International Journal of Crashworthiness
Vol. 13, No. 3, June 2008, 265–278
Design of ventilated helmets: computational fluid and impact dynamics studies
Praveen K. Pinnoji, Zafar Haider and Puneet Mahajan∗
Department of Applied Mechanics, Indian Institute of Technology Delhi, New Delhi, India
(Received 13 December 2007; final version received 14 January 2008)
The existing motorcycle helmets are thermally uncomfortable as there is no provision for air flow inside the helmet. A new
design of helmet, with grooves in the liner foam and slot in the outer shell and liner foam to improve the ventilation, is
proposed. Computational fluid dynamics studies show considerable improvement in air velocities inside the helmet in the
presence of grooves and slot. Impact dynamics and biomechanics studies on various motorcycle helmets with deformable
and rigid head show that the proposed design meets the requirements of the standards in the drop test. The optimum size of
the groove and number of grooves for a motorcycle helmet are decided on the basis of the above studies.
Keywords: helmets; design; fluid flow; impact; finite element; biomechanics
1. Introduction
A well-designed motorcycle crash helmet has proved to be a
very good protection device for the rider to prevent or minimise the head injuries in road accidents. If a helmet is not
worn, the head impact with any object would cause localised
high pressure on the skull, which leads to brain injury. The
helmet design can be divided into functional (like shockabsorbing capability, penetration resistance, retention and
reliability) and non-functional (like low cost, good aesthetics, comfort, light weight and good thermal characteristics)
categories. Though a helmet is well-designed for functional
characteristics, because of weak non-functional characteristics drivers sometimes dislike wearing it while riding. In
South Asia, excessive sweating and resulting discomfort
due to hot and humid weather conditions discourage motorcycle riders from using helmets unless it is mandatory
by law. The space between the head and helmet is small,
and both mass flow and air velocities in this gap are also
low; as a result, the sweat is unable to evaporate making the
rider uncomfortable. The discomfort caused by sweating
can be reduced by increasing the air velocities inside the
helmet so as to enhance the sweat evaporation rate. Air flow
in helmet can be improved by large ventilation openings
as in bicycle helmets, but unfortunately such ventilation
openings may be detrimental to the safety and structural
integrity of the helmet. Most of the studies on motorcycle helmets are based on the material and biomechanics
aspects, and few studies exist which investigate the effect
of ventilation on air flow inside the helmet or effect of
this (or ventilation openings) on the dynamic performance
of helmet. It is possible that helmets with ventilation are
∗
Corresponding author. Email: [email protected]
ISSN: 1358-8265
C 2008 Taylor & Francis
Copyright DOI: 10.1080/13588260801933626
http://www.informaworld.com
available in the market, but systematic studies on these
are not available in the literature. In first part of the paper, we investigate air velocities inside the helmet, with
and without ventilation using computational fluid dynamics (CFD) techniques. The fluid flow study was carried out
to examine the possibility of improving the ventilation in
motorcycle helmets. In second part of the paper, the biomechanics characteristics of head impact were studied for helmets with and without ventilation using finite element (FE)
analysis.
In a motorcycle helmet, the comfort foam apparently
helps in fitting the helmet on heads of different sizes, although it is rarely used in bicycle helmets. Because the
comfort foam always rests on the head and gives resistance
to air flow, it has not been included in the helmet designs
studied here. We found that, without comfort foam and if
tied properly, the helmet sits on the head and does not move.
Even without comfort foam, however, there are regions in
the central plane of the head, where this helmet rests on
the head, and there is no space for air to circulate. It was
decided to support the helmet on the head by comfort foam
of 2 mm thickness provided on the sides of the helmet only
(but not on the top). A groove was made in the central plane
to provide space for air to flow, and a slot was provided in
front of the helmet for air to enter. The groove and slot
in the helmet are shown in Figure 1. Flow velocity inside
the helmet was determined by varying the depth and width
of the groove keeping slot dimensions fixed. Four different
sizes of groove, listed in Table 1, were investigated. The dimensions of the slot present in the outer shell and liner foam
were fixed at 48 mm × 7 mm. These helmets did not have
266
P.K. Pinnoji et al.
Figure 1. Three parallel grooves of 14 mm × 7 mm with slot in liner foam.
a visor. Studies were also performed on helmets without a
slot. However, velocities inside the helmet were generally
lower than those observed with a slot, and therefore, only
the latter case is discussed here.
2. Fluid flow analysis
Reischl [20] carried out an investigation of helmet ventilation designs for firefighter helmets and found that a helmet
with side ventilation holes was cooler than unventilated
helmet, and increasing the gap between helmet and head
also enhanced the cooling action due to the improved air
circulation. Abeysekera and Shanavaz [2] investigated the
potential benefits of helmets with ventilation holes for industrial workers both in laboratory and field settings. Bruhwiler et al. [4] studied the heat transfer variations of bicycle
helmets. They carried out experiments in water channel
to study the efficiency and placement of vents in bicycle
helmets. They concluded that there is significant potential
Table 1. Various ventilation designs.
Design
1
2
3
4
5
No. of
grooves
Groove width
(mm)
Groove depth
(mm)
Slot
(mm)
1
1
1
1
3
14
28
42
14
14
7
7
7
14
7
48 × 7
48 × 7
48 × 7
48 × 7
48 × 7
within the basic helmet structure for improving the heat
transfer.
The above studies except that of Bruhwiler et al. [4]
were mainly in conditions where the velocities of air were
very small. All the above studies were experimental, and
no CFD analysis had been done for fluid flow inside the
helmet. Here we study air flow inside a motorcycle helmet
using CFD where the relative velocity of air outside the
helmet is about 15 m s−1 . Inside the helmet the air gap
varies from approximately 0 mm (where the helmet touches
the head) to 10 mm, and the velocities are small. One of
the ways of improving these velocities and resulting heat
transfer is by providing grooves in the central plane of the
helmet. This leads to increased evaporation of the sweat.
The present study used CFD to investigate air flow inside
the gap between the helmet and head. To get an idea of
how well CFD is able to simulate the air flow in a gap of
10 mm, experiments were performed in a wind tunnel on
two concentric cylinders with a gap of 10 mm between the
cylinders and then simulated with CFD. Such experiments
with a helmet can be performed, but measurements in small
gap regions would be difficult with the Pitot tube used here.
2.1.
Experiments in wind tunnel and validation
of CFD
Experiments were performed on concentric cylinders in
the wind tunnel of cross section 450 mm × 750 mm. The
cylindrical model was mounted with axis of the cylinders
International Journal of Crashworthiness
267
Figure 2. Schematic view of two concentric cylinders model.
perpendicular to the direction of flow as seen in Figure
2. The diameter of the inner cylinder is approximately
the same as that of the actual head, i.e. 140 mm, and the
diameter of the outer cylinder was 160 mm. Six pressure
tabs were placed on the surface of the inner cylinder, and
velocity within the gap at various points along the y-axis
was measured using a three-hole velocity probe connected
to the micro-manometer. Velocity of the wind at inlet was
measured by a Pitot-static tube. Experiments were carried
out at three different speeds of the wind, 11 m s−1 , 15.7
m s−1 and 23.5 m s−1 although only results for 15.7 m s−1
are shown here.
Fluid flow analysis for the above concentric cylindrical
model was carried out through CFD using FLUENT [6].
The upstream length was four times the diameter of the
inner cylinder, and downstream length was 20 times the
diameter of inner cylinder. The downstream length was
kept large to ensure that the flow was fully developed. Fluid
inside the gap was air with density of 1.225 kg m−3 and
viscosity of 1.789 × 10−5 kg (m s)−1 .
For numerical simulations the inlet velocity was taken
as 15.7 m s−1 , and outflow condition available in FLUENT
was used at the outlet. No slip condition was assumed at
the walls. Fluid flow was assumed as steady and incompressible. The standard k − ε model [7] was used as the
turbulence model with standard wall functions. Here ‘k’ is
the kinetic energy of the particle and ‘ε’ is its dissipation
rate. Segregated solver, which solves the non-linear equation set sequentially, was used. Convergence criteria were
defined by specifying a tolerance on all FLUENT residuals
such as velocities, turbulence kinetic energy and dissipation
rate, which appear while solving the transport equations.
The air velocities along a vertical section on the central
plane of the cylinder (i.e. from top to bottom of the gap
at the central plane) in CFD simulations and wind tunnel
experiments are compared and depicted in Figure 3. The air
velocities in this gap are relatively lower on top and bottom
of the gap compared to the centre. It can be observed from
this figure that the experimental and numerical results show
a reasonable match, thus validating the CFD results.
3.
CFD analysis for air flow inside the helmets
with ventilation
FLUENT was next used to study the flow between the head
and various ventilated helmets. Geometry of the head was
Figure 3. Comparison of air velocities with computational fluid
dynamics and wind tunnel experiments.
268
P.K. Pinnoji et al.
the same as the one used later in impact dynamics analysis.
The helmet geometry was taken from a commercially available helmet and had a liner foam thickness of 32 mm. The
computational domain was 4750-mm length and 600-mm
high. The upstream and the downstream lengths were 750
and 4000 mm, respectively. It was assumed that helmet and
head are symmetrical with respect to the central plane. On
account of symmetry one half of the domain was modelled.
Tetrahedral grid was used for meshing and the minimum
grid size was 1.2 mm. The total number of elements in
different models was approximately 750,000.
Boundary conditions, turbulent models and convergence criteria, here, are the same as used in the Section
2. The helmet without ventilation had regions touching the
head where the air velocity is zero as air could not circulate.
Figure 4 shows the velocity contours in three dimensions
within the helmet–head gap with 14 mm × 14 mm groove
and 48 mm × 7 mm slot. Here air velocities in the central
plane are improved to 8 m s−1 near the slot, 6 m s−1 on top
and 5 m s−1 at the end of the groove. In a helmet with a
groove of 42 mm × 7 mm in the central plane and a slot of
48 mm × 7 mm, air velocities in the central plane which
can be seen in Figure 5 are further improved to 11 m s−1
near the slot, 5 m s−1 on top and 5 m s−1 at the end of
the groove. With these ventilation models, air velocities are
enhanced not only in the central plane of the groove but
also in other regions such as C and D in Figures 4 and 5,
respectively. The air velocities in other regions are higher
in the 42 mm × 7 mm groove helmet than in the 14 mm ×
14 mm groove helmet. A comparison with Figure 4 shows
Figure 4. Velocity contour values within the foam-head gap with
half symmetry in a helmet of 14 mm × 14 mm groove and a slot
of 48 mm × 7 mm.
Figure 5. Velocity contour values within the foam-head gap with
half symmetry in a helmet of 42 mm × 7 mm groove and a slot
of 48 mm × 7 mm.
that a larger groove in the helmet leads to higher air velocities inside the helmet–head gap. In a helmet with three
14 mm × 7 mm parallel grooves with spacing of 35 mm
between them and a slot of 48 mm × 7 mm in the front, air
velocity is higher only in the central plane (i.e. the groove
region) and almost zero adjacent to the central plane (region G in Figure 6) because eddies are formed in this region.
It, therefore, seems that for ventilation purposes providing
three grooves is not very advantageous.
Figure 6. Velocity contour values within the foam–head gap with
half symmetry in a helmet of three 14 mm × 7 mm parallel grooves
with spacing of 35 mm and a slot of 48 mm × 7 mm.
International Journal of Crashworthiness
269
Figure 7. Streamlines on the central x–z plane in and outside the gap with 14 mm × 14 mm groove and 48 mm × 7 mm slot (slot at 30◦
to the horizontal).
In ventilated helmets, the orientation of the slot is an
important parameter that affects the flow inside the helmet.
The slot is provided to increase the mass flow of air in the
helmet.
Two slot orientations, at angle 30◦ to the horizontal
and tangential to the head surface, were studied. Higher air
velocities were observed when the slot is tangential to the
head. For better fluid flow visualisation, the streamlines,
which show the direction of fluid flow, on the central plane
for the 30◦ slot and tangential slot are outlined in Figures
7 and 8, respectively. For a slot at 30◦ to the horizontal, a
vortex zone near the slot (region E in Figure 7) is formed
Figure 8. Streamlines on the central x–z plane in and outside the gap with 42 mm × 7 mm groove and 48 mm × 7 mm slot (slot at
tangential to the head surface).
270
P.K. Pinnoji et al.
Figure 9. Comparison of air velocities in various helmet ventilation models.
because of obstruction to the flow, but when the slot is
tangential to the head surface, the vortex zone is not formed
and the flow is smooth as seen in region F in Figure 8.
Air velocities along the head surface on the central
plane in different ventilated helmet models are compared
and shown in Figure 9. Angle θ , in Figure 9, is 0◦ at the
entrance to the helmet, 90◦ at the top and 210◦ at the rear
end. It can be observed from Figure 9 that there is lot
of fluctuation in air velocities at the entrance. There is a
decrease in velocity up to 20◦ followed by an increase up
to 40◦ , along the head surface. This decrease in velocity of
air in the beginning is due to the interaction of air which
is entering through the slot and from below through the
groove. The direction of air entering from both the sides is
different, and hence the resultant magnitude of air is less.
The ventilated helmet model with 42 mm × 7 mm groove
and slot has a maximum air velocity of 12 m s−1 and a
minimum of 3.8 m s−1 . With 14 mm × 14 mm groove
and slot in the helmet, air has maximum velocity of 10.4
m s−1 and minimum of 2.0 m s−1 . Air flow is improved
and similar trends are observed for all the ventilated helmet
models. In the conventional helmet without ventilation, air
velocities decreased in the beginning of helmet–head gap
and are almost constant at the top at 1 m s−1 .
4.
Dynamics of motorcycle helmets with deformable
head model
In the second part of the work, dynamics of helmet impact
are studied by focusing on the biomechanic characteristics
of the head. In the past, the FE analysis of drop test of
helmet used a rigid head, and results were reported in the
form of head injury criterion (HIC) values and accelerations of the headform. Lately, there has been an increased
interest in the FE modelling of the human head to study
the various injuries [11,15,21,23]. Ruan et al. [21] studied
the dynamic response of the human head impact with 50
percentile of the actual human head model. Kleiven and
Hardy [15] studied the localised brain motion, intracerebral acceleration, intracranial pressures and HIC. Horgan
and Gilchrist [11] constructed an FE model of the human
head and predicted the brain motion and intracranial pressure changes in head impact in pedestrian accidents. Zong
et al. [25] used a structural intensity approach to study the
power flow distribution inside the head in frontal, rear and
side impacts. The results using human head models are presented in the form of pressures, stresses and strains in the
brain although a clear relation between stresses and brain
injury is still to be fully established.
Gilchrist and Mills [8] used a one-dimensional (1-D)
analytical model to examine the dynamic response of helmets, and Brands et al. [3] carried out numerical simulation
for predicting the head injury using a 3-D FE model of a
motorcycle helmet. Yettram et al. [24] performed an FE
parametric study of the impact response of the motorcycle
helmets and used HIC for judging the crashworthiness performance. Pinnoji and Mahajan [19] investigated the ventilation in motorcycle helmets by considering the helmet as
a hemisphere and the head as a cylinder. They showed that
the presence of grooves improved the air velocities in the
gap between helmet and head without having a detrimental
effect on the dynamic performance of the helmet.
Deck et al. [5] simulated a frontal impact for helmet optimisation against the biomechanical criteria using FE modelling of helmet–head. Their parametric study was based on
the dynamic behaviour of helmet–head components, and
the values of brain pressures exceeded the tolerance limits
proposed in the literature. Otte et al. [18] found that many
of the injuries in motorcycle crashes are a result of rotation
of the head in oblique impacts. Aare et al. [1] investigated
the oblique impacts with FE model of Hybrid III dummy
head and FE models of human head and helmet. Kleiven
[14] studied the effect of load directions and durations with
FE model of the human head and evaluated Head Impact
Power (HIP), HIC, peak accelerations and maximum principal strain in the brain. Halldin et al. [9] developed a new
oblique test method for motorcycle helmets. They measured the linear and rotational accelerations by impacting
the helmet at 28◦ to the horizontal axis and found helmet
deformation to be larger than the oblique impacts in British
Standard (BS) 6658 and Economic Commission for Europe
(ECE) Regulation 22/05.
Safety standards for helmets prescribe a drop test.
All three impact situations – front, side and oblique –
were considered here. The drop test with a flat anvil was
simulated in LS-DYNA [10] using the FE models of helmet
and head. LS-DYNA is an explicit FE code for non-linear
and dynamic analysis. The FE analysis of helmet–head
impact required input consisting of geometry, initial and
International Journal of Crashworthiness
Table 2. Material properties of head [23].
Part
Skin
Skull (outer table)
Skull (inner table)
Cerebrospinal
fluid
Face
Falx
Tentorium
271
Table 3. Material properties of brain [23].
Density
(kg m−3 )
Elastic
modulus
(N m−2 )
1200
1800
1500
1040
16.7 × 106
15.0 × 109
4.5 × 109
12.0 × 103
0.42
0.21
0.0
0.49
7
2
3
3000
1140
1140
5.0 × 10
31.5 × 106
31.5 × 106
0.21
0.23
0.23
5
2
1
Poisson’s Thickness
ratio
(mm)
9
Part
Brain
1040
Bulk
modulus
(N m−2 )
G0 (N m−2 ) G∞ (N m−2 ) β (s−1 )
1.125 × 109 49.0 × 103
16.7 × 103
4.2. Constitutive model for helmet
The motorcycle helmet is made of a shell, liner foam, comfort foam and strap. Outer shells are made either from
a moulded thermoplastic like acrylo-butadiene styrene,
polycarbonate or from a composite material with glass or
carbon or Kevlar fibres. The outer shell resists the penetration of any foreign object and distributes the localised
forces to a wider area causing the large volume of the liner
foam to deform thus increasing its energy-absorbing capacity. A full-face helmet without visor is used in helmet–head
impact analysis as it covers a larger part of the head. A nylon strap is attached to the outer shell and is used to tie the
helmet to the head. The strap was assumed as linear elastic,
and the properties are defined in Table 4. In the dynamics part, shell thickness and material, liner foam thickness
and density are kept as constant but groove sizes are varied to study their effect on the helmet impact performance.
Comfort foam is not modelled in this analysis. Material
model 3 (*MAT PLASTIC KINEMATIC) available in LSDYNA was used for the outer shell in FE analysis to model
the acrylo-butadiene styrene material, and the properties
defined are given in Table 4.
In motorcycle helmet impact, the liner foams are subjected to plastic deformations. The energy-absorbing liner
foam considered here is made of expanded polystyrene
(EPS), which is elasto-plastic in nature. The stress–
strain behaviour of EPS foam with 44 kg m−3 density
is taken from Yettram et al. [24]. Material model 63
(*MAT CRUSHABLE FOAM) available in LS-DYNA
was used to model the EPS foam. This material model is an
isotropic foam model and crushes one-dimensionally with
zero Poisson’s ratio. This model transforms the stresses
into the principal stress space where the yield function is
defined. If the principal stresses exceed the yield stress
they are scaled back to the yield surface and transformed
4.1. Constitutive model of head
The biomechanical response of the head, by considering
it as deformable, is studied in terms of forces, pressures
and stresses. A 3-D FE model of human head developed by
Willinger et al. [23], which was validated against Nahum’s
experiments [16], is used here. The FE model of head has
skin, face, skull, cerebrospinal fluid (CSF), falx, tentorium
and brain.
The various layers of the head are generally nonhomogeneous, anisotropic, non-linear and viscoelastic.
However, for modelling purposes here, they are assumed
as homogeneous, isotropic and linearly elastic, except for
the brain, which is assumed as viscoelastic in nature. The
properties of the various head parts are listed in Table 2.
The shear characteristics of viscoelastic behaviour of the
brain are expressed by
(1)
Here G∞ is the long-term shear modulus, G0 is the shortterm shear modulus and β is the decay factor. The values of
G0 , G∞ and β suggested by Willinger et al. [23] are used
and listed in Table 3. Skull was modelled by a three-layer
sandwich material, which represents the outer and inner
table along with a soft dipole layer. Shell elements are used
to model the face, skull, falx and tentorium; solid elements
are used to model the skin, brain and CSF. Mass of the FE
Table 4. Material properties of outer shell and strap in helmet.
Part
Outer shell (acrylo-butadiene styrene)
Strap (nylon)
145
model of head is 4.5 kg, and there are 11,939 nodes and
13,193 elements in the FE mesh.
boundary conditions, interface conditions and material
properties. The output was in the form of stresses for the
deformable head. To evaluate the HIC, however, a rigid
head model was used.
G(t) = G∞ + (G0 − G∞ ) e−βt
Density
(kg m−3 )
Density
(kg m−3 )
Elastic modulus
(N m−2 )
Yield stress
(N m−2 )
Poisson’s
ratio
Thickness
(mm)
1200
1100
2.0 × 109
3.0 × 109
34.3 × 106
–
0.37
0.42
3
1
272
P.K. Pinnoji et al.
back to the original stress space. The yield surface and
its evolution are defined by the Equations (2) and (3),
respectively, given below.
ft = |σi | − Y = 0
(2)
Y = Y +H (ev )
0
Yt = Yt0
(3)
Here Y is the yield stress, Y ◦ is initial compressive yield
stress, Yt is tensile cut off stress, σ i is the principal stresses
and H is strain hardening, which is a function of the volumetric strain, ev , defined by natural logarithm of relative
volume. An associative flow rule for a flow surface, which
is same as yield surface, is assumed and the plastic strains
are derived from Equation (4)
·
εij·p = λ
∂F
∂ σij
(4)
·
εij·p is the plastic flow rate tensor, and λ is the plasticity
consistency parameter.
In LS-DYNA, the stress versus volumetric strain data
for the liner foam is given in tabular form and it fits the
material model 63 to this curve. The EPS foam has Young’s
modulus of 1.8 × 107 N m−2 with zero Poisson’s ratio and
a compressive yield stress (Y ◦ ) of 0.6 × 106 N m−2 .
4.3.
Finite element mesh and injury criteria
To carry out the complete impact analysis of helmet–head
FE simulations were performed for front, side and oblique
impacts against a flat anvil. Figure 10 shows the FE model of
helmet–head used in front, side and oblique impacts. Fournoded Belytschko-Tsay shell elements were used to model
the outer shell of motorcycle helmet with 3 mm thickness.
Belytschko-Tsay shell element has five integration points
through the thickness and is computationally efficient. The
EPS liner foam and the Nylon strap were modelled with
eight-noded brick elements. The model had 2130 elements
for outer shell, 7360 elements for the liner foam and 858
elements for the strap. The mass of the helmet was 0.8 kg.
Surface-to-surface contact based on penalty formulation
with low coefficient of friction was modelled between the
head–helmet and between the helmet–flat anvil to prevent
interpenetration of these surfaces. Numerical simulations
were performed with initial velocity of the helmet–head
system varying from 7 to 10 m s−1 .
Head injury criterion characterises the injury of the head
under the impact by not only involving the peak acceleration
but also the distribution and duration of the acceleration
over the time of impact [17]. For finding HIC, the head
is considered as rigid. For numerical simulations in LSDYNA, the geometry of the head is the same as used in the
deformable head model. The HIC is calculated as


HIC = 
1
(t2 − t1 )
t2
2.5

a(t)dt 
∗ (t2 − t1 )
(5)
t1
where a is the resultant acceleration at the centre of gravity
of the rigid head in units of acceleration of gravity (g =
9.81 m s−2 ). t1 and t2 are the time points in seconds during
the crash for which HIC is maximum.
For a deformable head to predict the damage in brain
during head impact, von Mises stress has been used by
Kang et al. [13] and Shuaeib et al. [22]. Lower the von
Mises stress in the brain under helmet impact, the better
helmet it is. Intracranial pressure, which is the pressure
exerted by skull on the brain tissue and CSF, is used by Ruan
et al. [21], Zong et al. [25], Klieven [14] and Kang et al.
[13] for validation of human head models. One of the most
damaging aspects of the increase in intracranial pressure
is brain trauma. Normally, the intracranial pressures are
increased due to the brain swelling and the blockage of
CSF outflow in the brain ventricles at the base of brain.
In the head impact, skull may undergo deformation, and
change in volume of the skull is compensated by change
in volume of the CSF and brain. The skull deformation
compresses the CSF, which accumulates in the brain and
causes an increase in the intracranial pressure.
5. Results and discussion
Helmet standards use flat anvils as impact surface for determining the performance of helmet. For Indian road conditions, the IS 4151standard [12] prescribes four impact
points for impact tests with one at front, one on the rear and
Figure 10. Finite element model of helmet–head in front, side and oblique impacts.
International Journal of Crashworthiness
273
Figure 11. Force on the head with and without helmet (front
impact).
Figure 12. Intracranial pressures at coup and contra-coup with
and without helmet (front impact).
two on the either sides of the helmet and is compatible with
ECE R22:03 and SNELL motorcycle helmet standards. It
recommends that the peak acceleration of the head should
not exceed 300 g for an impact velocity of 7 m s−1 . First,
the numerical results were compared for front impact with
a bare head and with the helmet over it.
pressures at coup and contra-coup sites when head impacts
the flat rigid anvil without and with helmet. Without helmet
the intracranial pressures are 1.4 and 0.78 MPa at coup
and contra-coup, respectively. With helmet on head, the
intracranial pressures are reduced to 0.21 MPa at coup and
0.12 MPa at contra-coup site and closely match with the
values given by Deck et al. [5].
When the helmet–head impacted the flat rigid surface,
initially the polystyrene foam deformed from outside as the
outer shell came in contact with the rigid surface and came
to rest whereas liner foam continued to deform. Later, at
1.5 ms the head crushes the polystyrene foam permanently
from inside and the compression is highly local. At 10 m s−1
impact velocity, the liner foam in the helmet without ventilation has almost bottomed-out and the stresses in the foam
are high. Figure 13 shows the deformation in motorcycle
helmet under head impact at 10 m s−1 velocity. Force on
the head at different impact velocities during the frontal
impact is depicted in Figure 14, and it can be observed that
the sharpness of the force versus time curve rises with impact velocity. It means that the peak force on the head acts
5.1.
Front impact
Figure 11 shows the force versus time curve at 7 m s−1
impact velocity for head impact with and without a helmet.
In head impact without a helmet, the contact duration
between the head and rigid surface was approximately 2
ms. The contact force increased sharply to 47,000 N at
1.8 ms and dropped thereafter. For the head with a helmet,
the maximum force on head was found to be 7230 N at 6
ms, and the contact duration between the head and helmet
was approximately 7 ms. This force on the head matches
closely with the results of Deck et al. [5], who predicted a
force of 8000 N on head for front impact of helmet–head
at 7.5 m s−1 velocity. Figure 12 shows the intracranial
Figure 13. Deformation in helmet under head impact at 10 m s−1 velocity (front impact).
274
P.K. Pinnoji et al.
Figure 14. Force on the head without helmet ventilation at different velocities (front impact).
Figure 15. Force on the head with ventilated helmet at 7 m s−1
velocity (front impact).
for longer duration at low-velocity impacts and for smaller
duration at high-velocity impacts. The force on the head
creates positive pressures at the coup site in brain because
of compression, and negative pressures are developed at the
contra-coup site because of tension. With all impact velocities, the maximum force on head was reached between 5
and 6 ms time after the first contact between the helmet
and rigid surface. In front impact at 10 m s−1 velocity,
the helmet without ventilation predicted a maximum force
of 11,496 N on the head and is higher than the predicted
fracture force of skull 9900 N by Willinger et al. [23]. So,
this helmet would not protect the human head at 10 m s−1
impact velocity.
Figure 15 shows the force on the head with different
ventilation models in the helmet when it impacts a flat rigid
surface at 7 m s−1 velocity. It is observed that the trend in
the dynamic force with time is qualitatively same for all
the ventilated helmet models but the peak value differs.
Forces on the head were lower with some ventilated helmet
models compared with the non-ventilated helmet models.
In the region with grooves, the liner foam behaves locally
as low-density foam. Moreover, during the deformation
under the impact, the foam was deformed towards these
grooves and slot as there was some space to displace.
These probably account for reduction of forces on the head
with ventilated helmets. The minimum peak force on the
head was 6574 N for the helmet with a 14 mm × 7 mm
groove and 48 mm × 7 mm slot, and the maximum peak
force was 7364 N for the helmet with 42 mm × 7 mm
groove. All the helmet ventilation models except 42 mm
× 7 mm groove gave lower peak force than the helmet
without ventilation at 7 m s−1 impact velocity.
The biomechanical parameters such as force, intracranial pressure, von Mises stress, HIC and resultant acceleration of head with all ventilated helmet models for 7 m s−1
velocity in front impact and the maximum value for each
case are listed in Table 5. The first five parameters are calculated by considering the head as deformable, whereas
for last two, the head is considered as rigid. Higher forces
were predicted for a helmet model with a groove of size
Table 5. Biomechanical parameters with different helmet ventilation models in front impact at 7 m s−1 initial velocity.
Intracranial pressure (N m−2 )
Helmet type
No ventilation
14 mm× 7 mm groove
28 mm × 7 mm groove
42 mm × 7 mm groove
14 mm × 14 mm groove
14 mm × 7 mm –
3 grooves
Force on the
helmet (N)
Force on the
head (N)
Coup
7441
6846
7363
7739
7057
7137
7230
6574
7058
7364
6862
6850
2.1 × 105
1.87 × 105
2.1 × 105
2.36 × 105
1.94 × 105
1.95 × 105
Contra-coup
Von Mises
stress in
the brain (kPa)
Head injury
criterion
Peak
acceleration
(g)
−1.19 × 105
−1.1 × 105
−1.2 × 105
−1.28 × 105
−1.15 × 105
−1.14 × 105
47.4
45.7
50.9
55.6
46.1
46.9
867
774
868
1051
744
691
170
158
168
183
155
160
International Journal of Crashworthiness
275
Figure 16. Von Mises stress in brain with helmet of 14 mm × 14 mm groove and slot (front impact).
42 mm × 7 mm as compared with other ventilation models. At 7 m s−1 impact velocity, the variation of intracranial
pressures at coup and contra-coup sites for various helmet
models with ventilation was very similar to the one for helmet without grooves and slots (Figure 12), and the peak
values are almost the same with all the ventilation models
of helmet. Figure 16 shows the von Mises stress distribution
in the brain when helmet with 14 mm × 14 mm groove and
slot impacts the rigid surface at 7 m s−1 velocity. In front
impact, the maximum von Mises stress occurred in brain
stem in all models of helmet whether with or without ventilation. The maximum von Mises stresses are in the range of
45–55 kPa and are similar to the values observed by Deck
et al. [5] but are higher than the 20-kPa limit proposed for
the neurological injuries.
The HIC and peak acceleration are lowest for the helmet with 14 mm × 14 mm groove although force on the
head and von Mises stresses are lowest for the helmet with
14 mm × 7 mm groove. The difference in the values of
these quantities between the two helmets is quite small. The
resultant acceleration of head versus time for various ventilation models of helmets is shown in Figure 17. The magnitude of acceleration (170 g) for the non-ventilated helmet
is lower than that in Brands et al. [3] but closely matches
Figure 17. Resultant acceleration of the head with and without
helmet ventilation (front impact).
Figure 18. Force on the head without helmet ventilation at different velocities (side impact).
276
P.K. Pinnoji et al.
Figure 19. Force on the head with ventilated helmet at 7 m s−1
velocity (side impact).
with that in Deck et al. [5]. In all the helmet models with
grooves and slot, the liner foam was crushed totally for an
initial impact velocity of 9 m s−1 , and the forces on the head
increased, whereas in the helmet with no grooves and slot,
the liner foam was crushed at an impact velocity of 10 m s−1 .
5.2.
Side impact
The forces on the head, HIC and resultant acceleration of the
head are higher in side impact than in front impact for same
impact velocity, but the intracranial pressures and von Mises
stresses are almost same. Figure 18 shows the force on the
head under helmet impact on side at different velocities.
The trend of force versus time is almost similar to the front
impact although the magnitude is almost 33% higher. The
gap between head and liner foam is almost 10 mm on side
compared with 4 mm on front. So, the head took some time
to come into contact with liner foam, and it can be observed
from the figure that the forces rise after 3 ms. In side impact,
the liner foam was totally crushed at 10 m s−1 velocity.
Figure 20. Force on the head with various ventilation models
(oblique impact).
Figure 19 shows the force on the head with various
ventilation helmet models at 7 m s−1 impact velocity.
The variation in forces, intracranial pressures, von Mises
stresses, HIC and acceleration of the head between different
ventilation models is much smaller in side impact as compared with the front impact. The biomechanical parameters
for side impact with various helmet ventilation models are
given in Table 6.
5.3. Oblique impact
The study of biomechanics of helmet–head in oblique impact is also needed as it involves rotation and results in
large shear strains in the brain, which can cause traumatic
brain injuries like diffuse axonal injuries. Rotation occurs
in helmet–head impact because the impact point is not situated straight under the centre of gravity of the helmet–head.
For the oblique impact test, the mean angle between the
helmet direction and the horizontal axis considered in this
study was 45◦ . In FE analysis, the helmet–head impact was
simulated at 7 m s−1 impact velocity.
Table 6. Biomechanical parameters with various helmet ventilation models in side impact at 7 m s−1 initial velocity.
Intracranial pressure (N m−2 )
Helmet type
No ventilation
14 mm × 7 mm groove
28 mm × 7 mm groove
42 mm × 7 mm groove
14 mm × 14 mm groove
14 mm × 7 mm – 3 grooves
Force on the
helmet (N)
Force on the
head (N)
Coup
10,202
10,132
10,207
10,074
10,187
10,126
9513
9504
9552
9384
9531
9512
2.2 × 105
2.2 × 105
2.2 × 105
2.2 × 105
2.2 × 105
2.19 × 105
Contra-coup
Von Mises
stress in
the brain (kPa)
Head injury
criterion
Peak
acceleration
(g)
−1.43 × 105
−1.43 × 105
−1.43 × 105
−1.43 × 105
−1.43 × 105
−1.43 × 105
63.5
62.3
62.1
63.4
62.2
63.2
1680
1693
1701
1690
1730
1682
229
231
235
234
223
230
International Journal of Crashworthiness
277
Table 7. Biomechanical parameters with various ventilation models in oblique impact at 7 m s−1 initial velocity.
Intracranial pressure (N m−2 )
Helmet type
No ventilation
14 × 7 groove
28 × 7 groove
42 × 7 groove
14 × 14 groove
14 × 7 – 3 grooves
Force on the Force on the
helmet (N)
head (N)
7248
7663
7596
7483
7467
8043
6339
6657
6474
6506
6502
6691
Coup
Contra-coup
1.76 × 105
1.81 × 105
1.82 × 105
1.78 × 105
1.85 × 105
2.17 × 105
−1.57 × 105
−1.56 × 105
−1.56 × 105
−1.55 × 105
−1.57 × 105
−1.43 × 105
Figure 20 shows the force on the head with various ventilation models of helmet. It is observed that there is not
much variation in force on the head between various ventilation helmet models and non-ventilated helmet models.
During oblique impact the helmet seems to undergo both
compression and bending, and the forces on the helmet and
head are not as high as in front and side impacts. The compression in oblique impact is highly localised than in front
and side impacts. This could be due to the higher curvature
of helmet and head at 45◦ angle.
The maximum values for various biomechanical parameters in helmet–head oblique impact are listed in Table 7.
As in front impact the intracranial pressures are almost the
same in oblique impact with all the helmet models, and
a maximum compressive value of 0.21 MPa in the region
of the impact site and a maximum tension value of 0.15
MPa are predicted at the opposite point. The HIC and the
resultant acceleration of the head are higher compared with
front impact for helmet without ventilation. The maximum
principal strain (Almansi) occurred in the central region of
the brain and did not vary much with the type of ventilation
in helmet. The maximum strain, 0.36, in the brain tissue and
peak acceleration of the head, 198 g, at 7 m s−1 impact velocity for helmet without ventilation closely matches with
Aare et al. [1]. In oblique impact, the liner foam in the helmet model with three grooves of 14 mm × 7 mm and a slot
crushed totally near the impact site at 7 m s−1 velocity, and
the forces were found to be 8043 N on the helmet and 6691
N on the head. The HIC was not calculated for this helmet
model as the liner foam crushed totally after 6.5 ms, and the
simulation was stopped. The EPS foam in all the ventilation
models of this helmet was crushed totally at 9 m s−1 , and
the forces on the head increased to approximately 12,000
N, which is higher than the skull fracture force.
6.
Conclusions
The air flow inside a motorcycle helmet and impact dynamics of the helmet are studied. The CFD is used to determine
air velocities inside the gap between the head and helmet.
Air velocities inside the conventional helmet are very low.
Von Mises
Peak
in stress
Max. strain Head injury acceleration
the brain (kPa) in brain
criterion
(g)
33.9
33.5
33.3
33.6
33.5
35.4
0.361
0.367
0.366
0.366
0.366
0.363
1524
1195
1264
1314
1550
–
198
184
177
201
199
–
These velocities can be improved by providing a groove in
the central plane of the liner and a slot in front. All the
ventilated helmet models gave improved higher air velocities than the helmet without ventilation. This improvement
in air velocities is observed not only in the central plane of
the groove but also in other regions of the helmet–head gap.
For a helmet model with three parallel grooves and slot (of
fixed size used here), results show that air velocities vary at
each groove entrance and did not show increase in regions
other than the central plane. The slot orientation affects air
velocities, and the slot tangential to the head gives higher
values of air velocity in the helmet–head gap.
In the dynamics study, the various head biomechanical parameters are estimated for helmet with ventilation
through the FE analysis. The human head model is considered as deformable for finding the forces, intracranial
pressures, stresses and strains and as rigid for finding HIC
and peak acceleration. The analysis is carried out for front,
side and oblique impacts with all the ventilated helmet models along with head. The analysis predicts higher stresses
in the head than the tolerance limits given in the standards
although they match well with values reported by other researchers. In front impact at 7 m s−1 velocity, the von Mises
stresses for helmets with 14 mm × 7 mm and 14 mm ×
14 mm groove dimensions are lower as compared with the
conventional helmet without ventilation. In side and oblique
impacts, the von Mises stresses are almost same for all the
helmet models. All the models of helmet with ventilation
passed the standards in terms of HIC and peak acceleration
at 7 m s−1 impact velocity but failed at higher velocities.
During front impact, the liner foam in helmets with ventilation bottomed out at approximately 9 m s−1 as compared
to the 10 m s−1 observed in the conventional helmet. A
similar phenomenon was also observed in side and oblique
impacts.
Though the air velocities are higher in the gap between
head and helmet with 42 mm × 7 mm groove and forces on
the head are lower in side impact, the stresses in front impact exceed those of stresses in the conventional helmet and
helmet with 14 mm × 14 mm groove. For 42 mm × 14 mm
groove, the values in front impact are still higher and, therefore, this size of groove is not recommended. It is seen that
278
P.K. Pinnoji et al.
higher groove size gives higher air velocities inside the helmet but at the same time results in higher stresses in the
brain if the groove width and depth exceed a certain size.
The compromise design is to minimise the injury or harm to
the helmet wearers while providing ventilation. Among the
various helmet ventilation models, the one with a groove
of 14 mm × 14 mm along with a slot of 48 mm × 7 mm
is preferable at 7 m s−1 impact velocity as it enhances the
average air velocities in the helmet–head gap, and the head
experiences lower forces and stresses. As no literature is
available on the optimisation of ventilation in motorcycle
helmets, it is recommended that further studies should be
carried out to improve the ventilation as well as dynamics
performance of ventilated helmet impact.
Acknowledgment
The authors express gratitude to the Transportation Research and
Injury Prevention Programme, IIT-Delhi, India, for assistance and
financial support provided through a sponsored research project.
References
[1] M. Aare, S. Kleiven, and P. Halldin, Injury tolerances for
oblique impact helmet testing, Int. J. Crashworthiness 9
(2004), pp. 15–23.
[2] J.D.A. Abeysekera and H. Shanavaz, Ergonomic evaluation of modified industrial helmets for use in tropical environments, Ergonomics 31 (1988), pp. 1317–
1329.
[3] D.W.A. Brands, J.G.M. Thunnissen, and J.S.H.M. Wismans, Modelling head injury countermeasures: A 3D helmet model, AGARD Specialists’ Meeting on Impact Head
Injury: Responses, Mechanisms, Tolerance, Treatment and
Countermeasures, New Mexico, 7–9 November 1996.
[4] P.A. Bruhwiler, M. Buyan, R. Huber, C.P. Bogerd, J. Sznitman, S.F. Graf, and T. Rosgen, Heat transfer variations of
bicycle helmets, J. Sports Sci. 24 (2006), pp. 999–1011.
[5] C. Deck, D. Baumgartner, and R. Willinger, Helmet optimization based on head-helmet impact modeling, 8th International Conference on Computer Aided Optimum Design of Structures, OPTI VIII, MNP, Michigan, USA, 2003,
pp. 319–328.
[6] FLUENT 6.2 User’s Guide, FLUENT Inc., 2005.
[7] J.H. Ferziger and H. Peric, Computational Methods for Fluid
Dynamics, 3rd ed., Springer, Berlin, 2002.
[8] A. Gilchrist and N.J. Mills, Modelling of the impact performance of motorcycle helmets, Int. J. Impact Eng. 15 (1994),
pp. 201–218.
[9] P. Halldin, A. Gilchrist, and N.J. Mills, A new oblique impact
test for motorcycle helmets, IRCOBI Conference, Isle of
Man, UK, 2001, pp. 343–345.
[10] J.O. Hallquist, LS-DYNA Theoretical Manual, Livermore
Software Technology Corporation, 1998.
[11] T.J. Horgan and M.D. Gilchrist, Influence of FE model variability in predicting brain motion and intracranial pressure
changes in head impact simulations, Int. J. Crashworthiness
9 (2004), pp. 401–418.
[12] IS 4151, Indian standard specification for motorcycle helmet
users, Bureau of Indian Standards, New Delhi, 1993.
[13] H. Kang, R. Willinger, B.M. Diaw, and B. Chinn, Validation of a 3D anatomic human head model and replication of
head impact in motorcycle accident by finite element modeling, SAE Transactions, Paper No. 973339, 1997, pp. 329–
338.
[14] S. Kleiven, Evaluation of head injury criteria using an FE
model validated against experiments on localized brain motion, intra-cerebral acceleration, and intra-cranial pressure,
Int. J. Crashworthiness 11 (2006), pp. 65–79.
[15] S. Kleiven and W.N. Hardy, Correlation of an FE Model of
human head with experiments on localized motion of brain –
consequences for injury prediction, Society for Automotive
Engineers, 46th Stapp Car Crash Journal, 2002, SAE Paper
No. 02S-76.
[16] A.M. Nahum, R. Smith, and C.C. Ward, Intracranial pressure dynamics during head impact, Proceedings of 21st
Stapp Car Crash Conference, New Orleans, USA, 1977,
pp. 339–366.
[17] J.A. Newman, Head injury criteria in automotive crash
testing, Society of Automotive Engineers, Warrendale, PA,
1970, SAE 700426.
[18] D. Otte, B. Chinn, D. Doyle, S. Makitupa, K. Sturrock, and E.
Schuller, Contribution to Final Report of COST 327 Project,
University of Hanover, Germany, 1999.
[19] P.K. Pinnoji and P. Mahajan, Impact analysis of helmets for
improved ventilation with deformable head model, IRCOBI
Conference, Madrid, 2006, pp. 159–170.
[20] U. Reischl, Fire fighter helmet ventilation analysis, Am. Ind.
Hyg. Assoc. J. 47 (1986), pp. 546–551.
[21] J.S. Ruan, T.B. Khalil , and A.I. King, Dynamic response
of the human head impact by three-dimensional finite element analysis, J. Biomech. Eng. Trans. ASME 116 (1994),
pp. 44–50.
[22] F.M. Shuaeib, A.M.S. Hamouda, R.S. Radin Umar, M.M.
Hamdan, and M.S.J. Hashmi, Motorcycle helmet Part 1:
Biomechanics and computational issues, J. Mater. Process.
Technol. 123 (2002), pp. 406–421.
[23] R. Willinger, B.M. Diaw, and H.-S. Kang, Finite element
modeling of skull fractures caused by direct impact, Int. J.
Crashworthiness 5 (2000), pp. 249–258.
[24] A.L. Yettram, N.P.M. Godfrey, and B.P. Chinn, Materials
for motorcycle crash helmets – a finite element parametric
study, Plastics Rubber Compos. Process. Appl. 22 (1994),
pp. 215–221.
[25] Z. Zong, H. Lee, and C. Lu, A three-dimensional human
head finite element model and power flow in a human
head subjected to impact loading, J. Biomech. 39 (2006),
pp. 284–292.