ--+---- 2 CI avi cl e of Scawia 4

O

Humerus t I

Spine Of ScapuLCoraco id Process of Scapula Acromium Process

Glenohumeral Joint Head of Humerus

of Humerus

4 .- Humerus

Figure 2.2 (MDA,1997) Figure 2.3 (MDA,l997)

The most important ligaments of the shoulder complex are the costoclavicular ligament, between rib 1 and the clavicle near the SC-joint, and the coracoclavicular ligament, between the coracoid process of the scapula and the clavicle near the AC- joint. The coracoid process is a protrusion pointing forward next to the glenoid. At the posterior side of the scapula, there is the spine of the scapula. This is a horizontal rim, with the acromion at the lateral end, to which muscles are attached.

16

3 The Shoulder Complex and Torso Model

The basis for the shoulder complex model, described in this thesis, is the MADYMO human body model with multi-segment spine by Van den Kroonenberg (1 997). A more detailed description of this model can be found in 0 1.3.2. The shouldermodel itself has been based on the shoulder model by Van der Helm ( I 991), implemented in A A A n V A A n h x r Wil lPmc í7007ì. I.U u/ I L . I V "J I , I 1 I V I L I . J , a,, , ,

An image of the new human body model with detailed shoulder model can be found in appendix B.

The first and most important step in creating the MADYMO human shoulder model, is implementing the shouldermodel by Van der Helm in the human body model by Van den Kroonenberg. When this step has been completed, several things have to be changed, to enhance the biofidelity of the model, both from mechanical as from aesthetical point of view. All steps performed to create the final model, are described in the following paragraphs, divided in the spine, the thorax (ribcage) and the shoulder.

The model by Van den Kroonenberg (P997), is assumed to be good enough for the bodyparts not discussed in this thesis (legs, abdomen, frontal bending / axial elongation spine, neck and head), so that no further attention will be paid to these bodyparts.

3.1 Spine

Because the human body model by Van den Kroonenberg (1997) only has been validated for rear impact, first some changes have to be made to make it suitable for side impact. The most important part of the original model which is not suitable for side impact, is the spine. Each intervertebral joint has all six degrees of freedom, but only for forward / rearward bending and axial compression / elongation realistic stiffness and stop-angle values have been defined. For the other degrees of freedom (frontal and lateral translation and axial and laterai rotation) very high stiffnesses have been defined.

For kitera! iq32ct redistic vahes wi!l have to be defined fm at least lateral Sending. The stop-angles for lateral flexion have been derived from the range of motion given by Kapanji (1974), just as has been done for frontal bending (v.d.Kroonenberg,1997). The range of motion is given as 120 degrees (0.35 radians) for the lumbar and 120 degrees (0.35 radians) for the thoracic part of the spine. The stop-angles for each intervertebral joint can be found by dividing the range of motion by the number of joints. This results in stop-angles for the lumbar joints of 10.395 = 10.07 radians and for the thoracic part of &0.35/12 = 10.03 radians.

approximation the same stiffness until the stop-angle has been taken for lateral flexion as for frontal flexion. These stiffnesses are 678 Nm/rad for the lumbar spine and 1356 Ndrad for the thoracic spine (v.d. Kroonenberg, 1997; Prasad, 1974). After the stop- angle the stiffness has been chosen 50 times as large as before the stop-angle.

No data have been found concerning the lateral bending stiffness. As a first

17

The spinemodel is very stiff for frontal and lateral translation and axial rotation. In the model by Van den Kroonenberg no damping for these directions has been specified. The combination of high stiffness and no damping results in the occurrence of high- frequent vibrations. In the model, described in this thesis, these vibrations are being suppressed, by defining critical damping. This means that the damping parameter b

k k (damping force = b * velocity), equals 2Jmk = 2- = - . The undamped

wo @o eigenfrequency fo (Hz) has been estimated from a Fourier analysis of the undamped vibratiins;

Finally the ellipsoids representing the vertebrae have been changed. In the original model by Van den Kroonenberg (1997) the ellipsoids of the vertebrae spanned the whole back of the human body model. The lateral longitude of these ellipsoids has been changed, to make the ellipsoids represent only the vertebrae.

3.2 Thorax

In the model by Van den Kroonenberg, the ribs are represented by one single body, as in the Hybrid HII model, with next to each other two piles of six flat ellipsoids to visualise the ribs. So thorax deformation is not possible in this model, only thorax translation and rotation are allowed. For side impact, it is preferred to be able to have some lateral thorax compression. This implicates that at least two bodies next to each other are required. To make the thorax model more widely usable, also frontal thorax compression is required. To achieve this, the ribs-body from the model by Van den Kroonenberg is separated in four bodies, with each a pile of six ellipsoids. In figures 3.1 and 3.2 the original and the adapted thorax model (ribcage, sternum and thoracic spine) are shown.

Figure 3.1:Original thorax model Figure 3.2:Adapted thorax model

Each side (left or right) of the thorax consists of two bodies. These two bodies are connected by a revolute joint with a vertical rotation axis. The sternum and the spine are connected to the ribs using several Point-Restraints and Cardan-Restraints. Point- Restraints are elastic elements, with specified force-translation characteristics between the two connected bodies (3-D springs). Cardan-Restraints are elastic elements, with specified torque-rotation characteristics between the two connected bodies (MADYMU, 1996). Each back-ribs-body is connected to the spine with two Point-

18

Restraints; one between the ribs and TI and one between the ribs and T10. These two vertebrae are chosen, because they are the upper and lower vertebrae, to which true or false (but not floating) ribs are connected. The rotational degrees of freedom are restrained by a Cardan-Restraint between each back-ribs-body and T6. For the Cardan- Restraint T6 is chosen, because it is in the middle of the thorax.

The sternum is connected to each front-ribs-body by two Point-Restraints (at the upper and lower ends of the sternum) and a Cardan-Restraint.

It has to be noted, that because several Point-Restraints are defined between the ribs and the spine and between the stermm md the ribs, thc effective rotational stiffnesses of the sternum and the ribs are not only determined by fne Cardan-Restraints, büt also by the Point-Restraints.

The thorax model, as described till here, the ribcage could too easily rotate around the spine. This could not be prevented by adjusting rotational stiffnesses between the ribs and the spine, because these also affect lateral thorax compression. To prevent the rotation of the complete ribcage around the spine, a Point- and a Cardan Restraint have been added to the model between the sternum and thoracic vertebra T6.

Because in the current thorax model it is hard to obtain realistic values for the force-displacement and torque-rotation characteristics of the Point- and Cardan- Restraints, and because the modelling of the thorax is not the main subject of the study, these parameters have been tuned to give good results in validation (Chapter 4).

In the original model by Van den Qoonenberg (P997), the ribs are located lower than the thoracic vertebrae. In the new thorax model, the ribs have been moved upwards 5 cm, in order to be at the same height as the thoracic vertebrae. After this, the shape of the ellipsoids has been changed, to be appropriate for the new position.

In the original model by Van den Kroonenberg (1997), the back is formed by the vertebrae, while the ribs are located in front of the vertebrae. Due to the changes of the vertebral ellipsoids described at the end of 85.1, the back of the model no longer is formed by the vertebral ellipsoids. To solve this, the rib-ellipsoids have been changed, as can be seen in figures 3.1 and 3.2. The result of this is, that the back is formed by the ribs, just as in humans, thus enhancing both the aesthetical and mechanical biofidelity .

The thorax model described here will need more improvement, but because the main subject of this research is the shoulder girdle and not the thorax, this will not be a part of this stììdy. This thoïclx model is vqposed to be good enough for use I S i! basis for the shouldergirdle. Better thorax modelling can be done in other studies, using flexible bodies or E M . In the thorax model in this study thorax parameters have been tuned. A better thorax model will have to be based on more realistic parameter values.

19

3.3 Shoulder

The shoulder model by Van der Helm (1991) consists of a thorax-body, with a clavicle-body connected to it. To this clavicle-body the scapula-body, and to the scapula-body the humerus-body is connected. The connections between these bodies are ‘ball-and socket’-joints, with each three rotational degrees of freedom, representing the human synovial joints. Besides these joints, there is a direct connection between the scapula and the thorax; the scapuiothoracic gliding plme. In the ï ~ ~ ~ d d by Van der Helm, this connection has been defined by two points of the scapula, ‘gliding’ across the surface of the thorax. Finally most shoulder muscles have been defined in the shoulder model by Van der Helm

To make this model useful for impact e.g. geometries (for contacts) and joint characteristics will have to be defined and the shoulder will have to be attached to the thorax of the human body model. This will be discussed in the following paragraphs.

3.3.1 Scapulothoracic Connection

In the shoulder model by Van der Helm (1991) there is supposed to be contact between the thoracic gliding plane and two points of each scapula; Trigonum Spinae (TS, called node 37 by Van der Helm) and Angulus Inferior (AI, called node 38 by Van der Helm). To maintain this contact, a force pulling the nodes 37 and 38 against the gliding plane is required. This force must act perpendicular to the surface of the gliding plane ellipsoid, to prevent the shoulder to move spontaneously without an external force. This can be achieved in MADYMO using a Point-Restraint. This is an element with which stiffnesses in three perpendicular directions can be defined by specifying the force-displacement characteristics (MADYMO, 1996). A detailed description of this Point-Restraint and the scapulothoracic contact can be found in appendix C.

After the ribs were split up into four separate bodies (53.2) to each ribs body a scapulothoracic gliding plane ellipsoid has been attached. Each scapula contacts the corresponding gliding plane. In this way the scapulothoracic gliding plane will follow thorax deformation.

The thoracic gliding plane ellipsoids are not shown in appendix B, because they do not represent a geometry as whole, but only the shape of the ribcage near the scapuiae. In the image of the shouldermodel by Willems and Van der Helm (appendix A) the gliding plane ellipsoid has been shown once.

3.3.2 Clavicular Connection, AC-, SC- and GH-joints

Besides the connection at the two scapulothoracic gliding planes, the shoulder has a connection to the thorax through the clavicle. In the original model by Van der Helm (1991) the clavicle is connected to the same body as the gliding plane, the ‘thorax’. In the new combined model the clavicle is connected to the sternum, thus enhancing the mechanical biofidelity. Changing this connection did not affect the positions and orientations of any body in the initial position (seated).

20

The shouldermodel by Van der Helm (1991) contains no joint stiffness and joint stop characteristics for the three synovial joints of the shoulder: SternoClavicular (SC), AcromioClavicular (AC) and GlenoHumeral (GH). To make the shoulder model usable for impact simulations, these joint characteristics will have to be specified. This is done by means of Cardan-Restraints in the SC-, AC- and GH-joints. There is no information on the characteristics of separate joints. The research and measurements done on shoulder joint characteristics, mostly see the shoulder as only one (Ma,1995 ; Yang, I997 ; Engin, 1989a) or two (RAMSIS) joints. Irwin (1 994) has made a model with t'mee joints in the shorildeï, but the pzrz~~eters have been tiined and ate not based on direct measurements. Tümer and Engin (Engin, i98Sb;Lümer,l ray) nave caicülated joint sinus cones for the three different shoulder joints, based on the motion of the humerus with respect to the thorax, using an optimisation method with a "minimum joint motion" criterion. So the properties of the individual joints will have to be calculated, based on properties of the whole shouldergirdle, and can not be based on direct measurements.

The ranges of motion for the SC-, AC- and GH-joints are based on W S I S ( M S I S ) . This gives ranges of motion for the SC-joint and a combination of the AC- and GH-joints. For the left shoulder, these ranges of motion are given for an initial position with the humerus and the clavicle straight to the left (there is no scapula in this model). In this position the x-axis points in the direction of the clavicleíhumerus to the left, the y-axis points forward and the z-axis points down. The actual position is given by three successive rotations (y,&a) around the three axes (z,y,x). The ranges of motion follow from the minimal and maximal rotation angles:

~ n n n \ 1

angle min.(") max.(") a (x), around bone -15 5

min.(") max.(") -85 85

p (Y), up/down y (z), forwardhear

Table 3.1: RAMSIS ranges of motion (RAMSIS)

-25 15 -80 80 -15 35 -80 135

In the shouidermodei, these ranges of motion cue modelled, using Cardan Restraints. The actual orientation is defined by means of three successive rotations (cp,e,v) around the three z e s (x,y,z). These axes are oriented so tbat in the initial position with the left arm stretched to the left, the x-axis points upwards, the y-axis points forward and the z-axis points to the left. This differs from RAMSIS, which means that the axes (z,y,x) and angles (y,P,a) from RAMSIS, correspond to the axes (-x,y,z) and angles (-cp,û,v) from the model. For the right shoulder the orientation of the axes has been chosen to yield the same stop angles as the left shoulder: x-axis down, y-axis backward and z- axis left when the right arm is stretched to the right. So for both shoulders, cp represents rearward / forward bending, 8 represents up / down bending and represents rotation around the bone axis.

The range of motion for the AC-joint is chosen to be k0.35rad (k20") for all three directions. The range of motion of the GH-joint follows from the AC-joint and the sum of AC&GH-joints. This leads to the following ranges of motion for the three joints:

21

MADrnO I sc I sc I AC I AC I GH 1 GH I

Table 3.2: Ranges of motion for the shoulder model

The rotational stiffness for rotations outside the range of motion is chosen to be 4000 Ndrad . For rotations inside the range of motion, but within 0.05 rad (2.9") of the stop-angle (minimum or maximum rotation, as given in Table 3.2), the rotational stiffness is chosen to be 1000 Nm/rad. For other rotations the torque is zero.

In simulations of Wayne State University lateral sled tests with the model described in this thesis, using the joint characteristics as described here, the distance between the two acromions (shouldertips) did not change very much, while test data indicated a decrease of about 8 cm (Irwin,1994). Apparently, the connection between the two acromions, mainly formed by the two clavicles and AC- and SC joints, is much too stiff.

acromioclavicular (AC) and glenohumeral (GH)) have been purely rotational joints, with no translational degrees of freedom, and the clavicle and sternum have been purely rigid bodies. To decrease the stiffness of the connection between the two acromions, these assumptions must be altered. This can be done by changing the AC- and SC-joints, giving them some translational stiffness, or by making the clavicle or sternum deformable. It has been chosen to give the AC- and SC-joints some translational stiffness, because this fits best in the used multi-body approach. The translational stiffnesses in the AC- and SC-joint models represent not only the human AC- and SC-joint translational stiffnesses, but also clavicle and sternum deformation 'lumped' into the two joints.

To estimate the stiffness of the clavicle along its own axis, the clavicle has been split up in three straight beams, modelling the curvature of the clavicle. Using an E-modulus of 1,2E10 N/m2, used by Eummden for sibs (EkmameZen,d995), this resulted in a stiffness of about 1E6 N/m. It is assumed that the stiffness of the human AC- and SC- joints is lower than the clavicle stiffness. The combined stiffness of these joints and the clavicle (much lower than 1E6 N/m) are modelled as stiffnesses in the AC- and SC- joints. Based on occurring forces and needed deformation in the sled tests, the stiffness of the AC- and SC-joints has been chosen to be 2,5E5 N/m each, resulting in a combined stiffness of 1,25E5 N/m, which is eight times lower than the calculated clavicle stiffness.

Kroonenberg (1 997), the sum of the masses of the scapula and clavicle (including surrounding soft tissue) has been chosen about half the mass of the clavicle of the model by Van den Kroonenberg, resulting in 1 ,O kg each.

Until now, the synovial joints of the shoulder (sternoclavicular (SC),

To maintain the same total body mass as the original model by Van den

22

3.3.3 Arm

In the shoulder model by Van der Helm (1991) tRe elbow has been modelled by a spherical Joint, with tree successive rotations defining tRe actual orientation. This has been replaced by a revolute Joint, with only one rotation angle, because the human elbow has only one main degree of freedom.

TRe lower arm has separate bodies for the radius and the ulna. To maintain the same total body mass as the original model by Van den Kroonenberg (1997), the mass of the radius and the ulna (including surrounding soft tissue) has been chosen each about half the mass of the lower-arm-mass of the model by Van den Kroonenberg, resulting in 0,87 kg each.

3.3.4 Ellipsoids

To visualise the shoulder model bodies and to be able to define contact interactions between them, ellipsoids have to be defined for the clavicles, scapulae, arms and humeral heads.

SC-joints. The upper arm has been represented using an ellipsoid for the upper arm skin, and a sphere for the humeral head, to which shoulder impacts take place. The lower arm exists of an ellipsoid connected to the ulna.

The geometry of the scapulae has been represented using two ellipsoids for each scapula, one for the scapula itself, and one for the scapular spine with the acromion. The acromial end of the scapular spine ellipsoid decreases the range of motion of the GH-joint, because there can be contact between the humerus and the acromion. This is necessary because the range of motion as derived from RAMSIS is not entirely correct. In the human GH-joint combined upward and backward rotation of the arm, starting from stretched to the side, is not possible, because this is prevented by the acromion and the coraco-acromial ligament (Figures 2.1-2.3). To achieve this behaviour in the shoulder model, the acromion has been modelled and contact has been defined between the humerus and the acromion. The scapular spine (which is formed by the same ellipsoid as the acromion) has only a visual function.

and scapular spine, the humerus and the humeral head is shown in figure 3.3, with the point of view left, above and in front of tRe shoulder.

The clavicle has been modelled witR an ellipsoid along the axis between the AC- and

A picture of the left-shouldermodel, showing the clavicle, the scapula with acromion

Figure 3.3:Shoulder model

23

3.3.5 Muscles

In the model by Van der Helm muscles have been defined. Some large (multi- dimensional) muscles have been modelled using several one-dimensional muscle models (v.d.HeZm,1991). In table 3.3 all modelled muscles are shown with for each the number of one-dimensional muscle models (#)? the origin and the insertion:

Muscle I # I origin TraDezius (scamla) 16 I C7-T10

Rhomboideus I 3 I C7-T5 Serratus Anterior 6 thorax Deltoideus (scapula) 6 clavicle Deltoideus (clavicle) 6 clavicle Coracobrachialis I 6 I scapula Infraspinatus 6 scapula Teres Minor 6 scapula Teres Maior 6 scauula SuurasDinatus I 6 I scapula

insertion S

scamla scauula S

humerus

radius ulna I

Table 3.3: Modelled muscles

In the original model by Van der Helm, there is no difference between the spine and the thorax, so muscles originating at the thorax as well as muscles originating at the spine, are connected to the thorax. In the new shoulder model all muscles, originating at the spine, are connected to the corresponding vertebra.

The muscles originating at the thorax are connected to T1. These muscles are not connected to the corresponding ribs or to the sternum, because the thorax model will have to be changed in a later stage.

elongation characteristics are explained in appendix D. The shape of the force- elongation curve is the default MADYMO shape (MADYMO, 1996). For each muscle, only a force-parameter, taken from Van der Helm (1991), and the untensioned length, chosen equal to the initial length, are defined.

For all muscles only passive behaviour has been modelled. The passive force-

24

3.4 Total Model

In Figure 3.4 the final shoulder and torso model with muscles is shown. This model is a part of a complete human body model, but for the sake of clarity, only the adapted part is shown. The complete model can be found in appendix B.

Figure 3.4:Shouldes and torso model

25

26

4 Validation of the Model

For some parameters of the shoulder- and torsomodel no measurement values were available. These parameters concern the characteristics of the Point- and Cardan- Restraints of the thorax model, the contact stiffnesses of the contact between the impactor and the shoulder/thorax and the translational stiffness of the AC- and SC- joints (see Chapter 3). To obtin values for the unknown parameters, the available lateral impact validation data had to be used to tune these parameters. It was not possible to fulfil all requirements at the same time, so the parameters have been chosen to give the best overall fit. The results given in the graphs in appendix E represent this overall fit. A better fit would have been possible with less validation test, but the result would be less valuable. The fit of the response between the corridors and with the cadaver data is not very good, but it seems to be reasonable for a model in this stage. Model improvement will be necessary to enhance the fit.

Because the lateral validation data have been used for tuning of parameters, these datasets can not be used for real validation anymore. So this paragraph merely describes the tuning simulations instead of modell validation. But some validation is possible, because the extent in which the response can be fitted inside the corridors by tuning a parameter, is a measure for the quality of the remaining parts of the model. Hence, although not entirely correct, the simulations in this paragraph will be called wadidations.

The shoulder and thorax model has been validated for side impact, using human model (dummies and mathematical models) response requirements from IS0 WG5 (ZSO, 1996). In this document response requirements are given, based on cadaver tests by the APR (Bendjellal,1984), Ewing et al. (1977), Wayne State University (Viano, 1989 ; Irwin, 1993 ; Irwin, 1994), the Highway Safety Research Institute (Eppinger, 1978) and the University of Heidelberg (Marcus, 1983).

The in this study used response requirements give corridors for the contact force time history for three impact tests. These three tests are all lateral pendulum impact tests with a 150 mm diameter impactor of 23 kg. One test is against the shoulder, with an impact speed of 4.5 m/s (Shouldertest i), based on cadaver data by APR (Bendjellal,1984), normalised with the method by Mertz (1984). The test setup is shown in figure 4.1.

27

Figure 4J:Test Setup for Shouldertest 1 Initial (left) and impacted (right) position

The two other tests are impacts against the thorax, using the same pendulum (23 kg, 150 mm diameter) with impact speeds 4.3 m / s (Thoraxtest i) and 6.7 m / s (Thoraxtest 2>? based on cadaver data by HSRI (E’pinger,1978), normalised with the method by Mertz (1984). For these two thoraxtests, the force-corridors have been shifted 700 N up, to account for muscle tone (ISO,1996).For thoraxtest 1 also a corridor for the lateral thoracic acceleration versus time history has been specified. The test setup is

Figure 4.2:Test Setup for Thoraxtests 1 and 2 Initial (left) and impacted (right) position

All these validation simulations have been performed without gravity, because this did not seem to have much influence. The graphs of these lateral pendulum impact

28

validations can be found on the first page of appendix E, showing the ISO-corridors and the model results.

Besides the ISO-response requirements, as described in the previous paragraph, also sled test data from Wayne State University (Huang, 1994a ; Huang 1994b ; Irwin, 1993 ,- Irwin, 1994), have been used.

mounted laterally in the direction of movement to a sled. To this sled also an impact wall, consisting of seveïd box beams, was mowìted, approximately 750 mx from the side of the cadaver. This test setup is shown in figure 4.3.

In these lateral sled tests a human cadaver was seated on a rigid teflon seat,

Figure 4.3:Test Setup for the WSU Lateral Sled Tests Initial (left) and impacted (right) position

The sled was given an initial speed of 6.7 or 9.1 d s , and decelerated rapidly (over 203 mm) to a speed of O m / s , causing the cadaver to slide across the seat, its left side impacting the impact wdi. At the time of impact, the sled was supposed io have stopped moving. During the impact, contact forces between the cadaver and the beams 2nd displzcements of several bony lamharks have been recorded.

Irwin (1 994) has drawn corridors for the T5-displacement and shoulder&thorax contact force time histories for both impact speeds. These corridors have been used here for validation. The graphs of this validation can be found on the second page of appendix E, showing the corridors by Irwin and the model results.

acromions and the acceleration of the left (impacted) acromion and for 6.7 m / s of the displacement of the lower sternum, have been used. The graphs on the third page of appendix E show the comparison of the model results and the test data.

As additional validation, time-histories for both speeds of the displacements of both

Finally the model has been validated for a frontal pendulum impact on the thorax. This validation is based on guidelines by the Motor Vehicle Safety Systems Testing Committee (MVSSTC,1985), in which contact force - sternum deflection corridors are given for impacts with a 23 kg pendulum with a diameter of 150 mm, at 4,3 and 6,7 m / s , the same impactor and speeds as used in the lateral pendulum impacts on the

29

thorax. The setup for this test is shown in figure 4.4. The graphs resulting from this validation are shown on the fourth page of appendix E. In these graphs the acceleration of the pendulum has been given, equal to the contact force / 23 kg.

Figure 4.4:Test Setup for the Frontal Thorax Pendulum Test Initial (left) and impacted (right) position

30

5 Sensitivity Analysis

Some of the model parameters have not been based on measurement data, but have been tuned to give optimal results in the valiâation. To enhance the model, it might be useful to measure some of these parameters. A sensitivity analysis has been performed to obtain information about the sensitivity of the model to several model parameters

The default model has been chosen to be the described model (chapter 3)> but without the muscles. The muscles will be re-added as a variation. The sensitivity analysis concerns roughly all parameters except the geometry, like stiffnesses and masses. Besides model parameters, also some initial conditions have been varied, to study their influence on the model response and thus the importance of an accurate initial positioning.

In order to reduce the number of simulations to be performed and to reduce the data, several parameters have been changed simultaneously. The parameter variations that have been used are shown in table 5.1:

-I- ancl thiic --I-" the ..&I., A*&- irnpeguqcp ~f (hetter) m-eascrem-efit of these pzram-eters.

parameter set parameters variation( s) Thorax mass mass and moments of inertia of ribs "2

Thorax stiffness Stiffness and damping cardan- and point-restraints and joint-stiffness of the thorax mouel

ribs ar,d the hunerd head

(-Uk) and PESH (shape) (=default)

stiffness: "4 ; *0,25 damping: "2 ; *0,5

Contact stiffness Contact stiffness (and hysteresis) of the "2

Muscles All muscles models, parameters PEXM PEXM=0.8, PESH=10 " 8 3

PEXM=0.2, PESH=l (explained in appendix D)

Arm position Initial conditions GH-joint arm besides torso arm to the front

Shoulder position Initial conditions SC- and AC-joints SC-joint: clavicle 16 degrees to the front

Table 5.1: Parameters for sensitivity analysis (scapulothoracic contact maintained)

The influence of these parameters has been examined using the same validation tests as for the original model (chapter 4). To reduce the amount of data, only the graphs for the pendulum impacts and the graphs of the contact force and T5 and Acromion

31

displacements for the 9,l m i s sled test have been taken into account. For all described parameter variations these graphs are shown in appendix F and discussed below.

The Thorax mass hardy affects shoulder impact response. Thorax impact contact forces slightly increase (decrease) with 8.5% (-3.3%) due to increased (decreased thorax mass ("2, "0.5). The shoulder&thorax contact force in the sled test does not change for the first peak, but does for the following peaks. This indicates that the first peak is caüsed by shoulder contact, and tRe second by thorax contzct. InXuence OE acromion displacements is small (3%), indicating that acromion displacement is mainly caused by shoulder impact. As thorax mass is relatively well known and has little influence, no measurements seem to be necessary

An increase (decrease) of the clavicle and scapula mass ("2, "0.5) causes an increase (decrease) of peak contact force for thorax (117%) and shoulder (12.7%) impacts. Also the peak duration increased (decreased). An increased shoulder mass causes increased (decreased) Left and Right Acromion displacement and increased (decreased) acromion approach of 20% (14%). No good data are available for the effective mass of the clavicle and scapula (mass of bone and some of the surrounding soft tissue). As these masses clearly influence the results, better measurements are necessary.

The arm mass ("2, *0.5) has negligible influence on thorax impacts, but Shoulder impact contact forces increase 10% (decrease 6%) due to increased (decreased) arm mass. Increased (decreased) arm mass causes increased (decreased) Left and Right Acromion displacement and increased (decreased) acromion approach of 10%. As the arm mass has less influence on responses and is better known than the shoulder mass, measurements are less urgent, but can improve the model.

The influence of the AC&SC translational stiffness ("0.25, "4) on the thorax impacts is very small (peak contact forces 12%). The influence on the shoulder impacts is larger, but still not very large (112%) (also for sled test shoulder contacts). This stiffness influences mainly the right acromion displacements (130%) and thus the decrease of distance between the acromions (170%). Tnis parameter has been tuned arid has müch influence, so measurement is useful.

Limiting the AC-ROM from 1 0.35 to I O. 15 rad for all three directions has very little effect on the responses. Locking this joint mainly influences shoulder contact forces (+ 17%) and right acromion displacement (-25%). As limiting the ROM does not have much influence, better measurements are not very urgent. The modelling of the AC joint is important, because locking changes the response.

The stiffness of the thorax point- and cardan restraints ("4, "0.25) hardly affects shoulder impacts. The thorax impact contact force increases (decreases) with 21.9% (- 20.0%). Just as for thorax mass, the shoulder&thorax contact force in the sled test does not change for the first peak, but does for the following peaks, indicating that the first peak is caused by shoulder contact, and the second by thorax contact. Also

32

influence on acromion displacements is small (2.6%), again indicating that acromion displacement is mainly caused by shoulder impact. The thorax stiffness has been tuned and has considerable influence on thorax impact, so measurements are useful.

The contact stiffnesses ("2, "0.5) for the ribs and the humeral head affect all tests. An increased (decreased) contact stiffness decreases (increases) the width of all first force and acceleration peaks by approximately 30%. Peak contact forces increase (decrease) about 25% for the shoulder and 20% for the thorttx. Al1 displacements i s the sled tests decreased (increased) wit'n about 0.8 (i .U) cm, due to increased (decreased) penetration in the contact. The contact stiffness has large influence on the responses and has been tuned, so measurements are necessary.

The default muscles have no influence in the pendulum impacts and in the first 25 ms of the sled test. After 25 ms, there is mainly influence on the shape of the acromion displacement curves, causing a second maximum (lower than the first). The changed muscles have no influence in the first 10 ms of all tests. After 10 ms they cause higher contact forces (+17%), lower T5 displacement (-19%) and higher right acromion displacements (+25%). The muscles only have important influence if the stiffness is increased (changed muscles). Measurements are necessary to determine the actual passive (and active) muscle behaviour.

The initial position of the humerus has negligible influence on pendulum impacts (no contact in thorax tests). The initial position of the humerus more affects the shape of the contact-force-time-history curve than its magnitude, caused by the changed impact conditions. In general the arm besides torso causes lower displacements (till -22%) and the arm to the front higher displacements (till 9%). The initial position of the clavicle has very little influence on the pendulum impacts. The lateral displacement of T5 stays longer on the maximum of about 10 cm for the 6,7 m/s sled test, and is limited to 11,4 cm (original peak 13,4 cm) for the 9,l m / s sled test. Left acromion displacements decrease 9% and right acromion displacements 5%, causing the acromion approach not to change much. To obtain results matching cadaver tests, the initial shoulder and arm positions should be correct. Chaging these (md other) initia! conditions may affect iajury pwmeters.

33

34

6 Interaction with Side-Airbags

The human shoulder- and torsomodel, as described in the previous chapters, can be used for several simulations. In this chapter a simulation will be described, of the human body model, seated in a car that is struck from the side by a barrier with a velocity of 13,9 m/s. This car has a side-airbag. This simulation is based on a standard MADYMO application, described in the MADYMO Applications Manual (A4z4DKwQ, 1006). The SCtECup fer this Si,mu!ti^n is shown helew.

Figure 6.l:Simulation setup for the side impact application

In the original simulation an EuroSID- 1 model, described in the MADYMO Database Manual (MADYMO,1996), was seated in the car. For the simulation with the human body model, the EuroSID model has been replaced with the human body model. The initial position of the human body model has been adjusted to match the initial position of the EuroSID model as good as possible. The initial position of both the EuroSID and the human body model are shown in figure 6.2.

Figure 6.2:Initial position EuroSID (left) and human body model (right)

In appendix G graphs are shown of the lateral rotation of the head and the lateral acceleration of the head, T1 and T E , for simulations with the human body model as well as the EuroSD model.

35

The movement of the human body model looks quite realistic. While the side-airbag is being inflated, the shoulder is pushed upwards and the arm is thrown away by the airbag, possibly causing injuries. The upward movement of the shoulder does not occur in the EuroSID model, because its shoulders can not move. Also the movement of the lower arm can not be compared for the EuroSID model, because it has no lower arms.

At the moment the only injury criterium for the shoulder and upper extremity, is a reference value for humerus bending moment. No injury criteria are known for the scapula or clavicle, so the occurrence of injuries in these parts can not yet be predicted.

The most striking difference between the human body and EuroSU3 responses is the motion of the head. The EuroSID head rotates about 90" to the side, then contacts the shoulder, causing a high head acceleration peak (2000 m/s2), and rotates back towards the initial position. The human head does not contact the shoulder, but rotates about 160".

The lateral accelerations of T1 and T12 of the EuroSID and human body models show much resemblance. The main difference in the vertebral acceleration signals of the EuroSID and the human body model is, that the EuroSID signals have a peak of about 300 m/s2, caused by the head-shoulder contact.

shoulder is that apparently the model is suitable for this kind of simulations. No very strange deformations occur during the simulation. So an improved version of this model may be useful for studies in shoulder-airbag interactions.

A very important result of the simulation with the human body model with detailed

36

7 Conclusions

In this study, a three dimensional mathematical model of the human shoulder complex has been developed as a part of a complete human body model. A problem in developing such a model is, that there is very little information available about the characteristics of the shoulder complex. Measurements of shoulder joint stiffness and free range of motion mostly concern the whole shoulder complex, i.e. the combi~ed properties of the sternoclavicular, acromioclavicular and glenohumeral joints and the muscles. Individual joint properties have been calculated, based on these measurements.

For validation of the shouldermodel only lateral cadaver sled and pendulum test data are available, as used in chapter 4. Frontal validation of the shoulder model is not really possible, because no frontal tests are known, in which there is an impact against the shoulder and/or sensors have been mounted to the shoulder.

The shoulder model has been validated using pendulum and sled tests (chapter 4, appendix E). Some model parameters have been tuned to give the best possible fit on all tests. Unfortunately, it was not possible to stay inside all given corridors at the same time. Enhancing the fit in one test, caused a worse fit in another test. A more detailed model (e.g. Finite Elements thorax model, flexible bodies for shoulder bones) may be able to give responses better fitting inside the corridors, because more possible deformations (e.g. clavicle deformation) can be modelled.

In comparison with the shoulder model by Irwin (1994), which is the best available shoulder model known to the author, the model described in this thesis has some advantages.

the ribcage, formed by the scapulothoracic gliding plane, has been modelled, which is not the case in the model by Irwin. Also, the clavicle is connected to the sternum, connected to the ribs, connected to spine, which has separate bodies for each vertebra. In the model by Irwin the clavicle is connected to one single 'thorax' body. Furthermore, the model by Irwin has been validated only using the WSU sled corridors for contact forces and T5-displacements (chapter 4, appendix E). Due to this limited validation, Irwin is able to get a better fit, that also could have been achieved by the model from this thesis.

In the model, described in this thesis, the direct connection between the scapula and

The validated shoulder model has been used for a sensitivity analysis (chapter 5). The parameters with the most influence on the model responses are the clavicle and scapula mass, translational stiffness of the AC- and SC-joints, stiffnesses of the thorax model and contact stiffnesses. Muscles only have major influence if the passive muscle stiffness is increased. Parameters with less influence are the mass of the thorax and the arm and the range of motion of the AC-joint.

37

A side impact simulation, in which a barrier laterally strikes a car with a side-airbag, has been performed with a EuroSID model and with the human body model with detailed shoulder (chapter 6) as drivers. The main difference between these models was the rotation of the head, which was much lower in the EuroSID model, because of contact with the shoulder. The human body model showed a realistic movement of the shoulder; it was pushed upwards by the airbag.

Overall, the model seems to show the global kinematics of the shoulder in impact (direcî or by an airbag) ï&er well. Two load paths from the scapula tc the thorzx c m be distinguished in the model. This has been validatad with vertebral acceïeration data. For the model to have a really predictive function for the occurrence of injuries it will have to be validated more thoroughly, for which more measurements are required.

38

8 Recommendations

Based on the problems, encountered in the development of the human body shoulder and torso model, some recommendations can be done for model improvement:

The thorax response to a frontal impact does not fit the corridors at all. Also the connection between the ribcage and the spine does not seem to be very good. In the current model this cannot be improved, because the thorax is not able to follow spine deformation (frontal and lateral bending). So the thoraxmodel, consisting of five rigid bodies (the spine not included), will need improvement. It seems not possible to make a much better model using only rigid bodies. Further improvement of the thorax model may require flexible bodies andor FEM techniques.

The current shoulder model consists of one rigid body for each bone (clavicle, scapula, humerus). Using this assumption, the model cannot get much more detailed. Further improvement of the shouldermodel wil% need the use of flexible bodies for the clavicle and scapula. The shape of the scapulothoracic gliding plane can be improved using facet surfaces or a FEM-thorax. Also the shape of the scapula can be modelled using facet surfaces. The force that pulls the scapula against the thorax should be exerted by muscles.

Measurements are required of several model parameters, to be able to use realistic values in the model. Parameters that need more measurement are the effective clavicle and scapula mass (bone and some of the surrounding soft tissue), translational stiffnesses of the AC- and SC-joints (andor clavicle deformation), thorax stiffness, and contact stiffnesses. For other test conditions than used in this thesis, also individual joint range of motion and stiffness data may be required.

To be able to validate a shouldermodel for all possible impacts, more cadaver test data are required, like e.g. frontal impact on the shoulder, frontal sled tests, and tests in which the shoulder is pushed upwards. In all these tests displacements or accelerations of bony landmarks of the shoulder (e.g. acromion) should be measured.

If the default MADYMO parameters are used for the muscle models, the muscles do not have much influence. If these parameters are adjusted using recent data, causing the muscle stiffness to increase a lot, the muscles start to have some influence. Research will be required, to determine which parameter values are realistic for muscle behaviour (active as well as passive) in impact conditions. If the muscle stiffness seems to be much higher than the default in MADYMO, the muscles should be included in the shoulder model.

39

40

References

Bass, 1997 The Interaction of Air Bags with Upper Extremities C.R. Bass, S.M. Duma, J.R. Crandall, R. Morris, P. Martin, W.D. Pilkey, S. Hurwitz, N. Baewpong, R. Eppinger, E. Sun S A X 973324

Bendjellal, 11984 APR Biomechanical Data F. Bendjellal, G. Walfisch, A. Fayon, C. Tarriere Nanterre, France, 1984

CBS, 1997 Centraal Bureau voor de Statistiek http://statline.cbs.nl

Engin, 1989a A Statistical Investigation of the In Vivo Biomechanical Properties of the Human Shoulder Complex A.E. Engin, S.M. Chen Mathl. Comput. Modelling, Vol. 12, No. 12, 1989, pp. 1569-1582

Engin, 1989b Three-Dimensional Kinematic Modelling of the Human Shoulder Complex - Part I: Physical Model and Determination of Joint Sinus Cones A.E. Engin, S.T. Tümer Journal of Biomechanical Engineering, vol. 11 1, may 1989, pp. 107-1 12

Eppinger, 1978 Development of a Promising Universal Thoracic Trauma Prediction Methodology R.H. Eppinger, K. Augustyn, D.H. Robbins STAPP 780891

Eummelen, 1995 Adult to Child Scaling of the Human Thorax in Frontal Impact Conditions, Applied to a Hybrid Finite ElementMulti-Body Thorax Model in MADYMO P. Eummelen WFW-report nr. 95.004, Eindhoven University of Technology, 1995

Ewing, 1977 Measurement of Head, T 1, and Pelvic Response to -Gx Impact Acceleration C.L. Ewing, D.J. Thomas, P.L. Majewski, R. Black, L. Lustik STAPP 861893

41

Frampton, 1997 An overview of Upper Extremity Injuries to Car Occupants in UK Vehicle Crashes R.J. Frampton, A.P. Morris, P. Thomas, G.G. Bodiwala IRCOBI Conference - Hannover, September 1997, p.37-5 1

Happee, 1992 The Control of Shoulder Muscles During Goal Directed Movements R. Happee Ph.D. Thesis Delft University of Technology, 1992

v.d.Helrn, 1991 The Shoulder Mechanism - A Dynamic Approach F.C.T. van der HePm Ph.D. Thesis Delft University of Technology, I991

Huang, 19943 A MADYMO Model of Near-Side Human Occupants in Side Impacts Y. Huang, A.I. King, J.M. Cavanaugh Journal of Biomechanical Engineering, vol. 116, may 1994, p.228-235

Huang, 1994b Finite Element Modelling of Gross Motion of Human Cadavers in Side Impact Y. Huang, A.B. King, J.M. Cavanaugh STAPP 942207

Huelke, 1997 Upper-Extremity Injuries From Steering Wheel Airbag Deployments D.F. Huelke, R. Gilbert, L.W. Schneider STAPP 970493

Irwin, 1993 Displacement Responses of the Shoulder and Thorax in Lateral Sled Impacts A.L. Irwin, T.J. Walilko, J.M. Cavanaugh, Y. Zhu, A.I. King STAPT 933 124

Irwin, 1994 Analysis and CAL3D Model of the Shoulder and Thorax Response of Seven Cadavers Subjected to Lateral Impacts A.L. Irwin Ph.D. Thesis Wayne State University, 1994

ISO, 1996 Committee Correspondence ISO/TC22/SC 12íWG5 - Anthropomorphic Test Devices, Document N455 - Revision 2 Road Vehicles - Anthropomorphic Side Impact Dummy - Lateral Impact Response Requirements to Assess the Biofidelity of the Dummy May 1996

42

de Jager, 1996 Mathematical Head-Neck Models for Acceleration Impacts M.K.J. de Jager Thesis Eindhoven University of Technology, 1996

Kallieris, 1997 Response and Vulnerability of the Upper Arm Through Side Air Bag Deployment D. Kallieris, A. Rizetti, R. Mattern, S. Jost, P. Priemer, M. Unger S A E 973323

Kapanji, 1974 The Physiology of Joints. Volume 3: The Trunk and the Vertebral Column H.A. Kapanji Churchill Livingstone, 1974

v.d. Kroonenberg, 1997 A Human Model for Low Severity Rear-Impacts A.J. van den Kroonenberg, J.M.G. Thunnissen, J.S.H.M. Wismans International IRCOBI Conference on the Biomechanics of Impact, 1997, pp. 117-132

Ma, I995 Development of Human Articulating Joint Model Parameters for Crash Dynamics Simulations D. Ma, L.A. Obergefell, A.L. Rizer STAPP 952727

MADYMO, 1996 MADYMO manuals version 5.2 TNO Road-Vehicles Research Institute, 1996

Marcus, 1983 Human Response to and Injury from Lateral Impact J.H. Marcus, R.M. Morgan, R. Eppinger STAPP 83 1634

MDA, 1997 The Medical Disability Advisor http://www .asianconnect.com. sg/health/mdasample/anat9.html December, 1997

Mertz, 1984 A Procedure for Normalising Impact Response Data H.J. Mertz SAE 840884

MVSSTC, 1985 Human Mechanical Response Characteristics Report of the Motor Vehicle Safety Systems Testing Committee SAE J1460, March 1985

43

Prasad, 1974 An Experimentally Validated Dynamic Model of the Spine P. Prasad, A.I. King Journal of Applied Mechanics, 1974, pp. 546-550

R M S I S , 1998 Body Builder Users Guide version 1. 1 Tecmath human modelling, 1998 Kaiserslautern, Germany

Sobotta, 1977 Atlas of the Human Anatomy, Vol. 1: Regions, Bones, Ligaments, Joints and Muscles. 9th English Edition J. Sobotta, F.H.J. Figge Urban & Schwarzenberg, Baltimore-Munich, 1977

Turner, 1989 Three-Dimensional Kinematic Modelling of the Human Shoulder Complex - Part 11: Mathematical Modelling and Solution Via Optimisation S.T. Tumer, A.E. Engin Journal of Biomechanical Engineering, vol. 11 1, may 1989, pp. 113-121

Viano, 1989 Biomechanical Response and Injuries in Blunt Lateral Impact D.C. Viano STAPP 892432

Willems, 1997 Modelling Compensatory Postural Muscle Use in Subjects with a Thoracic Spinal Cord

M.M.M. Willems WFW report 97.068, Eindhoven University of Technology, 1997

Injury

Wismaíís, 1994 Injury Biomechanics, course notes 2nd printing J.S.H.M. Wismans, E.G. Janssen, M. Beusenberg, W.P. Koppens, H.A. Lupker Eindhoven University of Technology ! TNO Crash-Safety Research Centre Delft, 1994

Yang, 1997 Development and Validation of a Human Body Mathematical Model for Simulation of Car-Pedestrian Collisions J.K. Yang, P. Lövsund IRCOBI Conference - Hannover, September 1997, p. 133-149

44

Appendices

Appendix A : Images of available models . . . . . . . . . . . 46 Appendix B : Images of the model . . . . . . . . . . . . . . 49 Appendix C : Scapulothoracic contact . . . . . . . . . . . . 51 AppendixD: PassiveMuscle Characteristics . . . . . . . . . 52 Appendix E : Model Validation . . . . . . . . . . . . . . . 53 Appendix F : Parameter Analysis . . . . . . . . . . . . . . 57 Appendix G : Side Impact Application . . . . . . . . . . . . 77

45

Appendix A: Rasdel Wayne State University

46

Appendix A: Model van den Kroonenberg

47

Appendix A: Model Willems (Van des Helm)

48

Appendax Bc Images of the model

49

Appendix B: Images of the model (continued)

50

Appendix C: Scapnlothosaclc contact

The scapulothoracic contact consists of a gliding-plane ellipsoid, attached to the ribcage, small spheres for AI (Angulus Inferior, inferior point of the scapula, node 38) and TS (Trigonum Spinae, medial point of the scapula, node 37), contact interactions between these two and a Point Restraint, pulling the scapula towards the gliding plane, perpendicular to the ellipsoid surface.

The scagulothoracic gliding plane is formed by an ellipsoid with half-axis lengths in x- y- and z-direction lx, ly and lzp thus following the condition:

It can be found that the force on a point (xp,yp,zp) on the surface of this ellipsoid is perpendicular to the ellipsoid surface if the force vector (Fx,Fy,Fz) follows the following condition:

F F F

In this condition the parameter F can be chosen freely and is allowed to be dependent of xp, y, and zp. To create this force using a Point-Restraint, the force in one particular direction must be a function of only the displacement in that direction. This means, that the parameter F cannot be dependent of xp, yp and zp. So F is a constant factor, which can be chosen freely.

A Point-Restraint with this characteristic is defined between the gliding plane and each node. So the Point-Restraint generates a force acting from the gliding plane on the node, perpendicular to the ellipsoid surface. In the initial position this force must be compensated exactly by the contact force of the contact between the gliding plane and the node. This contact force equals the product of the contact stiffness k and the penetration. Because the centre of each node is located on the surface of the ellipsoid, the penetration equals the radius of the node-sphere. So the radius of this sphere has to be chosen equal to the resulting Point- Restraint force divided by the contact stiffness.

The force parameter F is chosen to be 5 Nm. The resulting Point-Restraint force in the initial position of the nodes is then 27.63 N for node 37 and 34.95 N for node 38. This is close to the contact force in rest of about 25 N, given by Happee (Happee,92) .

The contact stiffness is chosen to be 3.0E+04 N/m. To make the contact force equal to the Point-Restraint force, the radii of the node-spheres (equal to penetration) must be 0.000921 m for node 37 and 0.001 165 m for node 38.

It would also have been possible to define the same radius for all nodes, and different Point- Restraints or contact stiffnesses.

51

Appendix D : Passive Muscle Characteristics

All information in this appendix is taken from the MADYMQ Theory Manual (MADYMO, 1996).

The passive muscle force F, at a relative length 1, (= actual length / lref) is determined by:

f, is the standard MADYMQ passive force-lengtR relation:

k (expek (I -I)]-I) forl , >i P 0 for Ir I 1

f = I 2 "

with constants: k =lil(exp(PE 1-1} 1 ska

The parameter PE,h determines the shape of the force lengt curve The parameter PE, determines the relative elongation ((l-lr&lref = 1,-1) inducing a passive muscle force equal to Fmm, so that fp equals 1.

The user has to specify values for: lref the reference (untensioned) length F,, the maximum active force the muscle can exert PE,h shape parameter PE, elongation parameter

A graph of the function fp(lr) for several values of PE,h and PE, is shown below.

I .o

0.8

Q

a,

O

a, >

u-

L 0.6 cc

.- U

- 0.4 L

0.2

0.0 1

I I I I I I

/ PEsh=l P Ex m = 0.2

1 ---- I I I

1.2 1.4 1.6 relative length Ir

52

Lateral Pendulum Tests IS0

vi w

Thoraxtest 1 ................. Corridor Acceleration Impactor (m/sA2)

Corridor Acceleralion Iwipactor ( m i A Z ) Acceleralion Imc>aclor ( d s ? ? ) Model Rmporiye

.................

1x0

I60

140

120

I O0

no

60

4c

2c

C

Time (ms)

Thoraxtes t 2 300

250

200

150

LOO

50

O

................. Corridor Acceleration Impactor (misAZ) Corridor Acceleration Impactor (dr'2) Acceleration Imc>acior (ds"2) Model Rerponse

.................

10 20 30 40 50

Time (ms)

Thoraxtest 1 300

200

I O0

o

-100

-2oc

................. Corridor Acceleraliwi T6 (nW2) Corridor Acceleration T6 (tn/s^2)

Acceleralion T6 ( d s A Z ) Model Rmportse

.................

R

10 20 30 40 50

Time (ms)

S houldertest 1 ................. Corridor hripaclor-Stioiilder Confacl Force (N)

Corridor Impaclor-Shoulder Coritact Force (N) ...................

4000

3000

2000

1000

O

Time (ms)

Model Validation

vi P

Lateral Sled Tests Wayne State University Corridor ï%orar&Shoulder Contact Force (N) 6.7 rds Trsl Corridor ïkorar&Shoulder Conlact Force (N) 6.7 d s Test Conlacl Force (Nj Model Response 6,7mfs Test

______........... ______...........

*le3

I I I O 10 20 30 40 so 60

Time (ms)

______........... .................

Corridor 7liorar&Shorilder Conlacl ForLe (Nj 9,l ~ d s Tesl Corrrdor ï%orax&Shoulder Contact Force (Nj 9,l r d s Ter1 Conlacl Force (Nj Model Response 9,l nvs Tesl

*le3

12.

10

8

6

4

2

O I I I I I 1 10 20 40 so 60 30

Time (ms)

o. I2

o. I

0.08

0.06

0.04

0.02

O

-0.02

................. Corridor La18ral Dirplncerrterrl T5 (rnj 6,7rrds Tesl

................. Corridor LaleralDirplacemenl T5 (nij 6,7111,s Te.11 Laleid Dirplacemenl T5 (mj Model Response 6,7 ttds Tesl

I I I 1- 10 20 30 40 50 60

Time (ms)

0.14

0.12

o. 1

0.08

0.06

0.04

0.02

O

-0.02

Corridor LaleralDisplncement T5 (m) 9.1 I I I ~ Tal Conidor LaleralDispiacemenl T5 (In) 9.1 1d5- Tesl UiternlDisplacemetirT5 (m) Model Response 9.J ndsTesl

.................

......._.........

10 20 30 i 0

Time (rns)

@ Model Validation -

Lateral Sled Tests Wayne State University

vi vi

LateralDisplacement Left Acromiori (mJ Model Response 4 7 d s Test LateralDisplacemenl Left Acromion (mJ Model Respome 9.1 rds Test TesfDafa LaleralD~placemenlLeft Acromion (in) 4 7 r d s Test Test Data LateralDh~placementLeft A ~ r ~ m i o n (m) 9,l d s Test

........._...__._

*le-3

so

40.

30.

20.

10

U

I /

I I I 10 20 30 40 50 60

Time (ms)

IS00

1000

500

O

-500

-1000

-1500

-2000

-25-00

-3000

Lateral Acceleration Left Acromion (dr"') Model Respome 6,7rW/s Test lateral Acceleration Left Acronuon (m/s^Z) Model Response 9.1 d s Test Test Data Lateral Acceleration Left Acromion ( d s A Z ) 6,7m/s Test Test Data Lateral Acceleration Left Acromion ( d s " 2 ) 9,l d s Tesl

.................

I 7 5 4 10 12

Time (ms)

0.14

0.12

o. 1

0.08

0.06

0.04

0.02

O

LateralDLsplacemnnt Right Acromion (mi Model Response 6.7 i d s Test lateral DLsplacerfinnl Right Acromion (in) Model Resporm 9.1 d s Test Test Data Lateral Dmplaceinerit Right Acromion (m) 6.7 i ds Test Test Data lateral Disylacenurit Riplit Acromion (m) 9.1 mis Test

..........._.____ - . - . - . -

/ /

I I I I -- 10 20 30 40 50 60

Time (ms)

o. I

0.08

0.06

0.04

0.02

O

lateral Displacemerit Lower Sternurn (mJ Model Respome 6,7 rrds Test TestDala LaferaiDisplacemenl Lower Sternum (mJ 6,7trds Test I . . . . . . .. . . . . . . . . .

10 20 30 40 so 60

Time (ms)

Frontal Thorax Pendulum Test 1

n 180- cv c

160- E v

C 140 O .I ci L 120

0 100 a

a a I

E 80 3 < 60 C Q)

I

e 40

20

O

New Model with gravity, 4,27 d s Test New Model without Gravity, 4,27 d s Test Original Model, 4,27 i d s Test Corridor 4,27 d s Test Corridor 4,27 d s Test

-. -. - . - .................

.................

I I I I I I I I I I I I I I I ---+

-0.01 0 0.01 0.02 0.03 0.04 0.05 0.06 0.07

Sternum Deflection (m)

- New Model with Gravity, 6,71 m/s Test New Model without Gravity, 6,71 m/s Test Original Model, 6, :7l d s Test Corridor 6,71 d s Test Corridor 6,71 d s Test

-. -. -. -

.................

Sternum Deflection (m)

I

e

Lateral Pendu ................. Corridor Accelerafion Impacfor (ds"2J

Corridor Acceleration Irnpncfor (rds"2) Accelerafion Impacfor (rrJs"2) Original Model Acceleration Impacfor (rds"2) Model wifh Higher Mass Accelerafion Impacfor (rrJs"2) Model wifh Lower Mass

.................

Thoraxtest 1 200.

1 so

I00

50

0

/ - \

Time (ms)

................. Corridor Accelerafion Impador (dY2) Corridor Acceleration Impactor (m/sA2) Acceleration lrnpaclor (ds"2) Origuzal Model Accelerafion Impacfor (rrJs"2) Model wifh Higher Mars Acceleration Imacfor f ds "2J Model wifh Lower Mass

.................

Thoraxtest 2 300

250

200

1 SC

1 O(

5(

I

I

/ - .

Time (ms)

lum Tests IS0 ................. Corridor Accelerafioii T6 (rds"2)

Corridor Accelerafion T6 (m/s"Z) Accelerafion T6 (rdsA2J Original Model Acceleration T6 (dY'2) Model wifh Higher Mass Accelerafion T6 (rrdsA2J Model wifh Lower Mass

.................

Thoraxtest 1 200

I00

Time (msb

..... Corridor Impacfor-Shoulder Confacf Force (NJ I 3 houldertest 1

.................. Corridor Impactor-Shoulder Con facl Force (N) Confacf Force (NJ Original Model Confacf Force (N) Model wifh Higher Mass Confacf Force (NJ Model wifh Lower Mass

4000.

3000

2000

1000

O 10 20 30 40 60

Time (ms)

Lateral Sled Tests Wayne State University

9,l m/s *le3

................. Corridor Thoru&ShouMer Contact Force (N) Corridor Thoru&Shoulder Conlacl Force ( N ) Contact Force (N) Original Model Contact Force (N) Model with Higher Mass Contact Force (N) Model wifh Lower Mass

.................

- . - . -. -

I I I 10 20 30

Time (ms)

I 40 50

9,B m/s so

40

30

20

10

O

................. Cadaver Dal0 uiteral Displacement Left Acromion (m) lateral Dirplacermrit Left Acromon (m) Origrrul Model lateral Dirplacenent Left Acromion (In) Model wilh Higher Mass lateral Dnplacernenl Lef1 Acromion (m) Model with Lower Mass - . -. - . -

I 40

I 50

Time (nis)

9,l m / s O. 14

0.12

o. 1

0.08

0.06

0.04

0.02

O

-0.02

................. Corridor Laleral Displacetmnf TS (m)

LateralDisplaceinent TS (in) Origiiuil Mob1 lateralDnplaceirurf T5 (mi Model with Higher Mass lateral Dirplaceirunf TS (in) Model wifh Lower M m s

................. Corridor laleralDnplflcemenf T5 (In)

-. -. - . -

-.-_ - - . -_

... 1-._

-.. 3..

........... ........ ..... .._. --._

......

10 20 30 40

Time (ms)

9,l rn/s O. 14

0.12

o. 1

0.08

0.06

0.04

0.02

O

................. Cadaver Data LaleralDisplacer,~iit Righf Acromion (in)

IaleralDirplaceinenf Righl Acronuon (m) Origrml Model IateralDirplncement Right Acromion (In) Model with Higher Mars uileral Accebrnlion Righf Acromion (in) Model wilh Lower Mass

............ .......... ....... .............

.._ --._.

I I I I 1 IO 20 30 40 50

Time (mc)

Model with Higher / Lower Thorax Mass -

Lateral Pendulum Tests IS0 ................. Corridor Acceleration Impactor (dsAZJ

Corridor Accelerafioii Impactor (rn/r'ZJ hl Acceleration Acceleration Irrpaclor Impactor (m/sn2J (rris"2J Model Origuzal with Model Higher Masses

.................

horaxtest 1 Acceleratiori Impactor (rrisVJ Model with Lower Marses - . -. _ . -

180

i 60

140

I20

100

8C

6C

4c

2(

(

Time (ms)

................. Corridor Acceleration Impador (nus"2J Corrrdor Acceleration Impactor (I~V'ZJ Accelerrilion Impactor (rrisAZJ Origural Modef Accelerarion Impactor (dsA2J Model with Higher Masses Acceleration Irmactor Irrisi? J Model with Lower Masses

.................

Thoraxtest 2

Time (ms)

- ................. Corridor Acceleration T6 (m/sA2J

Corridor Acceleration T6 (rds"2J

Acceleration T6 (ridr'2) Model with Higher Masses Acceleratiori T6 (rrdsAZJ Model with Lower Masses

................. - Acceleraliori T6 (rrdsAZJ Original Model Thoraxtest 1

300

200

100

C

-IOC

-20(

-30( I I I I l 10 20 30 40 so

Time (ms)

j houldertest 1 SO00

4000

3000

2000

1000

O

........................... ii" ,.' I I .... 1,' I I ' b

................. Corrrdor Impactor-Shoulder Contact Force (NJ Corridor hpacror-Shoulder Contact Force (NJ Coriracf Force (NJ Original Model Contacf Force (NJ Model wilh Higher Mflssm Coritact Force (NJ Model with Lower Mmses

................. -- _ . - . .- . - --

... .-..

I I 30 I 40 '58 60 10 20

Time (ms)

&@ Model with Hiqher / Lower Clavicle & Scapula Mass -

Lateral Sled Tests Wayne State University

9,l m / s

................. Corridor ï%orar&Shoulder Contact Farce (NJ Corridor Thorar&Shouider Conlacl Force ( N ) ................. Contact Force (Nj Origuial Model Confacl Force (Nj Model with Higher Massa Conlact Force (Ni Model wilh Lower Marses

9,l d s *le3

................. Cadaver Dala LateralDisplaccmenf Left Acromion (mj LzleralDlrplacement Lqï Acromion (m) Original Model LaferalDisplacemertlL~l A c r ~ ~ ~ n (mj Model wrth Higher Mawe8 LateralDirplacement L@ Acromion (mj Model with Lower Marses _ . - . - . -

14

I2

10

8

6

4

2

(

Time (ms)

0.06.

0.05

0.04

0.03

0.02

0.01

O

Time (nis)

9,l m / s 0.14

0.12

o. 1

0.08

0.06

0.04

0.02

c

-0.0;

................. Corridor Lateral Dirplacemenl T5 (m) Co'rridor Lateral Displacemenl TS (4 Laleraf Dirplacemenf TS (rri) Original Model LaleralDisplacemeril TS (in) Model wifh Higher Masses LaferalDirplacemenl TS (mj Model with Lower M a r s ~

.................

-. _ . -. -

io Z O 30

Time (ms)

40

Cadaver Data LaferalDi.PplaceWIenf Right Acromion (In)

LaferalDisplacemeni Right Acromion (m) Original Model Lateral Dirplacement Righl Acromion (mj Model with Iligher Marses Lateral Acceleratmn Right ACIOIIUO~ (in) Model wilh Lower Marsm _ . -. _ . -

o r I O

I -I 10 20 30 40 so

Time (ms)

Model with Hiqher / Lower Clavicle & Scapula Mass

Lateral Pendulum Tests IS0

Thoraxtest 1 ................. Corridor Acceleration Impaclor ( d s ^ Z J ................. Corridor Acceleration Impaclor ( d s " 2 )

Acceleration Impaclor (ds"2) Original Model Acceleralion Impuclor (m/sYJ Model with Higher Mass Acceleration Impactor (rrJsA2) Model with Lower Mass

180

I60

140

120

100

80

60

40

20

C I I I *I I 10 20 30 40 so

Time (ms)

................. Corridor Acceleralron Impaclor (ds"2)

Acceleralion Zny>actor (rrJssY) Origutal Model Acceleration Impnclor (dsAZJ Model with Higher Muss

-. - . - . - Acceleralion Iwwacior ím/sA2) Model wilh Lower Mass Thoraxtest 2

300

250

200

1%

1 O(

S(

[

Time (ms)

Thoraxtest 1 ................. Corridor Accelmzlron T6 (rw3"') ................. Corridor Acceleralion T6 (~ds^Z)

Acceleration T6 ( I~u 's~Z) Origuul Model Acceleration T6 (nUs^Z) Model wilh Higher Mass Acceleration T6 (nvs"2) Model wilh Lower Mam

300

200

1 O0

0

I

I 'h

-100-

-200 I I I I I 10 20 30 40 50 O

Time (ms)

S houldertest 1 .................. Corridor Impaclor-Shoulder Contact Force (N)

Contad Force (N) Origmal Model Contucf Force (N) Model wilh Higher Mass Coniacl Force (NJ Model with Lower Mars

.................. Corridor Impactor-Shoulder Coritacf Force (N)

so00

4000

300C

ZOO(

loo(

(

I'

Time (ms)

@ Model with Hiqher 1 Lower Arm Mass

Lateral Sled Tests Wayne State University ................. Corridor Thorar&Shoulder Contact Force (Ni

Corridor Thorax&Shoulder Contact Force (N) Contact Force (N) Origmal Model Conlact Force (N) Model wilh Higher Mass Conluct Force (N) Model with Lower Mars

.................

9,l d s *le3

I I I I + O 10 20 30 40 so

Time (ms)

9,l d s 0.06.

0.05

O 04

0.03

0.02

0.01

O

................. Cadaver Data LateralDisplacemnl Lefi Acroniion (m) LateralDiPplacemnl L e t Acromion (m) Original Model LateralDirplacemnt Left Acromion (m) Model with Higher Mars LaferalDirplucemnt L@ Acromion (m) Model with Lower Mars - . - . - . -

I I I ,I 10 20 30 40 so

Time (ms)

- ................. iCorridor Lateral Dirylacement T5 (tri) ................. iCorndor LateralDirylacenzenl T5 (m)

.laferalDirplacemenf T5 (ni) Original Model lateralDirp/acer!xnt T5 (in) Model with Higher Mars 9,l d s

0.14

0.12

o. I

0.08

0.06

0.04

0.02

c

-0.0:

...........

I I - 10 20 30 40

Time (ms)

9,l d s ................. Cadaver D a l LateralDisplacemerrt Righl Acromion (rri)

lateralDisplace>mril Right Acromion (in) O r i g i w l Model LateralDuplaceiwnt Right Acroinron (m) Model wilh Higher Mars laleral Acceleration Rinht Acromion fmj Model with Lower Mars

0.14

0.12

o. I

0.08

0.06

0.04

0.02

O

............ ---. ......... ...... ...................

I I I -- 10 20 30 40 so

Time (ms)

Lateral Pendulu m Tests IS0

Thoraxtest 1 ................. Corridor Acceleration Impactor (m/sAZJ

Corridor Acceleration Impactor ( d W J Acceleration Impactor (m/sA2J Original Model Acceleration Impactor (ds"2J Model with AC&SC StYyer Accelerntron Impactor (dWJ Model with AC&SC Weaker

.................

I80

160

140

120

100

80

60

40

20

O

Time (rns)

Thoraxtest 2 300

250

200

I50

1 oc

SC

(

~~

................. Corridor Accelerafion Impactor (W"^ZJ Corrrdor Acceleration Zmpactor (m/s"ZJ Acceleration Impactor (m/sA2J Origulal Model Accelcratron Impactor (m/sAZJ Model wllh ACBSC Stiffer

.................

- . - . _. - Acceleration Impactor (m/sA2J Model with ACBSC Weaker

Time (ms)

- ................. Corridor Acceleration T6 (dS"2J

Corridor Acceleration T6 (irhA2J Acceleration T6 (nds"2J O r l g m l Model Acceleration T6 (ndsAZJ Model with AC&SC Stirer Acceleration T6 (m/s%'J Model with ACBSC Weaker

.................

Thoraxtest 1 300

200

100

O

-100

-200

-3OC io I

20 30 40 so

Time (ms)

3 houldertest 1 ................. Corridor Impactor-Shoulder Contact Force (NJ

Corridor Impactor-Shoulder Conlact Force (NJ -- Confad Force (NJ Original Model

Contact Force (NJ Model with A C G C Stiffer Contact Force IN) Model wilh A C G C Weaker

...................

- . - . __ . - 5000

4000

3000

200c

loo(

(

f \ ' I M!

Time (ms)

~~

Lateral Sled Tests Wayne State Univers ................. Corrrdor Thorax&Shou!der Contact Force (Nj

Corridor Thorax&Shoulder Contact Force (Nj ................. Contact Force (Ni Origmal Model Contact Force ( N ) Model with A C M C Stirer Contacl Force (NI Model with A C M C Weaker - . -. -. -

9,l m / s *le3

-. I

O 10 20 30 40 50 I I I

Time (ms)

9,l m / s ................. Cadaver Dntn LateralDisplacement Left Acromion (m)

lateralDisylacerient L-ft Acromion (m) Original Mudel LateralDLrplacemnf Left Acromion (mi Model with ACBSC Stiffer breral Displacement Left Acromion (m) Model with AC&SC Weuker -. - . -. -

so

40

30

20

10

C I IO 20 30 40 i 0

Time (ms)

9,l m / s 0.14

0.12

o. 1

0.08

0.06

0.04

0.02

O

-0.02

................. Covidor Lateral Displacemerit T5 (m) Cowidor LateralD~rplacement T5 (»I)

lateral Displacement T5 (rij Original Model Latt,ralDirplacement T5 [id Model with ACLSC St#er LaiirnlDisplacernent T.5 (m) Model with AC&SC Weaker

.................

-. -. -. -

I I I I 10 20 30 40

Time (ms)

Cadaver Data lateral Displacement Right Acromion [In) LaleralDispla<emnt Right Acromion (m) Original Model LaferalDoplacemerif Right Acromrorr (In) Model with AC&SCStk7er

- . - . -. - laleral Accelerntm Right Acromion (in) Model with AC&SC Weaker

0.14

0.12

o. I

0.08

0.06

0.04

0.02

C

............ ... .......... ... ........... ._..e.

10 20 30 40 so

Time (ms)

50 Model with AC&SC Pointrestraints Stiffer / Weaker

Lateral Pendulum Tests IS0 ................. Corridor Acceleration Impactor (m/sAZJ ................. Corridor Acceleration Impactor (rds"2) Thoraxtest 1 r-1 Acceleration Acceleralion Ifnpaclor Impactor (m/sAZJ (m/s"Z) Model Origvral with Model AC Locked

- . -. -. - Acceleration Impactor (ni/s^ZJ Model with AC ROM 0.15 rad

1 8 0 1

Time (ms)

................. Corridor Acceleration Impactor (m/s^ZJ

................. Corridor Acceleration Impactor (m/sAZJ

Acceleration Irnpactor (m/s*ZJ Origuial Model Acceleration lnyactor (m/sAZJ Model with AC Locked Acceleration Impactor ( d s A Z J Model with AC ROM 0.15 rad

Thoraxtest 2 i - . - . - . -

300

2so

200

1 SO

I O0

so

O

Time (ms)

Thoraxtest 1 300

200

100

O

-100

-200

- ................. Corrfdor Acceleration T6 (fds"2J ................. Corridor Acceleratiofi T6 (frús*ZJ

Acceleration T6 (rdsAZ) Original Model Acceleration T6 (dsAZJ Model with AC Locked Acceleration T6 (m/sA2J Model with AC ROM 015 rad I I - . - . -. -

7

1 - 1 I I I 10 20 30 40 50

Time (ms)

5 houldertest 1 -- ................. Corridor Impactor-Shoulder Contact Force (NJ

Corridor Impactor-Shoulder Conlacl Force (N) ................. -- Contact Force (NJ Original Model

Contact Force (NJ Model wrth AC Locked - . - . - . - Contact Force (NJ Model with AC ROM 0.1 5 rad

/ \ 40001 / \

.... .._.

't.-- A-+- - -i I I I

0 10 20 30 40 50 60

Time (rns)

Model with AC Ranqe of motion 0.15 rad / AC Locked -

Appendix F: Parameter Analysis (continued)

L 2

4

m-

I I I

h

E .- ì! u

I-

CA

2 4

m-

m -

n

E .- E" W

I-

* , l i I '

I I . I I

n

E ì! u

i=

n

E .- ì! W

I-

Lateral Pendulum Tests IS0

Thoraxtest 1 250.

200-

1%-

100.

so

O

................. Corridor Acceleration Impactor (ds"2) Corridor Acceleration Impactor (dr '2) Acceleration Impactor (idsA2) Original Model Acceleration Impactor ( d s Y ) Model with Stiffer Thorax Acceleration Impactor (d2"2) Model with Weaker Thorax

.................

- . _ . _. -

10 20 30 40 so

,h ..Li ....................... /, /.\.,.\ '. '. .' ...... ~

...................................

-.. \ ... .... .: . _ _ _ _ - . ---. .

\ \ \

I I I 10 20 30 40 so

Time (ms)

Thoraxtest 2 ................. Corridor Acceleration Impactor (ds"2)

Corridor Acceleration Impactor (ds"2) Acceleration Iny>actor ( d s " 2 ) Original Model Acceleration Impactor (dW) Model with Stifer Thorax Acceleratron Imuactor ídsA2J Model wilh Weaker Thorax

.................

- . - . -. -

300

250

200

1 so

1 O0

so

O

I . ' ' I

I I I 1 10 20 30 40 so

Time (ms)

Thoraxtest 1 300

200

I O0

O

-100

-200

-300

................. Corridor Acceleration T6 ( 1 ~ 2 ~ 2 )

Corridor Acceleration T6 (in/S*2) Acceleratron T6 ( r r W 2 ) Originul Model Accskroiion T6 ( d @ Z ) Model wrlh Sl~#j%r ïh iora

Acceleration T6 ( I I ~ Y ? ~ ) Model with Weaker Thorax

.................

li\ \

I I I -- 10 20 30 40 sa

Time (ms)

7 houldertest 1 ................. Corridor Impactor-Shoulder Contact Force (NI

Corridor Impacior-Shoulder Contact Force (N) Conlact Force (N) Original Model Contact Force (N) Model with StLfler îñorax Contact Force fN) Model wrlh Weaker ïïiorax

................. -_ - . - . __ . -

4000

3000

2000

1000

O

Time (mc)

Model with Stiffer /Weaker Thorax -

Lateral Sled Tests Wayne State University

9,1 m / s

1 m / s

................. Cadaver Dala LaleralDisplacement Lefl Acromion (mj LaferalDisplacemenl Leji Acromion (mJ Original Model LaterulDirplacewirnt Lef1 Acromon (m) Model with Stifler Thorax LateralDbplacemenf Left Acromori (m) Model with Weaker T h o r a

* le3

121

................. Corridor Thiora&Shoulder Contad Force (Nj Corridor Thora&Shoulder Contact Force (NJ Conlact Force (Nj Origaal Model Contact Force (Nj Model with Sl@w Thorax Conlact Force (N) Model wilh Weaker Thorax

.................

O i o 30 40

Time (ms)

I I I 1 -

10 20 30 40 50

Time (ms)

9,l m/s ................. Corridor LaleralDisplacemnr T5 (»ij

Corridor LaleralDwplacemenf T5 (mJ LaleralDirplacemeiif T5 (in) Origirurl Model LateralDisplacemenf T5 (mJ Model with Stffer T7wrax LateralDisplacemett T5 (tnJ Model with Weaker Thora

.................

0.12

............................ I

\ -i

F- - - - - - -.. ... //

....

....... 0.06

,/ .,' 0.04 .....

O 10 20 $0 40

Time (ms)

9,1 mis ................. Cadaver Data LaieralDisplacement Right Acromion (mJ

Lateral Dbplacemenl Right Acromon (mj Origrrul Model LaleralDirplacemnt Right Acromion (mj Model with Stifler T?mm Lateral Acceleration Right Acromion (IE) Model with Weaker Thorax

O i o 20 30 40 i 0

Time (ms)

Model with Stiffer /Weaker Thorax

Lateral Pendulum Tests IS0

Thoraxtest 2 ................. Corridor Acceleration Impactor (m/ f lZJ

Corridor Acceleration Impactor (rn/sA2J Accelerafion Inpactor (m/sA2J Original Model Acceleration Impaclor Im/sA2) Model wifh Conlacis Sliffer

.................

................. Corridor Acceleration Impactor (m/sAZJ Corridor Acceleration Impactor (m/s^Z) Acceleration Ifr!gactor (dW) Origlnal Model Accelerafion Impactor (rdsAZJ Model wrth Contacts Stiffer

................. ................. Corridor Acceleration T6 (rrdsAZJ

Corridor Acceleration T6 (rn/sA2J Acceleration T6 (rids"2) Origuml Model Acceleration T6 (rrr/SAZJ Model wifh Contacfs Sfiffw

.................

Thoraxtest 1 Thoraxtest 1 Acceleration Impactor (rrJS"2J Model with Contacts Weaker I I - . -. - . - Acceleratron T6 (rids"2J Model wifh Corifacfs Weaker I I - . -. -. -

c I \

200.

I 50

1 O0

50

O

300

200

100

0

-100

-200

-300

h

I 10

I I 20 30

I I 40 50

Time (ms) Time (ms)

5 houldertest 1 ................. Corridor linpaclor-Shudder Contact Force (NJ

Corridor Impactor-Shoidder Contact Force (NJ Confact Force (NJ Origmal Model Coiilact Force (NJ Model with CoritacLr Stirer Contact Force (NI Model wilh Coiilacls Weaker

................. ~-

- . - . -. . - Acceleraiwii I r ~ a c l o r (mWZJ Model wilh Conlacls Weaker - . - . - . -

10 20 30 40 50

350

300

250

200

150

1 O0

50

0

5000

4000

3000

2000

1000

O I I 10 20 30 40 50 60

Time (ms) Time (rns)

with Contact Ribs & Model St iff nesses Stiffer Humeral Head Weaker

4 O

Lateral Sled Tests Wayne State University

9,l d s *le3

14

12

10

8

6

4

2

O

\ !I

................. Corridor Thorax&Shoulder Contact Force (N) Corridor Thorax&Shoulder Contact Force IN) Corilact Force (N) Origmal Model Confact Force (N) Model wilh Contflcls Sl fe r Conlact Force (N) Model wdh CorilacLp Weaker

.................

- . -. - . -

IO 20 30 40 i 0

Time (ms)

9,l d s ................. Cadaver Data Lateral Displacement Lefr Acromion (m)

Lateral DlsplacertIent L@t Acromion (m) Original Model LaferalDlsplncementL~t Acromion (m) Model wilh Conlacls Sliffer LateralDlsulacemenl Left Acromon fm) Model with Conlaclr Weaker - . -. -. -

0.06

0.05

0.04

0.03

0.02

0.01

O I I I I 7 IO 20 30 40 SO

Time (ms)

- ................. Corridor LateralDirplacernenl T5 (In)

Corridor Lateral Dmplacetnenl T5 (mi Infe,dDirplacernenf T5 (mj Original Model Late~mlDirplacernerrt T5 (mi Model with Coritacts Slifeer Late,'-nlD~~placernenl T5 (mj Model with Contacts Weaker

.................

O. 16

0.14

0.12

o. 1

0.08

0.06

0.04

0.02

O

-0.02 I I -- 10 20 30 40

Time (ms)

9,l d s 0.14

0.12

O. I

0.08

0.06

0.04

0.02

O

~

................. Cadaver Data ILiteralDisplacemnl Right Acromion (m) Lateral Dcrplacemeril Right Acromion (mj Original Model LzteralDcrplncement Righl Acrormon (m) Model with ConlactsJtifer Inferal Acceleraliori Righl Acroniion (rnj Model with Corilacls Weaker -. -. -. -

I I I I 1 10 20 30 40 so

Time (rns)

Lateral Pendulum Tests IS0

..............................

................. Corridor Acceleration Iwipactor (m/sA2) I I Corridor Acceleration Impactor (m/sA2J Acceleralion Impactor (nds"2) Origural Model Acceleration Impactor (nds"2) Model with Defadt Murcles Accelerafion Impactor (nds"2) Model with Changed Muscles I I -. - . - . -

180

160

140

120

1 O0

80

60

40

20

O

Time (ms)

Thoraxtest 2 ................. Corridor Acceleration Itripnctor (m/sA2)

Corridor Accelerntion Impnctor (misAZ) Acceleration Impactor (rdsAZJ Origuiul Model Acceleration Impactor (idsl?) Model with Default Muscles

.................

-. _. - . - 300

250

200

I50

100

50

O

Time (ms)

Thoraxtest 1 200

100

O

-100

-2oc

................. Corridor Asceleralion T6 (rdsA2) Corridor Asceleralion T6 (tn/sA2) Acceleration T6 (rn/SYJ Origmal Model iiccelerution T6 ( r rJSY) Model with Default Muscles

.................

Acceleration T6 (ids"2) Model with ChangedMuscles I I -. - . _ . -

n

Time (ms)

S houldertest 1 .................. Corridor Iinpactor-Shoulder Coirtacl Force (N)

Corridor Impactor-Shoulder Conlact Force (N) Contact Force (N) Origmal Model Contact Force (N) Model with Default Muscles Contacl Force (NI Model with Chamed Miiscler

.................. ___-

_ . - . -. - 4000

3000

2000

1000

O

Time (nis)

Model with Default / Chanqed Muscles

Lateral Sled Tests Wayne State University

9,l m/s * le3

14

12

10

8

6

4

2

O

Corridor Iliorax&ShonI&r Contact Force (Nj Corridor Thorar&Shoulder Conlacl Force (N) Contact Force (Nj Origmal Model Contact Force (Ni Model with Default Muscler

.................

.................

10 20 30 40 so

Time (ms)

Cadaver Data LateraiDhrpíacemenl Lefr Acromion (mj Lateral Dlsplacemeiii L-fl Acromion (mi Orig ia l Model Lateral Dfsplacement Left Acromion (mj Model with Default Murclm LaferalDhrplacementLeft Acromion (mj Model wilh Changed Murdus

*le-3

so

40

30

20

10

O 10 20 30 40

Time (ms)

I so

9,l mis ~

Corrrdor Laleral Dlsplacemnt T5 (rnj

Corridor Laleraf Dcrplacemenf T5 (mj Laferal DLrplacemenf T5 (mij Original Model LateralDlsplacement T5 (in) Model wilh Defadl Murclus LaleialDLsplacemenf T5 (in) Model wilh Changed Murclus

.................

.................

0.14

0.12

o. 1

0.08

0.06

0.04

0.02

O

-0.02 i o i0 i 0

Time (ms)

I 40

~~

Cadaver Data Lateral Displacemenl Right Acroiniori (in)

LaleralDlsplacermiir Right Acronuon (mj Origiiuil Model Lateral Dlsplaceimnt Righl Acrotniori (in) Model with Defaull Murcles Lateral Acceleration Rqhl Acromion (inJ Model with Chmiged Murcles -. -. -. -

O. 14

0.12

o. 1

0.08

0.06

0.04

0.02

O

\ ...

10 20 30 40 so

Time (rns)

Model with Default / Changed Muscles

4 w

Lateral Pendulum Tests IS0 ................. Corridor A ccelerafion lmpncior (mis"2) I I

Corridor Accelerntion lmpnctor (mis"2j Acceleration Impactor ( d s A 2 j Original Model II_L_1 Acceleration impnctor (misA2j Model with Arm Besides Torso Thoraxtest 1 Acceleration lmpactor (Iris'z) Model wrth Arm to the Front I I - . -. -. -

180.

160.

140-

I 20

100.

80

60.

40.

20

O

Time (ms)

Thoraxtest 2 300

250

200

I so

1 O0

so

O

~~~~ ~~

.................

................ Corridor Acceleration Impactor (mLv*Z) Acceleration Impactor ( m i s Y j Origutnl Model Acceleration Iinpnctor ( d s " 2 ) Model with Arm Besides Torso Accelaratwn lwactor (ds"2j Model wifh Arm l o the Fronr I I -. - . - . -

10 20 30 40 50

Time (ms)

Thoraxtest 1 300

200

100

O

-100

-200

-300

~~ ~

.................

................. Corridor Acceleralion T6 (mísA2)

Acceleration T6 (tdsA2j Origuul Model Acceleration T6 (misAZ) Model with Arm Besides Torso Acceleration T6 (nds"21 Model wilh Arm l o lhe Front I I - . - . -. -

10 20 30 40 i 0

Time (ms)

S houldertest 1 ................. Corridor Impactor-Shoulder Contact Force (Nj

Corridor Impaclor-Shoulder Contact Force (Nj Coritact Force (Nj Original Model Contact Force (NI Model with Arm Boido lor,^

.................

4000

3000

2000

1000

0

I - . - . - ' - Conlncl Force (Nj Model with Arm to the Front I

A!

Time (ms)

Model with Arm Besides Torso / to the Front

4 P

Laterall Sled Tests Wayne State U

9,l m/s *le3

................. Corridor ïhorar&Jhoulder Contact Force (Nj Corridor Iliorax&Shoulder Contact Force (Nj Contact Force (Nj Origuial Model Contact Force (N) Model with Arm Berdas Torso

.................

(NJ Model with Arm to the Front 111 fl.\- ;\ ..............................

I I I O 10 20 30 40 so

Time (ms)

9,l m i s 0.06

0.05

0.04

0.03

0.02

0.01

O

................. Cadaver Data Lateral LateralDisplacemenrLefl Acramon (mj Orlgrnal Model LaleralDLsplacemenr Lef1 Acromon (mj Model with Arm Berider Torso LateralD~placemenrLefrAcromion (mj Model with Arm to the From - . _. - . -

I 10 20 30 40

Time (ms)

1 so

niversity

9,l mis ................. Corridor Lateral Displacemmr TS (m)

Corri'ior LaleralDisplacemetit TS (mJ Lateral Displacement TS (mJ Original Mo& Latend Displacemen1 TS (mJ Model with A ~ I H Bmides Torso

.................

TS (mj Model with Arm to the Front

0.16

0.14

0.12

o. I

0.08

0.06

0.04

0.02

O

-0.02 I I I I 10 20 30 40

Time (ms)

9,l m/s O. 14

0.12

o. I

0.08

0.06

0.04

0.02

O

................. Cadaver Data Lateral Displacement Right Acromion (m) Lateral DLsplacernenl Right Acromiori (raj Original Model Iateral DLsplacemnt Right Acromion (in) Model with Arm Berides Torso LaleralAccebratio,n Right Acrommri (mJ Model with Arm to the Front -. -. -. -

10 20 I I 30 40

Time (ms)

1 so

Model with Arm Besides Torso /to the Front

-

Lateral Pendulum Tests IS0 ................. Corridor A ccelerntm Impactor (ds"2)

Corridor Acceleration Impactor ( d s " 2 ) Accelernrion Impactor ( r r J s 9 ) Origmal Model Accelernliori Iwncror ( d s A 2 ) Model with Clavicle 10 the Front

........ Thoraxtest 1 1 L.... 1 180

160.

140.

120.

100.

80.

60

40.

20.

0.

Time (ms)

................. Corridor Accelerntiori Impactor (ds "2 ) Acceleration Impnctor (rn/sA2J Origmnl Model Accelerafion Impactor i ds "2 J Model wilh Clavicle to the Front Thoraxtest 2

300

250

200

I so

1 O0

so

o

Time (ms)

Thoraxtest 1 200

1 O0

o

~ ~~

.................

................. Corridor Acceleration T6 (rds^Z) Accelerafion T6(dsA2) Origuial Model Accelerntiori T6 ( d s ^ Z J Model wilh Clavicle lo the Front

-100-

-200 I I I -- o 10 20 30 40 so

Time (ms)

S houldertest 1 ................. Corridor Impador-Shoulder Contact Force (N)

Corridor Irnpaclor-Shoulder Contact Force (N) ................. - Contact Force (N) Origmnl Model

Contact Force IN) Model with Clavicle to the Frorrt

400C

300C

2000

1000

O

w

Time (ms)

Model with Clavicle 16 degrees to the front -

Lateral Sled Tests Wayne State University

9,l d s *Ie3

................. Corridor Thorax&Shoulder Contact Force (N) Corridor Thorax&Shoulder Contact Force ( N ) Corifact Force (NI Original Modei Contact Force (N) Model with Clavicle to the Front

.................

I I I I 1 10 20 30 40 so

Time (nis)

................. Cadaver Data Lateral Disdacement Leït Acromion /mi lateral Dhrpbcemenl Lefr Acrormon (m) Orrgrnal Model l a f e r a l D i s p l a c e m e ~ L ~ 1 Acrormon (mi Model with Clavicle lo the Front 9,l rn/s '

*le-3

so

40

30

20.

10.

0. 1 I v 10 20 30 40 50

Time (ms)

- - ................. Corridnr Lareral Dwplacemerit T5 (m)

Corridor IateralDirplacemenf T5 (m) IaferalDirplacenwnt T5 (m) O n g i r d Model Iaterai Dirplacemenl T5 (In) Model with Clavicle to the Fro181

.................

/ - - - - - - - - - - ................................. r . - - - 0.12 o.li

-0.02 I I I -- O IO 20 30 40

Time (ms)

Cadaver Data Lalera~lDi,iplacemerit Right Acromion (m) laferalDisplacemen1 Right Acromion (in) Origrnal Model Ialeralüisplacemenl Right Acromion (in) Model with Clavicle /o /he Front

0.14

0.12

o. 1

0.08

0.06

0.04

0.02

O

-0.02.

............ .... ...... ........ ........... ._..I

-..

I I I I I 10 20 30 40 so

Time (ms)

Model with Clavicle 16 deqrees to the Front

Side Impact with Barrier, including Airbag

160

140

120

1 O0

80

60

40

20

O

-20

I lateral Head Acceleration (ds"2J EuruSW Lateral Head Acceleratron (dsA2'2) Human

e& 2000 nils"2

I I I I I1 20 40 60 80 1 O0 120

Time (rns)

Lateral TI Acceleration ( d s A 2 ) EwoSID LateralTl Acceleration fdsA21 Human

500,

400-

300.

200.

100-

0-

-100-

-200-

-300- I I I I I 7

20 40 60 80 100 I20

Time (ms)

0.5

O

-0.5

I

-1.5

2

-2.5

-3

Time (ms)

lateral TI2 Acceleration (misAZ) EuruSiû lateral TI2 Acceleration (rrúsA2J H m m

600

500

400

300

200

1 O0

O

1 O0

-200.

lateral Head Rotation (rad) EwoSID Lateral Head Rotation (rad) IJmmrr E- -_ __ - - .

--y \ '.J \

/ \

/ \ / \

/ \

.. _.. ' \

/' /

/ /

I I -- so 1 O0 150 200

\/ I I I I --

20 40 60 80 1 O0 I20

Time (ms)

78