Vol. 18, No. 4, 2016 DOI: 10.5277/ABB-00633-2016-03
Biomechanical analysis of injuries of rally driver with head supporting device
KAMIL JOSZKO1*, WOJCIECH WOLAŃSKI1, MICHAŁ BURKACKI1, SŁAWOMIR SUCHOŃ1, KAROL ZIELONKA2, ANDRZEJ MUSZYŃSKI2, MAREK GZIK1
1 Silesian University of Technology, Faculty of Biomedical Engineering.
2 Automotive Industry Institute.
Purpose: The aim of the study was to develop and verify a model of rally driver with a safety system HANS (head supporting de- vice), which will enable biomechanical analysis of injuries in rally accident. Methods: Simulations were carried out in Madymo® soft- ware, the results of which were verified based on sled test performed in the Automotive Industry Institute (PIMOT) in Warsaw. The model being verified allowed us to perform a multivariate simulation of rally accident in terms of assessing effectiveness of protection and usefulness of HANS system. Results: Acceleration waveforms of the head and chest were obtained from numerical experiment and also forces and moments occurring in the upper cervical spine. The results obtained allowed driver injuries to be analyzed based on injury criteria of the head and neck: HIC15, NTE, NTF, NCE and NCF. Conclusions: The analysis enabled assessment of the driver safety while using 4 and 5 point harness with HANS system. In further studies the model developed was used to identify factors affecting the safety of a rally driver.
Key words: safety systems, rally driver, injury assessment of driver, crash test, HANS device
1. Introduction
Safety systems used in rally cars are significantly different with respect to safety systems in passenger car. This is a consequence of specific conditions in motorsports where risk of high speed accident is sub- stantial. The rally car compared to commonly used car is deeply modified. One example of modifications is safety roll cage which increase stiffness of construc- tion for crew protection. Other examples are bucket racing seats and multipoint harness for reducing body displacement during accident. During accident a re- duced crumple zone and tighter fixation of torso may have a negative effect on head and neck spine. During head’s forward movement its inertia is increased by helmet mass (around 1.4 kg, depending on the model and size of helmet). These facts show that rally car safety systems need to be complemented with extra
protection for head and neck. An answer to this de- mand is HANS (Head and Neck Support device) de- signed by Dr. Robert Hubbard for NASCAR drivers in 1980s [12]. Since 2003, after some modifications, HANS device is obligatory for rally drivers in Poland, since 2009 every driver participating in international events is obligated by FIA (Fédération Internation- ale de l’Automobile) to use HANS [13]. From the beginning, some drivers keep questioning the need of using head restrain systems as they are uncomfort- able to use [3].
There are many papers concerning investigation of car accidents with experimental and numerical stud- ies. However, majority of publications in the field of head and neck safety describe whiplash effect and head restraint, e.g., [1], [7], [10]. Only few papers investigate motorsports and sport car drivers. Numeri- cal experiments were applicate in construction analy- sis, i.e., F1 bolide [11] and rally car [17], [18]. In [5],
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* Corresponding author: Kamil Joszko, Silesian University of Technology, ul. Roosevelta 40, 41-800 Zabrze, Poland. Tel: 698002547, e-mail: Kamil.Joszko@polsl.pl
Received: April 29th, 2016
Accepted for publication: June 24th, 2016
[6], [14], the authors confirm, based on experimental methods, helpfulness of HANS device. The present paper describes work with complex numerical model as a supplement for existing state of knowledge about rally driver safety.
2. Materials and methods
The implementation of numerical models was car- ried out in MADYMO® software, which allows analy- sis of dynamic phenomena, depending on the input parameters. In addition, this software is predisposed to simulate car accidents due to the library of dummy models Hybrid III. As a result, it was possible to de- velop a model of the driver along with the HANS safety system (Fig. 1) and conduct biomechanical analysis of injuries in rally accident. The validation of model was based on experimental tests, the results of which were compared with model results. The choice
Fig. 1. Numerical model
of numerical simulation as a research method is justi- fied by the fact that conducting impact tests in rally conditions is very difficult, mainly due to the consid- erable high speeds achieved in rallies.
2.1. Numerical modeling
The numerical model of the driver along with the HANS system was developed using the Madymo® soft- ware. It is widely used in the automotive industry for numerical modeling in projects including crash analysis.
The authors chose this environment with consideration of several of the numerous advantages, such as:
the calculation speed, due to the application of the , “multibody” method [16],
library with models of Hybrid III dummies,
prepared functions for calculating injury criteria during crash test simulation.
For conducting the study of numerical models Hybrid III 50th, the seat (test stand on sled cart PIMOT – Fig. 2) and models of seat belts (4-point and 5-point), helmet and HANS device were used. The configuration of particular models (the geometry of the seat, the position of the dummy, place of belt an- chors, etc.) is an accurate reflection of the experiment test used for the model validation.
For the input parameters for simulation cart accelera- tion pulse with peak acceleration of 13 G from experi- ment (Fig. 3) was used. The acceleration peak was intro- duced into the model with LOAD.SYSTEM_ACC element which allows time dependent acceleration field to be added on model elements. In this case,
Fig. 2. Fully equipped sled in PIMOT Institute
acceleration was added on sled body and acceleration signal was inserted from sled experiment, which ensures precise control of sled speed in numerical model.
Gravity load was implemented also with this method except that it was added on the whole model with value of 1 g. The results of model analysis were vali-
dated based on the results obtained during the first attempt of the experiment. The boundary conditions form the model corresponded with real experiment.
Contacts in the model are described in Table 1. In model characteristics of helmet foam, seat belts and HANS belts are presented in Fig. 4.
As a result of numerical simulations parameters describing the dynamics and kinematics of the dummy (the driver) were obtained. The data obtained allowed us to compare it with the corresponding data sources and sensors used in the experiment which also provided an opportunity to compare the kine- matics of the video recorded with high frame rate camera.
Mannequins from of Hybrid III (HIII) family are some of the best human surrogates used for crash tests by Euro NCAP. Different versions correspond to the human body of various sizes and ages, as well as pro- vide the ability to modify the application in specific tests (e.g., a mannequin for frontal impact is different in design from the side impact dummy). To ensure the most precise reflection of biological characteristics of
Fig. 4. Model characteristics of helmet foam, seat belts and HANS belts
Table 1. Description of numerical contacts in the model
Contact/Description Master surface Slave surface Contact type Characteristics Friction Between dummy and seat Dummy: thorax, pelvis,
femur, shoe Seat Master From dummy Generic friction: 0.4
Between dummy and belts Dummy: pelvis, thorax Belts Master From dummy
Orthotropic friction coefficient function:
Fx: 0.4 Fy: 0.6 Between dummy and HANS Dummy: thorax, neck HANS Master From dummy Generic friction: 0.6
Between HANS and belts HANS Belts Master Rigid
Orthotropic friction coefficient function:
Fx: 0.4 Fy: 0.6 Between belt buckle and dummy Dummy: pelvis, thorax Buckle Master From dummy Generic friction: 0.8 Between head
and helmet foam Dummy: head Foam Slave
Seat_foam_char (from software library)
Generic friction: 0.8
Between helmet
and HANS Helmet HANS Master Rigid Generic friction: 0.6
Fig. 3. Sled acceleration pulse registered during experiment, solid line first experiment, dotted line: second experiment
the human body, featured mannequins are created based on cadavers research. Many authors create complex models of human body based on cadaver research [8], [15].
In this paper, dummy Hybrid III 50th was se- lected, which corresponds to the dimensions and weight of the 50-percentile male with weight of approximately 77.7 kg. The mannequin is equipped with force and torque sensors, accelerometers, which allows the risk of injury during the test to be estimated, in which the main work load is the sag- ittal plane. HIII model used in Madymo® is vali- dated and its characteristics correspond to the actual weight of HIII, which makes it possible to reflect experiment involving real dummy.
2.2. Safety systems
For elementary driver safety devices with rally ap- proval the following systems were used:
4-point harness – 50 mm in width and 1 mm in thickness;
OMP Star rally helmet;
HANS device.
Safety belts were modeled as Finite Elements with adequate to real parameters and material characteris- tics. Madymo® software built-in function Belt Fitting Tool allows belts to route with no gaps, adjusting FE elements to tighten the belt to dummy parts as they do in sled experiment setup.
Helmet model was created by reverse engineering with RevScan handheld 3D laser scanner. Thanks to 3D scanning technology, the obtained model geome- try was almost identical to real object. Helmet shell was adopted as rigid facet type object and pads inside were modeled as multibody elements with foam char- acteristic.
During HANS device modeling process 3D scan- ning technology was also used. Shoulder support part, made of durable synthetic material, was created as rigid facet element. Tether was modeled as Belt using Finite Elements with dimensions and material char- acteristic of HANS tether.
2.3. Experimental research
Model validation was made during two tests in Vehicle Safety Laboratory of PIMOT Institute. The tests were carried out on sled which allows desirable
parameter of speed and deceleration to be achieved according to requirements described in [19]. Test- ing stand was made from sled placed on runner.
Sled was accelerated by rubber wire. Decelera- tion with required impulse was obtained by braking mechanism. This mechanism stops sled by press- ing metal knob fixed to ground through polyure- thane sleeve. A detailed description is to be found in [19].
Steel seat and floor elements were bolted to sled.
Steel crate allows seat belts to be mounted like in a rally car. Hybrid III dummy was set up with OMP Star hel- met and HANS device as described in [9]. A 4-point harness was fastened as thigh as possible without any tools – just like driver does it.
In the experiment, two tests were carried out, both with a speed of 30 km/h and a 60 cm braking distance.
In the first test, maximum acceleration was 12.67g, in the second test, it was 11.49 g (Fig. 3).
During the tests, the following data were regis- tered, which were also used in numerical model vali- dation:
sled acceleration,
head acceleration in x, y, z axis,
torso acceleration in x, y, z axis,
compression force in cervical spine,
torque in cervical spine,
sled and dummy kinematic recorded with high speed camera at 1000 fps rate.
Data was registered at a rate of 10 kHz and fil- tered using CFC 600 filter excluding sensors in dummy’s head and neck where CFC 1000 filtration was applied.
2.4. Identification of HANS belts mechanical properties
Identification of HANS device elements dumping and stiffness characteristics was essential for accurate model response. MTS Bionix strength test machine was used for this purpose. HANS tether was clipped by metal buckles, preliminary distance was set at 360 mm (30 mm closer than belt length) which reflects real life condition of use where tether is loose to let driver rotate his head (Fig. 5). The test was carried out with machine’s highest displacement speed of 0.4 m/s.
Force-elongation characteristic of belt was introduced into numerical model (Fig. 6). After the test, the in- spection showed that the belt was damaged near metal buckle which was deformed (Fig. 7). Belt stitching was not damaged.
Fig. 5. Stages of dynamic test, from the left before test – strap is not tight, then moment in which the greatest amount of force is achieved after which strap breaks
Fig. 6. Strap displacement vs. force characteristics
Fig. 7. The damage of the tether; on the left: deformation of ear buckles;
on the right: rupture of the material
2.5. Model validation
Numerical model validation was based on com- parison of plots and values of head and thorax ac- celeration (Fig. 8). Validation was further based on forces and torques in the upper and lower neck, forces in thorax, acceleration in pelvis. The model
was also validated by comparing dummy kinematics with high speed video record of experiment (Fig. 9).
Because main loads occur in sagittal plane only values in x axis were considered for validation.
Comparative analysis shows high correlation of kinematic and dynamic parameters between nu- merical model and experiments carried out in PIMOT laboratory.
Fig. 8. Model validation simulation (dashed line) vs. experiment (solid line), top: head acceleration in x axis, bottom: thorax acceleration
Fig. 9. Dummy kinematics comparison between experiment and simulation
2.6. Computer simulations
Validated model was subjected to several calcula- tions using increased value of acceleration in the fal- lowing variants:
W13g/4 – simulation with peak acceleration of 13g (validated with an experiment with 4-point belt system),
W45g/4 – simulation with peak acceleration of 45g with 4-point belt system,
W60g/4 – simulation with peak acceleration of 60g with 4-point belt system.
Variant W60g/4 was omitted from further analysis because during the simulation driver’s pelvis “slipped”
beneath the seat belt (Fig. 10).
Fig. 10. The moment when pelvis slips under the belt in simulation W60G4
Therefore, to simulate conditions in major accel- erations 5-point belt systems were used:
W60g/5 – simulation with peak acceleration of 60g with 5-point belt system,
W70g/5 – simulation with peak acceleration of 70g with 5-point belt system
In addition, in order to compare the impact of the HANS device on the driver safety at the highest ac- celeration of 70 g the authors prepared a variant with- out its use:
W70g/5-H – simulation with peak acceleration of 70g with 5-point belt system without HANS device.
3. Results
From simulations: W13g/4, W45g/4, W60g/5 and W70g/5 accelerations of head and torso of the dummy were obtained and also compressive force and bending torque acting on cervical spine. The maximum accelera-
tion for head was acquired for W45g/4 (Fig. 11). The maximum acceleration for the thorax was obtained for W70g/5 (Fig. 11). Maximum torque and compressive force were observed for variant W45g/4 (Fig. 12). Based on the data obtained injury criteria were determined:
HIC15, NTE, NTF, NCE, NCF (Table 2) [3]. The highest parameter of HIC15, NTE, NTF and NCF was obtained for W70g/5-H which was 3316 for HIC15, NTE – 1.36, NTF – 1.4, NCF – 1.37 (Table 2). Maxi- mum parameter for the criterion NCE was obtained for variant W70g/5 which was 0.17.
Comparing the results of the last two variants:
W70g/5 and W70g/5-H (Fig. 13) it can be noticed that the maximum acceleration in x axis is greater than for variant W70g/5. However, resultant acceleration was greater for W70g/5-H and HIC15 parameter was three times larger than in W70g/5. Maximum force acting on cervical spine in W70g/5-H was 4430 N, the torque during bending was equal to 134 Nm and for the ex- tension – 190 Nm. Maximum force in HANS belt was 4210 N.
Table 2. Injury Criteria in simulation scenarios (limits from [4])
Criteria/scenario Limit [4] W13G/4 W45G/4 W60G/5 W70G/5 W70G/5BH
HIC15 700 31 54 706 1174 3316
NTE 1 0 0.27 0.63 0.75 1.37
NTF 1 0 0.13 0.41 0.58 1.4
NCE 1 0 0 0 0.17 0
NCF 1 0.2 1.17 0.76 0.93 1.37
4. Discussion
The method of modeling presented in this paper is very useful for the analysis of rally driver injuries during crash. Simulation results are convergent with experimental results, shown in the diagram (Fig. 8). In addition, due to the repeatability of the test, numerical model has an advantage over crash test [16], including cost reduction. What is more, numerical modeling is not subjected to restrictions, such as crash test due to the risk of potential damage of the Hybrid III dummy.
Therefore, research conducted in PIMOT was carried out at a low speed of 30 km/h resulting in lower ac- celeration of highest 13 g peak which was enough to validate the model. Model created in previous re-
search [2] was validated based only on kinematics from high speed camera recording sled test. The methodology presented in this article allows more credible model validation than in the previous work.
The numerical model created allowed us to analyze numbers of scenarios with different parameters during high accelerations without damaging measuring stand like in F1 car sled test [6], which would result in in- correct results by changing experiment conditions.
Also testing dummy during high accelerations without
head supporting device could result in neck section damage.
Numerical model contains dynamic characteristics of dummy and other elements like seat belts. These characteristics are provided with software. However, in the authors’ model there was a need to define char- acteristics of head restrain including straps. It is es- sential to determine characteristics of material during dynamic test similar to conditions of further experi- ment. MTS Bionix strength test machine allowed
Fig. 11. Comparison of simulations W13G4, W45G4, W60G5, W70G5, from the top: head acceleration in x axis, thorax acceleration in x axis
characteristics to be gained with displacement speed of 0.4 m/s which was the fastest possible displacement speed of the machine. Certainly, a comparison of the test with higher speed would be suitable for deter- mining the level of characteristic validation.
The model allowed us to perform multiple simula- tions with different parameters. Based on comparative analysis of variants W16g/4, W45g/4, W60g/5 and
W70g/5 it has been shown that the 4-point belt system is an insufficient protection in an accident with an acceleration above 45g. In this situation, there is a high risk of displacement of the pelvis under the harness, which may increase the risk of injury of the head and neck as well as decreasing the effectiveness of the HANS system. In the case of using 5-point harness injury parameters for acceleration of 60g and 70g are
Fig. 12. Comparison of simulations W13g4, W45g4, W60g5, W70g5, from the top: force in cervical spine in z axis, torque in upper neck in y axis
smaller than in the case of using 4-point belt in variant with acceleration of 45g.
Comparing the results of simulation W70g/5 and W70g/5-H higher resultant head acceleration was no- ticed in variant W70g/5-H. This happens due to the high stiffness of the helmet-HANS-belt system, which reduces the displacement of the head towards the chest (Fig. 14). Bending moment and the compressive force acting on the cervical spine are significantly reduced by the use of HANS device. This is indicated by three times smaller criterion HIC15 obtained in simulation with peak acceleration of 70g. Studies have confirmed the effectiveness of the HANS supporting device, even at high accelerations of up to 70g. The maximum force in HANS belt (4210 N) was below maximum tension force (6973 N) obtained during experiment.
Fig. 14. Comparison of simulations with and without HANS device
Head restrain systems are commonly used in rally and race cars because of high possibility of crash with tremendous acceleration. Assembly of safety systems has to provide minimum body displacement to minimize the possibility of impact with surround- ing objects. It is a different approach with respect to safety systems like airbags with three point safety harness commonly used in civil cars. Wheel airbag is designed to suppress body impact. Alternatively head and neck supporting device with five point belt har- ness reduce head displacement and transfer force to other body parts, for example, increasing chest de- flection. This kind of approach seems to be adequate because of cervical spine vulnerability to flexion and extension during high speed impact [7], [8]. The following argument against airbags are ergonomic adjustment problems – typical car systems are de- signed for 50-percentile human and people outside this range may suffer from reduction of safety level during accident. It is also widely known that only proper seat, steering wheel and belt setup provide expected safety level.
The authors want to continue the development of the model towards identifying the factors that may affect the safety of rally driver. Experimental sled tests are carried out in isolated conditions, mainly in frontal crashes, which is required to validate numeri- cal model. However, this approach does not consider
Fig. 13. Comparison of simulations W70g5 and W70g5-H.
From the top: head resultant acceleration, force in upper neck, torque in upper neck (dashed line W70G5, solid line W70G5-H)
the specificity of rally crashes which occur with an- gular collisions and rolling of the vehicle. Therefore, future studies will be undertaken in order to create a model that reflects the interior of the vehicle and its equipment components. This would allow analysis of driver safety in conditions similar to the real ones.
Additional simulations are planned for accelerations affecting in different directions and also simulations of rally car rollover. Also, the authors plan to compare safety systems used in commonly used car with rally car safety features. This kind of diversity of simula- tions can only be achieved with low cost numerical simulations.
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