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Strength analysis of human skull on high speed impact

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Strength analysis of human skull on high speed impact

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This research has been carried out to identify the strongest and weakest part of the human skull on high speed impact condition by using finite element simulation approach. There are nine parts of skull model to be examined, i.e., frontal bone, front mandible, side mandible, maxilla, nasal bone, occipital bone, parietal bone, temporal bone and zygomatic bone. The parts are impacted by a small rigid ball at constant velocity of 20 m/s. Three dimensional Finite Element solver FEBio is employed to do the simulation. Four mechanical responses; acceleration, displacement, effective stress and effective strain are monitored after 1 ms the ball contacts with skull. When the response parameter shows the strongest part, it is scored 100 whereas the weakest the weakest is scored of 10. The total score from four parameters are then summed up. The results show the weakest among the cranial bones is the parietal bone. On the facial section, the front mandible and the nasal bone are the weakest parts among other facial bones
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International Review of Mechanical Engineering (I.RE.M.E.), Vol. 6, N. 7
ISSN 1970 - 8734 November 2012
Manuscript received and revised October 2012, accepted November 2012 Copyright © 2012 Praise Worthy Prize S.r.l. - All rights reserved
1508
Strength Analysis of Human Skull on High Speed Impact
Waluyo Adi Siswanto1, Chong Szu Hua2
AbstractThis research has been carried out to identify the strongest and weakest part of the
human skull on high speed impact condition by using finite element simulation approach. There
are nine parts of skull model to be examined, i.e., frontal bone, front mandible, side mandible,
maxilla, nasal bone, occipital bone, parietal bone, temporal bone and zygomatic bone. The parts
are impacted by a small rigid ball at constant velocity of 20 m/s. Three dimensional Finite
Element solver FEBio is employed to do the simulation. Four mechanical responses; acceleration,
displacement, effective stress and effective strain are monitored after 1 ms the ball contacts with
skull. When the response parameter shows the strongest part, it is scored 100 whereas the weakest
the weakest is scored of 10. The total score from four parameters are then summed up. The results
show the weakest among the cranial bones is the parietal bone. On the facial section, the front
mandible and the nasal bone are the weakest parts among other facial bones. Copyright © 2012
Praise Worthy Prize S.r.l. - All rights reserved.
Keywords: Human Skull Model, Finite Element, Strength Analysis, Impact Simulation, FEBio
I. Introduction
Head as well as knee [1] can be considered as the
most important part of the human body that must be
protected well. In the human head, a bony structure and
part of the skeleton which is human skull supports the
structures of the face and forms a cavity for the brain.
The adult skull is made up of 22 bones. These bones are
separated into two categories which form the cranium
and facial bones. All of the bones of the skull are joined
together by structures which rigid articulations
permitting very little movement except for the mandible.
Eight bones include one frontal bone, two parietal bones,
one occipital bone, one sphenoid, two temporals and one
ethmoid form the brain case. Other fourteen bones form
the splanchnocranium which is the bone supporting the
face.
The function of the skull are protect the brain, fix the
distance between the eyes to allow stereoscopic vision
and fix the position of the ears to help the brain use
auditory cues to judge direction and distance of sounds.
The skull protects the brain from damage through its
hard unyielding property and less deformable substances
in nature. The bruised or injured of brain can be life-
threatening.
A break in one or more of the bones in the skull that
cause by a result of blunt force trauma is skull fracture.
Skull fractures occur with head injuries. The direct
impact force that excessive the bone will cause fracture
at site of impact and damage the underlying physical
structures contained within the skull.
The brain function can be affected directly by damage
to the nervous system tissue and bleeding. The blood
clots under the skull and then compress the underlying
brain tissue which is subdural or epidural hematoma can
disturb the brain’s function.
Skull fracture is a frequently observed type of severe
head trauma caused by blunt impact. Kleiven and von
Hols [2] proposed skull fracture as an indicator of brain
injury. The parameters included peak impact force, local
skull deformation and absorbed energy until skull
fracture. Experimentally, the fracture of the skull was
first performed by Gurdjian et al. [3]. They used a free-
fall method on dry human skulls instrumented with
stress-coat, a strain-sensitive lacquer. Now, in this
present research, the nine parts of skull are investigated
to find the strongest and the weakest parts when the skull
is impacted. Finite element approach is used to do the
impact simulation and to observe the response of the
parts on impact.
Although there has been a lot of research of head on
impact but most of the skull fracture test is for impact
from front, rear or side on skull [4]–[9]. Skull thickness
is not uniform and the impact force that causes a fracture
depends on the site of impact. The structure and the
shape of the bones can have the effect on damping.
Hence, the objective of this research is to conduct an
investigation to see the strongest and the weakest parts of
the skull on impact. A protection system will then
consider the results and put an extra attention to the
weakest part of the skull.
II. Literature Review
The head impact response in terms of head
acceleration and impact force depend on the inertia
properties of the head and surface impacted. For a 50th
W. A. Siswanto, C. Szu Hua
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percentile male, the average head mass is 4.54 kg and the
average mass moments of inertia are Ixx = 22.0 × 10-3 kg
m2, Iyy = 24.2 × 10-3 kg m2, and Izz = 15.9 × 10-3 kg m2
[10].
Drop tests against a rigid flat surface were performed
in many studies. The peak force for fracture at different
regions of head are summarized in Table I.
Experimentally, accelerometers cannot be mounted at the
centre of gravity of the head and the head does not
behave as a rigid body. Measuring the acceleration
response of the head is almost impossible. According to
Padgaonka et al. [11], measurement on the head
rotational acceleration is recommended so that the
acceleration of the head’s centre of gravity can be
computed.
TABLE I
PEAK FORCE FOR FRACTURE AT DIFFERENT REGION OF THE SKULL
Impact Area Force (kN) Reference
Frontal 4.2 [4]
5.5 [6]
4.0 [7]
6.2 [12]
4.7 [9]
Lateral 3.6 [4]
2.0 [7]
5.2 [9]
Occipital 12.5 [13]
In order to focus on head acceleration tests, Wayne
State University Cerebral Concussion Tolerance Curve
abbreviated as the Wayne State Tolerance Curve
(WSTC) was established to indicate a relationship
between the duration and the average anteroposterior
translational acceleration level of the pulse that accounts
for similar head injury severity in head contact impact
[3].
Clinically it is observed that prevalence of
concomitant concussion in skull fracture cases was used
to relate cadaver impacts to brain injury which 80% of
all concussion cases had linear skull fracture [14]. The
tolerance of the skull to fracture loads is the parameter to
infer the tolerance to the brain injury that used.
The combination of acceleration level and pulse
duration that above the curve exceed the human
tolerance that cause severe, irreversible brain injury.
Combinations below the curve that do not exceed human
tolerance will cause reversible injury. Original WSTC
covers a time duration range of 6 ms only, the durations
longer than 6 ms was extended using animal and
volunteer data that represent as modified curve. Test
conditions used to obtained the data and seen that the
head can withstand higher acceleration for shorter
duration are shown in Table II [15].
The WSTC was supported by experiments conducted
in Japan which led to the Japan Head Tolerance Curve
(JHTC) obtained from experiments with primates and
scaling of results to humans [16]. The difference between
the WSTC and JHTC can be negligible for time intervals
up to 10 ms and only minor differences shown for longer
durations. The WSTC becomes a straight line with a
slope of -2.5 in a logarithmic scale.
TABLE II
TEST CONDITIONS OF THE EXPERIMENTS THE WSTC [15]
Pulse
duration Test
objects Test
setup Response
measured Injury
criterion
2-6 ms cadavers drop test acceleration
at the back
of the head
skull
fracture
6-20ms cadavers
and
animals
impact
test
acceleration
of skull,
brain
pressure
pathologica
l
changes
>20ms volunteers sled tests
whole body
acceleration
without
head impact
concussion,
state of
consciousn
ess
The restrictions for test conditions have to be
considered no matter using the WSTC or any criterion
developed thereof such as the paucity of the data points,
the position of the accelerometer (back of the head), the
fact that the rotation acceleration is not considered and
the techniques used to scale the animal data. However,
the main criticism concerns the correspondence of the
skull fracture and brain injury was assumed from
biomechanical point of view. This hypothesis remains to
be verified since there was no direct demonstration of
functional brain damage in an experiment which
biomechanical parameters sufficient to determine a
failure mechanism in the tissue were measured [14].
WSTC is based on direct frontal impact tests and cannot
be applied to non-contact loading conditions and other
impact directions. Furthermore, diffuse brain injury and
subdural hematoma caused by rotational acceleration was
studied. The head rotation acceleration was measured
and the resulting degree of injury on primates was
assessed by Ommaya et al. [17]. It was found that the
angular acceleration and the injury threshold are related
to the mass of the brain. Figure 1 shows the tolerance
limit for the human obtained from the primates tests
while Table III gives the tolerance values that use
commonly.
Fig. 1. Rotational acceleration adapted from [19]
W. A. Siswanto, C. Szu Hua
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TABLE III
TOLERANCE THRESHOLD FOR ROTATIONAL ACCELERATION AND
VELOCITY OF THE BRAIN
Tolerance threshold
50% probability Type of brain
injury Reference
α = 1800 rad/s2
for t < 20ms
α = 30 rad/s2
for t 20 ms
Cerebral concussion [17]
α < 4500 rad/s2 and/or
α < 30 rad/s2 Rupture of bridging
vein [20]
2000 < α < 3000 rad/s2 Brain surface
shearing [13]
AIS 2: α = 1700 rad/s2
AIS 3: α = 3000 rad/s2
AIS 4: α = 3900 rad/s2
AIS 5: α = 4500 rad/s2
General disturbance [21]
Tolerance values up to 2500 rad/s2 for short durations
was suggested on volunteers [18].
Even though the several experimental studies predict
head injury from one specific input parameter either in
translational or rotational acceleration, head impact
situations can be expected in both translational and
rotational accelerations combine to cause brain injury.
Comprehensive brain injury prediction include the
various responses of the brain tissue for any combination
of mechanical loading [15].
III. Methodology
In this present research, the strength investigation of
human skull is conducted on finite element numerical
approach. A three-dimensional of human skull model is
impacted by a rigid object with a constant velocity on
various part of skull model at normal direction. The
response of the skull parts are then assessed. There are
four parameters used for the assessment, i.e. acceleration,
displacement, effective stress and effective strain. Every
part is then scored on each parameter and the total score
represents the strength’s score of the part. Te parts of the
skull are compared each other in terms of the strength’s
score. There are three stages in conducting this research.
The first stage is human skull finite element modeling
followed by simulating of impact loading to the skull
then analyzing and scoring of assessment parameters.
III.1. Three Dimensional Geometrical Data
of Skull Model
The physical model of the skull model is taken from
the human-emulated skull model ZL-DF227,
manufactured by Fujian China. This life-size human
skull model contains three removable parts; lower part,
calvarium and movable jaw.
In order to obtain geometrical digital data, a three-
dimensional digitizing system ATOS (Advanced
Topometric Sensor) is used. After the scanning process a
multitude of X, Y, Z coordinates on the surface of
physical object are recorded. Each discrete X, Y, Z
coordinate is referred to as a point. The conglomeration
of all all points of the object provides point clouds and
saved as STL (STereoLithography) file format.
III.2. Finite Element Meshing
Converting the STL geometrical data and generating
finite element meshing are conducted in GiD finite
element pre-processor. To be able to read successfully
the geometrical data in GiD, the format is converted to
DXF format in MeshLab before sending the file to GiD.
The head skull geometrical data is then meshed to
provide three-dimensional solid tetrahedral elements in
NASTRAN format. The lower part consists of 68999
elements, calvarium has been modelled as 34327
elements and jaw has 87399 elements. Figures 2 show
the sections of skull model in meshing condition in GiD
pre-processor.
(a) Lower part, 68999 tetrahedral
(b) Calvarium, 34327 tetrahedral (c) Jaw, 87399 tetrahedral
Figs. 2. Skull finite element model
This finite element NASTRAN format is then
converted to Gmsh format for the finite element solver
FEBio. The definition of material properties, boundary
conditions and other simulation parameters is controlled
under FEBio PreView.
III.3. Finite Element Analysis Modeling
The skull bone material properties follows Donnelly
and Medige [22] as listed in Table IV. In FEBio, the
material properties are implemented under neo-Hookean
material. A spherical rigid object with the diameter
15mm is used as the impacting object to the skull. The
velocity of the impactor is at constant velocity of 20 m/s.
During the impact simulation, the skull is in
unconstrained condition except hold at the bottom, fixed
around occipital condyle.
W. A. Siswanto, C. Szu Hua
Copyright © 2012 Praise Worthy Prize S.r.l. - All rights reserved International Review of Mechanical Engineering, Vol. 6, N. 7
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TABLE IV
MATERIAL PROPERTIES (NEO-HOOKEAN) FOR SKULL MODEL
Components E (kPa) ν ρ (kg/m3)
Skull upper 7.3 106 0.22 3000
Skull lower 7.3 106 0.22 3000
Facial bone 7.3 106 0.22 2700
mandible 7.3 106 0.22 2700
TABLE V
CONTACT AND CONSTRAINTS FOR SKULL MODEL
Part Elements Material Contact Constraints
Calvarium 34327 Neo-
Hookean Sliding
and tied -
Lower part 68999 Neo-
Hookean tied
Fixed
around
occipital
condyle
Jaw 87399
Neo-
Hookean tied -
Impactor 2048 Rigid
Body sliding Rigid body
with initial
velocity
Tied contact algorithm is used to connect the
calvarium, lower part and jaw. This tied connection glues
tightly the nonconforming finite elements. The material,
contact pairs and the boundary conditions are listed in
Table V.
A sliding contact is defined between the impactor
surface and calvarium surface. The sliding contact makes
sure that the contacting surfaces may slide across each
other but they do not penetrate.
There are three tied interface used as the contact
interface within connection of skull model. The tied
contact connects the calvarium, face and jaw at the
connecting surfaces. The properties of the contact
parameters during the simulation are documented in
Table VI.
IV. Results and Discussion
The total displacement, total acceleration, effective
stress and effective stress were recorded after 1.0 ms the
rigid impactor contacted with the part. This time duration
of 1.0 ms was selected to represent a condition just after
being contacted. The surrounding structures were not
effected in this short duration so that the strength of
individual part can be examined.
The simulation of impact test using a rigid ball
impactor is shown in Fig. 3.
A typical displacement record of front mandible is
shown in Fig. 4. The finite element that was contact with
the impactor is element 40686. The displacement was
tracked in this element. Since the impactor contacted
with the front mandible (represented element 40686) at
0.5 ms, then the displacement of the front mandible is
recorded after 1 ms. The impact displacement was
obtained at 1.5 ms.
TABLE VI
PROPERTIES OF CONTACT
Contact Sliding Tied 1 Tied 2 Tied 3
Master Impactor
surface
Lower
surface
aroud
calvarium
Upper
left hand
side tooth
surface
Upper
right hand
side tooth
surface
Slave Skull
surface
Upper
surface
around face
Lower
left hand
side tooth
surface
Lower
right hand
side tooth
surface
Augmented
Lagrangian No No No No
Aug.
tolerance 0.01 1 1 1
Penalty 40 1 1 1
Two-pass No - - -
Auto-
penalty No - - -
Friction
coefficient 0 - - -
Penalty
factor
search
tolerance
0.01 - - -
Contact
type Node-on-
facet - - -
The illustration of front mandible in impact is
captured in Fig. 5.
Similar technique was applied to obtain other
assessment parameters; acceleration, effective stress and
effective strain of every part. The element number to
represent the part in contact is documented in the second
column in Table VII. The complete results are collated in
Table VII and the comparison is shown in Fig. 6.
Fig. 3. Impact location
W. A. Siswanto, C. Szu Hua
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Fig. 4. Recording displacement front mandible
IV.1. Displacement
The results show that side mandible which is
commonly said as jaw is very flexible. In 1 ms after the
contact, it displaced 4.089 mm. Nasal bone is also easily
deformed. In real life, the nasal bone is the weakest part
among the parts of skull where easily fracture or injuries
when impact on it. Maxilla on the lower jaw shows the
smallest displacement on impact. This means the maxilla
part does not easily deform compared to other side part,
side mandible. The lower jaw is stiff on the front but less
stiff on the side part.
Fig. 5. Displacement response of front mandible
IV.2. Acceleration
The side mandible does not move with high
acceleration. Compared to other parts, the side mandible
is the lowest in its acceleration motion. Zygomatic bone
in human upper cheek shows similar characteristics with
low acceleration. The front mandible was recorded with
the highest acceleration. The occipital bone shows
similar response with acceleration movement. The
acceleration at the maxilla and nasal bone are
approximately the same because of they are close to each
other and impact condition are same.
IV.3. Effective Stress
Since the side mandible is flexible showing high
deformation, therefore this part is not under high stress.
The level of effective stress is the lowest among others.
The zygomatic bone has shown the highest stress.
IV.4. Effective Strain
The characteristics of the strain proportionally relate
to the stress. The effective strain results are consistent
with the effective stress results. The highest strain is on
the zygomatic bone whereas the lowest strain is at the
side mandible.
IV.5. Strength Assessment
Since the strength considers four parameters, then the
scoring criterion was applied to each parameter. For
every parameter, when the level of the assessment
parameter was at the lowest compared to others, a full
score of 100 was assigned to the part. This full score
indicated the part is the strongest one. On the contrary,
when the level was at the highest level among others, the
lowest score of 10 was assigned to the part. Other parts
were given scores in between 10 and 100 according the
ranking.
The scoring results of the skull parts are shown in
Table VIII based on data results from Table VII. The
comparison of the strength parameters are shown in Fig.
6. The total scores of skull parts are also summarized in
the table.
TABLE VII
ASSESSMENT RESULT
Part Repre-
senting
Element
Total
accel
(m/s2)
Total
displ
(mm)
Effective
stress
(GPa)
Effective
strain
Frontal
bone 99479 303.1 3.724 0.3462 0.055
Front
mandible 40686 6785 3.122 1.19 0.1635
Side
mandible 182588 112.1 4.089 0.0991 0.0164
Maxilla 116965 458.3 2.371 0.7386 0.1095
Nasal
bone 59248 491.7 4.03 0.8089 0.13
Occipital
bone 102844 2326 2.384 0.4279 0.06721
Parietal
bone 22769 686.2 2.858 0.4639 0.07224
Temporal
bone 42654 931 3.296 0.2395 0.03941
Zygomati
c bone 88476 202.5 3.29 1.278 0.18
The first strength rank is the side mandible, followed
by second rank the frontal mandible and the temporal
bone. The following ranks are the maxilla (rank 3), the
occipital bone (rank 4), the parietal bone (rank 5), the
zygomatic bone (rank 6), the nasal bone (rank 7) and the
last one (rank 8) is the front mandible.
The skull is divided into two regions which are the
cranial section and facial section. The cranial bones
consist of the bones in the top of the skull which protect
the brain while the facial bones consist of the bones that
make up human face.
W. A. Siswanto, C. Szu Hua
Copyright © 2012 Praise Worthy Prize S.r.l. - All rights reserved International Review of Mechanical Engineering, Vol. 6, N. 7
1513
TABLE VIII
ASSESSMENT STRENGTH SCORE
Part Accel. Displ.
Eff.
stress Eff.
strain Total
score
Stren
gth
Rank
Front
mandible 10 70 30 30 140 8
Side
mandible 100 30 100 100 330 1
Maxilla 70 100 50 50 270 3
Nasal
bone 60 10 40 40 150 7
Zygomati
c bone 90 50 10 10 160 6
Parietal
bone 50 80 60 60 250 5
Occipital
bone 30 90 70 70 260 4
Temporal
bone 40 60 90 90 280 2
Frontal
bone 80 40 80 80 280 2
Fig. 6. Comparison result
The cranial bones tested in this study are parietal
bone, frontal bone, occipital bone and temporal bone.
The facial section consists of nasal bone, zygomatic
bone, side mandible, front mandible, and maxilla.
In cranical bones that protect the brain, the parietal
bone is the weakest part, whereas the temporal bone and
the frontal bone are the strongest parts. In facial section,
the front mandible and the nasal bone are the weakest
part. This result indicates that on high impact, the impact
from the top of the head and the front of the head are
critical.
V. Conclusion
According to the impact analysis of four assessment
parameters, the weakest part of head at front mandible
having the lowest score of 140 and the strongest part of
skull with the overall score of 330. The skull is divided
into two regions which are the cranial section and facial
section.
The weakest among the cranial bones is the parietal
bone. Therefore, parietal bone is the part of skull that
should be protected well to avoid from brain damage
since the structure is not as strong as other cranial bones.
Furthermore, the front mandible and the nasal bone are
the weakest parts among other facial bones. This result
indicates that impact on the top and the front of the head
is critical. Head protective system should consider these
parts for more protected.
This study is important for the prevention and
investigation of head on impact. The study also can help
by giving the information in manufacturing the helmet
for rider and the combat sport or the precaution during
the accidents.
For the future research, the assessment on impact
should be considering the fracture criteria to see the
strength of the skull parts. Increasing the number of
element will be another consideration to refine the
acquired response on impact.
References
[1] A.S. Ismail, “Mathematical description for human knee joint
geometry and its effect on lubrication mechanisms,” International
Review of Mechanical Engineering (IREME), Vol 4, Issue 6, pp.
761-779, 2010.
[2] S. Kleiven and H. von Holst, “Consequences of head size
following trauma to the human head,” Journal of Biomechanics,
vol. 35, no. 2, pp. 153–160, 2002.
[3] E. Gurdjian, H. Lissner, R. Latimer, B. Hadded, and J. Webster,
“Quantitative determination of acceleration and intercranial in
experimental head injury,” Neurology, vol. 3, no. 6, pp. 417–423,
1953.
[4] A. Nahum, J. Gatts, C. Gadd, and J. Danforth, “Impact tolerance
of the skull and face,” SAE Technical Papers, Tech. Rep. SAE
680785, 1968.
[5] R. Nahum, A.and Smith and C. Ward, “Intracranial pressure
dynamics during head impact,” SAE Technical Report, Tech. Rep.
SAE 770922, 1977.
[6] R. Voigt and L. M. Thomas, “Breaking strength of the human
skull vs impact surface curvature,” Wayne State University
School of Medicine, Dept. of Neurosurgery, Tech. Rep. Accession
Number 00262859, 1974.
[7] D. C. Schneider and A. Nahum, “Impact studies of facial bones
and skull,” SAE Technical Papers, Tech. Rep. SAE 720965, 1972.
[8] S. Advani, A. Ommaya, and W. Yang, Human Body Dynamics.
Oxford Clarendon Press, New York, 1982, ch. Head injury
mechanisms-characterisations and clinical evaluation, pp. 3–37.
[9] D. Allsop, C. Warner, M. Wille, D. Schneider, and A. Nahum,
“Facial impact response - a comparison of the hybrid iii dummy
and human cadaver,” SAE Technical Papers, Tech. Rep. SAE
881719, 1988.
[10] G. Beier, E. Schuller, M. Schuck, C. Ewing, E. Becker, and D.
Thomas, “Center of gravity and moments of inertia of human
head,” in 5th International Conference on the Biokinetics of
Impacts, 1980, pp. 218–228.
[11] A. J. Padgaonka, K. W. Krieger, and A. I. King, “Measurement of
angular acceleration of a rigid body using linear accelerometers,”
Journal of Applied Mechanics, vol. 42, pp. 552–556, 1975.
[12] S. H. Advani, J. , Huston, and S. Ojala, “Human head impact
response experimental data and analytical simulations,”
Transportation Research Board, Tech. Rep. Accession Number
00134132, 1975.
[13] S. H. Advani, A. K. Ommaya, and W. J. Yang, Human Body
Dynamics. Oxford University Press, 1982, ch. Head Injury
Mechanisms, pp. 38–50.
[14] J. Melvin and J. Lighthall, Accidental Injury - Biomechnics and
Prevention. Springer Verlag, New York, 2002, ch. Brain injury
biomechanics, pp. 301–317.
[15] K.-U. Schmitt, P. F. Niederer, M. H. Muser, and F. Walz, Trauma
Biomechanics Accidental injury in traffic and sports 2nd edition.
Springer, 2007.
W. A. Siswanto, C. Szu Hua
Copyright © 2012 Praise Worthy Prize S.r.l. - All rights reserved International Review of Mechanical Engineering, Vol. 6, N. 7
1514
[16] K. Ono, A. Kikuchi, M. Nakamura, H. Kobayashi, and N.
Nakamura, “Human head tolerance to sagittal impact reliable
estimation deduced from experimental head injury using
subhuman primates and human cadaver skulls,” SAE Technical
Papers, Tech. Rep. SAE 801303, 1980.
[17] A. Ommaya, P. Yarnell, A. E. Hirsch, and E. H. Harris, “Scaling
of experimental data on cerebral concussion on sub-human
primates to concussion threshold for men,” SAE Technical papers,
Tech. Rep. SAE 670906, 1967.
[18] X. Trosseille, C. Tarriére, F. Lavaste, F. Guillon, and A. Domont,
“Development of a fem of the human head according to a specific
test protocol,” SAE Technical Papers, Tech. Rep. SAE 922527,
1992.
[19] G. Krabbel, “Ein rechnerisches schädel-hirn-modell zur
untersuchung dynamischer belastungen des kopfes, dissertation,”
Ph.D. dissertation, TU Berlin, 1997.
[20] P. Löwenhielm, “Mathematical simulation of gliding contusions,”
Journal of Biomechanics, vol. 8, no. 6, pp. 351–356, 1975.
[21] Ommaya, Biomechanics of head injury in Biomechanics of Trau,
M. Nahum, Ed. Appleton-Century-Crofts, Norwalk, 1984.
[22] B. Donnelly and J. Medige, “Shear properties of human brain
tissue,” Journal of Biomechanical Engineering, vol. 119, pp. 423–
432, 1997.
Authors’ information
1BioMSE, Dept. Engineering Mechanics,
Faculty of Mechanical and Manufacturing Engineering,
Universiti Tun Hussein Onn Malaysia.
2Dept. Engineering Mechanics,
Faculty of Mechanical and Manufacturing Engineering,
Universiti Tun Hussein Onn Malaysia.
Waluyo Adi Siswanto is an Associate Professor
at the Dept Engineering Mechanics and at
Biomechanics and Sport Engineering Research
Center (BioMSE), Faculty of Mechanical and
Manufacturing Engineering, Universiti Tun
Hussein Onn Malaysia (UTHM). He received his
bachelor degree in mechanical engineering from
Gadjah Mada University, Indonesia, and
completed his M.Eng and PhD from RMIT, Australia. He is one of the
contributors in an open source Finite Element Program Suite, Impact
(http://impact.sourceforge.net/). Dr. Siswanto is a member of SAE
Australasia, ESAFORM and senior member IACSIT.
Chong Szu Hua received her bachelor degree in
Mechanical Engineering with honors in 2012.
She was one of researchers at Biomechanics and
Sport Engineering Research Center (BioMSE),
Faculty of Mechanical and Manufacturing
Engineering, Universiti Tun Hussein Onn
Malaysia (UTHM).
... Therefore, defining fixed values for limits of fracture tolerance [N/m 2 ] does not seem suitable for case assessment without accurate differentiation of the fracture localization, e.g. Siswanto and Hua present a summarizing table on the peak force for fracture at different regions of head [21]. The authors differentiate between frontal, lateral, and occipital fracture localizations, possibly limited by the fact that the results (frontal 4.2-6.2 ...
... hammer blow is more than just a theoretical one [3]. If the results of Siswanto et al. [21] are used as a basis for these considerations and if a fracture risk begins in fact at any stroke with >2 kN impact force, then a skull fracture must be expected in more than 90% of cases with lateral location of the impact on the head, regardless of the (a) age and (b) sex of the aggressor. ...
Article
Hammer blows cause serious, often fatal injuries, especially when massive blunt violence is targeted at the head region. The evaluation of the injury potential depends not only on the body region hit, but also on the characteristics of the hammer as a weapon and on the physical characteristics of the attacker. This study aimed at elucidating the dependency between the physical constitution of a perpetrator and the intensity of hammer blows, thus to verify or refute this seemingly obvious interrelation sometimes expressed in the saying that a “strong hand strikes harder”. For this purpose, 113 volunteers of different ages and sexes took part in different experimental settings. While, as expected, clear differences between male and female were detectable in the striking power of single and multiple strokes, there were no age or sex differences with regard to the maximum number of strokes per time unit. Strength differences in slamming with a hammer between men and women exceeded expectations in this study. Using the fracture forces as described by Sharkey et al. in an exemplary manner, one can expect a fracture of the skull in 9 out of 10 cases with a 300 g hammer by men for intensively executed single strokes, whereas this was only the case for approx. 2/10 women in this study. The maximum circumference of the upper arm and the width of the shoulder girdle correlate significantly with the achievable impact forces of individual hammer blows in both sexes. A simple measurement of the hand force with a manometer using the regression formula y [kN] = 0.144 × manual grip force −1.08 can be used as a rough estimation parameter for the theoretically achievable impact force. If one strikes repeatedly with the same hammer for 1 min, the magnitude of a single strike decreases continuously from 4.5 kN to 2.6 kN on average. If a 1500 g hammer is used instead of a 300 g hammer, one does not get the fivefold impact force you might expect at first sight, but only on the order of twice the impact force, about 14 kN on average. The results prove the importance of physical experiments, whose results can help to better interpret the magnitude and effects of hammer blows as a form of potentially life-threatening violence.
... Direct impact of the head on the tatami is a major cause of head and neck injury [52]; it accounts for approximately 60% of ASDH in judo [38]. Impact responses of the head have frequently been described in terms of acceleration in cadaver and mechanical studies [53,54]. The current gold standard to assess head injury is the head injury criterion (HIC), determined by translational rotation. ...
Article
Full-text available
Objectives: To investigate the biomechanics of Ukemi in relation to head and neck injury in adult judokas with varying skill sets. Design: Narrative systematic review. Methods: An extensive literature search was performed using PubMed, Google Scholar, Science direct and EMBASE from inception to April 2021. Studies were included if they: (1) reported biomechanical analysis of judo throws and Ukemi; (2) were on adult judoka populations; (3) discussed injury related to judo technique. The included studies were assessed for risk of bias using a five-part modified STROBE checklist. A narrative synthesis was performed due to the heterogeneity of included studies. Results: 173 titles and abstracts were screened with 16 studies (158 judokas, 9 of which were female) included. All studies used 3D biomechanical analysis to assess Ukemi. Ukemi implementation produced reduced kinematic data in comparison to direct occipital contact, which was always below the injury threshold. Analysis of lower limb and trunk kinematics revealed variances in Ukemi between novice and experienced judoka. Whilst no significant differences were seen in neck flexion angles, hip, knee and trunk angle time plots revealed greater extension angles in experienced judokas. Conclusions: Ukemi is essential in preventing head and neck injuries; however, technique differs between experienced and novice judoka. Larger flexion angles of the hip, knee and trunk are seen in novice judoka, which correlate with increased kinematic data. The association of greater neck muscle strength with improved Ukemi is weak. However, a negative correlation was established between fatigue and breakfall skill by one study.
... However, race-dependent changes in parietal skull bone thickness have also been described [30,31]. The parietal bone is not considered as strong as other cranial bones and should be well protected [32]. The biomechanics of the cranial bone varies greatly depending on, for example, method of force application, impact speed, the integrity of the skull, age, sex, and race. ...
Article
Full-text available
Cranioplasty with freehand-molded polymethylmethacrylate implants is based on decades of experience and is still frequently used in clinical practice. However, data confirming the fracture toughness and standard biomechanical tests are rare. This study aimed to determine the amount of force that could be applied to virtually planned, template-molded, patient-specific implants (n = 10) with an implant thickness of 3 mm, used in the treatment of a temporoparietal skull defect (91.87 cm2), until the implant cracks and finally breaks. Furthermore, the influence of the weight and porosity of the implant on its force resistance was investigated. The primary outcome showed that a high force was required to break the implant (mean and standard deviation 1484.6 ± 167.7 N), and this was very strongly correlated with implant weight (Pearson’s correlation coefficient 0.97; p < 0.001). Secondary outcomes were force application at the implant's first, second, and third crack. Only a moderate correlation could be found between fracture force and the volume of porosities (Pearson’s correlation coefficient 0.59; p = 0.073). The present study demonstrates that an implant thickness of 3 mm for a temporoparietal skull defect can withstand sufficient force to protect the brain. Greater implant weight and, thus, higher material content increases thickness, resulting in more resistance. Porosities that occur during the described workflow do not seem to reduce resistance. Therefore, precise knowledge of the fracture force of polymethylmethacrylate cranial implants provides insight into brain injury prevention and serves as a reference for the virtual design process.
... The biomechanical characteristics of skull fracture have been studied widely. Some important mechanical parameters have been discussed such as force, sectional elastic modulus, energy absorbed, shear stress, acceleration deformation and bending stress [3] [4][5] [6]. It was found that velocity, strain rate, sampling position and intercranial variation have a marked effect on some of the mechanical parameters [5]. ...
... Among different types of sport injuries, one of the most frequent, and the most dangerous at the same time, are head injuries (Bullock et al., 1999;Rhee et al., 2002;Rubin and Winograd, 2002;Echlin et al., 2005;Gomes et al., 2006;Ceallaigh et al., 2007;Siswanto and Hua, 2012). The consequences of these injuries can be very hazardous to health and even life of competitors, because the head is a place where vitally important organs, like brain or organs of sight, hearing and balance, are located. ...
... Thus, the pushout strength provided by TTCP-PS in this study (> 3000 N) would provide strength closer to that of the native skull than would hardware fixation. 11,12 In this study, animals were not euthanized early to compare differences in flap fixation strength between the test groups and control groups before 12 weeks. However, prior (unpublished) testing in sheep cadaver skulls demonstrated that the condition in test group 1 (100% kerf fill) produced greater initial fixation strength than did plates and screws. ...
Article
OBJECTIVE The authors’ goal in this study was to investigate the use of a novel, bioresorbable, osteoconductive, wet-field mineral-organic bone adhesive composed of tetracalcium phosphate and phosphoserine (TTCP-PS) for cranial bone flap fixation and compare it with conventional low-profile titanium plates and self-drilling screws. METHODS An ovine craniotomy surgical model was used to evaluate the safety and efficacy of TTCP-PS over 2 years. Bilateral cranial defects were created in 41 sheep and were replaced in their original position. The gaps (kerfs) were completely filled with TTCP-PS (T1 group), half-filled with TTCP-PS (T2 group), or left empty and the flaps fixated by plates and screws as a control (C group). At 12 weeks, 1 year, and 2 years following surgery, the extent of bone healing, local tissue effects, and remodeling of the TTCP-PS were analyzed using macroscopic observations and histopathological and histomorphometric analyses. Flap fixation strength was evaluated by biomechanical testing at 12 weeks and 1 year postoperatively. RESULTS No adverse local tissue effects were observed in any group. At 12 weeks, the bone flap fixation strengths in test group 1 (1689 ± 574 N) and test group 2 (1611 ± 501 N) were both statistically greater (p = 0.01) than that in the control group (663 ± 385 N). From 12 weeks to 1 year, the bone flap fixation strengths increased significantly (p < 0.05) for all groups. At 1 year, the flap fixation strength in test group 1 (3240 ± 423 N) and test group 2 (3212 ± 662 N) were both statistically greater (p = 0.04 and p = 0.02, respectively) than that in the control group (2418 ± 1463 N); however, there was no statistically significant difference in the strengths when comparing the test groups at both timepoints. Test group 1 had the best overall performance based on histomorphometric evaluation and biomechanical testing. At 2 years postoperatively, the kerfs filled with TTCP-PS had histological evidence of osteoconduction and replacement of TTCP-PS by bone with nearly complete osteointegration. CONCLUSIONS TTCP-PS was demonstrated to be safe and effective for cranial flap fixation in an ovine model. In this study, the bioresorbable, osteoconductive bone adhesive appeared to have multiple advantages over standard plate-and-screw bone flap fixation, including biomechanical superiority, more complete and faster bony healing across the flap kerfs without fibrosis, and the minimization of bone flap and/or hardware migration and loosening. These properties of TTCP-PS may improve human cranial bone flap fixation and cranioplasty.
... Over the years, numerous numerical models of the human head have been developed to understand the skull fracture and injury mechanism in the brain. [1][2][3][4][5] TBI affects a huge amount of the global population, ranging from mild TBI (concussion) to coma. Globally, 50 million people are injured every year with a projected death count of 1.2 million. ...
Article
Traumatic brain injuries (TBI) are life threatening injuries that can lead to long term incapacitation and death. Over the years, numerous finite element human head models have been developed to understand the injury mechanisms of TBI. Many of these models are erroneous and used ellipsoidal or spherical geometries to represent brain. The present work is focused on the development of high quality, comprehensive three-dimensional finite element human head model with accurate representation of cerebral sulci and gyri structures in order to study traumatic brain injury mechanisms. Present geometry, predicated on Magnetic Resonance Imaging (MRI) data consist of three rudimentary components i.e. Skull, Cerebrospinal Fluid (CSF) with the ventricular system, and the soft tissues comprising of the Cerebrum, Cerebellum, and Brain Stem. The brain is modeled as a hyperviscoelastic material. Meshed model with ten nodes modified tetrahedral type element (C3D10M) is validated against two cadaver-based impact experiments by comparing the intracranial pressures (ICP) at different locations of the head. Our results indicate a better agreement with cadaver results, specifically for the case of frontal and parietal ICP values. Existing literature focuses mostly on ICP validation, while the effects of von Mises stress on brain injury are not analyzed in detail. In the present work, a detailed interpretation of neurological damage resulting from impact injury is performed by analyzing von Mises stress and ICP distribution across numerous segments of the brain. A reasonably good correlation with experimental data signifies the robustness of the model for predicting injury mechanisms based on clinical predictions of injury tolerance criteria.
... Abaqus/Explicit has been chosen for running simulations with increment value of 56ns and above. The skull modeled as a single layer is considered as linear elastic material with high material density and elastic modulus [3] as presented in Table I. Ideally, to study the mechanism of CSF leakages [11] in brain the cerebrospinal fluid should be considered as newtonian biological fluid which has consistency similar to water (as 99% of CSF is water) [12]. ...
... Approximately 11-40% of all sports injuries involve the face. These injuries are most often due to direct hits with a ball or player-to-player contacts [8,12,14]. The most common types of sports-related facial trauma are the soft tissue injuries and the fractures of the "T-Zone" bones (the nose, the zygoma, and the mandible). ...
Chapter
Full-text available
Head injuries, due to the presence of the brain and sense organs, especially of the sense of sight, constitute a very serious threat to the health, and sometimes even life. The main causes of these injuries are road traffic accidents, physical violence, as well as different sport activities. The article presents a study on the effects of dynamic forces, acting on the bones of the skull around the eye socket, while hitted by a baseball, golf or tennis ball. In the research the variability in the force magnitude during the strike, as well as its various action pathways have been taken into account. For determination of deformations and stresses arising in bone structures of the skull in the vicinity of the operations, the finite element method has been used. The effect of numerical simulations is an indication of the places with the highest fracture probability, as well as the moment of the highest stresses occurrence. The results obtained during the investigations can be useful for the development of the construction of the specialized skull protectors, dedicated for people participating in different sport activities.
Chapter
This study proposes a detailed biomechanical model of the human head to study the effect of non-penetrating (blunt) head impacts of different durations on intracranial organs. A patient-specific high biofidelity three-dimensional finite element human head model is developed from the magnetic resonance imaging (MRI) data and segmented into five volumes namely, skull, cerebrospinal fluid (CSF) with ventricular system, cerebrum, cerebellum, and brain stem and each segment is assigned with appropriate material properties. The model validated against the impact experiment based on human cadaver is used to perform simulation with a range of blunt impact durations and biomechanical analysis is performed by investigating the maximum intracranial pressure (ICP) and von Mises stress distribution across the brain. The probability of loss of consciousness and tissue damage is studied based on the ICP and von Mises stress values. The coup and contrecoup phenomena is also studied with the localization, extension and, intensity of tissue damage based on the injury tolerance criteria present in the literature.
Article
Full-text available
The computation of angular acceleration of a rigid body from measured linear accelerations is a simple procedure, based on well-known kinematic principles. It can be shown that, in theory, a minimum of six linear accelerometers are required for a complete definition of the kinematics of a rigid body. However, recent attempts in impact biomechanics to determine general three-dimensional motion of body segments were unsuccessful when only six accelerometers were used. This paper demonstrates the cause for this inconsistency between theory and practice and specifies the conditions under which the method fails. In addition, an alternate method based on a special nine-accelerometer configuration is proposed. The stability and superiority of this approach are shown by the use of hypothetical as well as experimental data.
Article
The geometry of the human-knee joint has been described mathematically by dividing the surface of the femur and tibia into principal shapes through the aid of MR images. The cartilage geometry is described to be a good guidance for the production of artificial one. The lubrication under squeeze action in flexion position is described by Reynolds equation to predict the pressure generation between femur and tibia. The rheology of the synovial fluid is described by using data from the literature. The pressure distribution is assumed in the form of power function in the radial direction, and the power of the function depends on the circumferential direction and the meniscus contour shape. The pressure power function substituted in Reynolds equation and solved, using Mathcad software, to obtain the power as a function in the maximum pressure, the radii of curvatures of meniscus contour, the average film thickness, the synovial fluid dynamic viscosity, and the approaching velocity in axial direction between the femur and tibia. The purpose of the work is to improve the artificial knee geometry for the sake of reliable joint lubrication and performance.
Book
Trauma biomechanics uses the principles of mechanics to study the response and tolerance level of biological tissues under extreme loading conditions. Through an understanding of mechanical factors that influence the function and structure of human tissues, countermeasures can be developed to alleviate or even eliminate such injuries. Trauma Biomechanics surveys a wide variety of topics in injury biomechanics including injury classification, injury mechanisms and injury criteria. Both injuries sustained in automotive accidents and in sports are addressed. The interdisciplinary approach necessary in trauma biomechanics is stressed by showing the span from anatomy for each body region to engineering solutions for protection against injury. Injury tolerance values are listed, either currently in use or proposed by both the U.S. and European countries. Although the book is meant as a first introduction for engineers and medical doctors, sufficient references for scientific research are provided also. © Springer-Verlag Berlin Heidelberg 2010. All rights are reserved.