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HUMAN FACIAL SOFT TISSUE THICKNESS AND MECHANICAL PROPERTIES-A
Human-Centric Design Research Lab
Department of Mechanical Engineering
Texas Tech University
Lubbock, TX 79409, USA
Human-Centric Design Research Lab
Department of Mechanical Engineering
Texas Tech University
Lubbock, TX 79409, USA
The finite element method (FEM) has been used in human
facial modeling both in clinical and engineering fields for
decades. Applications of human head modeling include the
interaction of personal protective equipment with the human
head and modeling head impact. In human head modeling, it is
critical to have a high fidelity model including accurate
thicknesses of each layer and accurate material properties.
Various experiments have been performed but do not report
consistent results; therefore, it is difficult to find reliable
parameter values to create an effective model of the human
head. This paper attempts to review and summarize the state of
the art of human facial studies including experimental
measurements of different layer thicknesses and the mechanical
properties of these layers.
Keywords: Human head modeling; facial soft tissue; material
properties; finite element method.
The human head houses several of the most important
components of the body, including the brain, mouth, nose, and
eyes. Significant research has been dedicated to the creation of
personal protective equipment (PPE) such as helmets and
respirators which protect against brain trauma and restrict
contaminants from entering the respiratory system.
Additionally, a significant portion of vehicle design is aimed at
preventing injury to the head during a collision. Traditional
evaluation of PPE tests physical prototypes using experiments.
However, with high speed computational tools and advanced
software, it is much easier to assess new protective equipment
designs at an early stage without physical prototypes. To create
effective simulations, it is critical to have high fidelity human
head models which include accurate segment geometries and
mechanical properties. In the literature, various thicknesses of
human facial skin and soft tissues and linear and nonlinear
mechanical properties have been reported. It is difficult to
determine which parameter values are acceptable for an
interaction simulation between the head and PPE. The objective
of this paper is to provide a summary of the literature reporting
mechanical properties and thicknesses of human facial skin and
soft tissues which will be useful for the creation of human head
models. In literature, finite element facial models simulating
facial soft tissues are mainly divided into two groups: linear
elastic models and nonlinear elastic models. Linear elastic
models treat the facial tissue as a homogeneous, linear elastic
material. For a linear elastic model, Hooke’s law is applied to
describe the material properties using two parameters: Young’s
modulus and Poisson’s ratio. The stress-strain response is
assumed to be linear under the condition that the tissues have
small deformations. Several researchers (Keeve et al., 1998;
Koch et al., 1999; Zachow et al., 2000) have used this method
to model the deformation of the facial tissue. These simplified
linear models produce solutions much more quickly; however,
some researchers have reported that linear facial models
become inaccurate when the deformations are large (Picinbono
et al., 2000, Gladilin et al., 2003). Fung (1993) showed in an
experiment that stress increases much faster than linearly after
relative strain becomes larger than 10% or 15%. The skin has
been proved to be a non-linear, anistropic and viscoelastic
material, which is too complex for a linear elastic model to
describe (Daly, 1982; Silver et al., 2001). Gerard et al. (2005)
created FE hyperelastic models for the cheek and tongue. Other
researchers created nonlinear isotropic models to describe the
mechanical behavior of the human skin. These include Ogden
(Lapeer et al., 2010; Shergold et al., 2006; Evans and Holt,
Proceedings of the ASME 2015 International Design Engineering Technical Conferences &
Computers and Information in Engineering Conference
August 2-5, 2015, Boston, Massachusetts, USA
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2009; Flyn et al., 2011), Mooney-Rivlin (Hendriks, 2003) and
neo-Hookean (Delalleau et al., 2008) material models. For the
muscle, hyperplastic (Cox et al., 2007; Jia et al., 2006; Linder-
Ganz and Gefen, 2004) and viscoelastic (Aritan et al., 2008;
Best et al., 1994; Bosboom et al., 2001; Van Loocke et al.,
2008) models have been introduced in literature.
HUMAN HEAD ANATOMY
The facial soft tissues are those from the skin surface to the
most superficial surface of the underlying bone (Stephan and
Simpson, 2008). The facial soft tissues consist of three layers:
skin, subcutaneous tissue and muscle (Gladilin, 2003), where
the subcutaneous layer is mainly composed of fatty tissue. The
structure of these layers is shown in Figure 1. The thickness of
facial soft tissues varies for different sites, and in the literature,
most researchers have tested the thickness of facial soft tissues
at several different sites (Pelican and Seidenari, 1999). Figure 2
shows the names of different sites on human face. The major
facial muscles include the masseter, levator labii superioris, and
zygomaticus major muscles shown in Figure 3.
Figure 1. Facial soft tissue structure
Figure 2. Facial sites
Facial skin thickness
Accurate measurements for the thickness of human facial skin
are valuable in both clinical and scientific fields. The thickness
of human facial skin varies with age, race, and exposure to
sunlight (Warren et al., 1991). The measured thickness of
human facial skin depends on the method of measurement. In
the literature, methods of measuring skin thickness can be
divided into two groups: in vivo which is on living people
(Takema et al., 1994) and in vitro which is on cadavers (Ha et
al., 2005). Generally, in vivo studies are considered more
accurate than in vitro studies. However, there is no significant
difference between the results of these two kinds of studies (Ha
et al., 2005; Takema et al., 1994). Most researchers have used
in vivo methods to measure skin thickness because of the lower
cost and higher efficiency compared to the in vitro method. In
vivo, pulsed ultrasound, micrometer screw gauges (Newton et
al., 1984), skinfold caliper (Dykes et al., 1976), and
radiography methods (Bilznak and Staple, 1975; Cho et al.,
2011; Newton et al., 1984) have been used to provide absolute
values for facial skin thickness. It is well believed that the
radiography method might pose a biomedical hazard to the
human body. The skinfold caliper is not popular among
researchers because it can’t provide sufficient results compared
to other measurement methods. Most of the researchers used
ultrasound to measure the skin thickness (Denda and
Takahashi, 1990; Diridollou et al., 2001; Gibney et al., 2010;
Laurent et al., 2007; Mak et al., 2014; Ploin et al., 2011; Rigal
et al., 1989; Seidenari et al., 1994). There are two forms of
ultrasound measuring methods: A mode and B mode. A-mode
ultrasound devices are believed to be easier to operate than B-
mode ultrasound devices (Seidenari et al., 1994). Tan et al.
(1982) reported that the skin thicknesses obtained from in vitro
measurement are larger than those obtained from in vivo
measurement because of the loss of resting dermal tension and
potential distortion of the sample during biopsy. Denda and
Takahasi (1990) reported that the facial skin thickness on the
forehead and cheek decreases with age. However, Takema et al.
(1994) reported that the facial skin thickness has a linear
increase with age on the forehead and cheek while in the corner
of the mouth the skin thickness remains constant. Both groups
used and A-mode scanner (Dermascan A, Cortex Technology,
Denmark), and their conclusions are contradictory. A potential
reason is that the A-mode scanner cannot determine the
interface between the dermis and subcutaneous tissue.
Pellacani and Seidenari (1999) used a 20 MHz B-mode scanner
(Dermascan C, Cortex Technology, Denmark) to obtain the in
vivo skin thicknesses of 40 healthy Caucasian women for the
cheeks, upper lips, chin and nose as shown in Table 1. The B-
mode scanner can provide cross-sectional imaging of the skin
and therefore produces a more reliable measurement of skin
thickness where the dermis-hypodermis interface is hard to
Ha et al. (2004) conducted an in vitro study on 3 fresh adult
cadavers and examined the facial skin thickness on the similar
facial sites tested by Pelican and Seidenari (1999). The result
showed that the skin on the eyelid is thinnest, and results are
very close to those obtained by Pelican and Seidenari (1999).
Copyright © 2015 by ASME
Figure 3. Major facial muscles (red parts): (a) masseter
muscles; (b) levator labii superioris; (c) zygomaticus major
Table 1. Mean and standard deviation (SD) of skin thickness at
different facial sites in young and elderly subjects (Pelican and
Facial Sites Mean (SD) for Mean (SD) for
Young Subjects (mm) Elderly Subjects (mm)
Central Forehead 1.95 (0.22) 2.03 (0.37)
Cheeks 1.71 (0.35) 1.86 (0.25)
Upper Lip 2.62 (0.35) 3.05 (0.45)
Lower Lip 2.15 (0.40) 2.05 (0.42)
Chin 2.55 (0.55) 2.61 (0.45)
Nose 1.95 (0.32) 2.19 (0.34)
Devices measuring facial skin thickness
Diridollou et al. (2001) tested human skin thickness using a B-
mode real-time ultrasonic scanner (INSERM, Unit 316, G.I.P
Ultrasound Group, Ultrasound Technology Company, Tours,
France) and an imager (DermCup 2020, Atys Medical, Soucieu
en Jarrest, France). Gibney et al. (2010) used high frequency
Cortex Dermascan C ultrasound unit with a 20 MHz probe and
a GE LOGIQ e ultrasound unit (General Electric Healthcare,
Little Chalfont, Buckinghamshire, United Kingdom) and a 4.5-
13.0 MHz probe. Laurent et al. (2007) and Ploin et al. (2011)
used a high frequency Cortex Dermascan C ultrasound unit.
Mak et al. (2014) used an ultrasound biomicroscopic system,
Vevo 770 (Visualsonics Inc., Toronto, Canada) with a
transducer of 55 MHz.
Facial skin mechanical properties
In literature, there are few studies providing direct experimental
data for human facial skin mechanical properties. The skin
mechanical properties are different when measured using in
vivo and in vitro methods (Fung, 1993; Wilkes et al., 1973).
Barbarino et al. (2009) gained results for human facial skin
mechanical properties from an in vitro experiment. Diridollou
et al. (1998) claimed that in vitro tests for skin properties are
limited because the characteristics of the skin are different after
isolating skin from the human body. Different non-invasive
techniques have been developed to measure the skin
mechanical properties using in vivo methods. The in vivo
methods for testing mechanical properties can be generally
divided into two categories: imaging techniques and
mechanical testing methods (Hendriks, 2001). Imaging
techniques have mainly been used to observe deformation and
calculate the mechanical properties of skin. Imaging techniques
in literature include motion analysis (Mahmud et al., 2010),
digital image correlation (Evans and Holt, 2009; Staloff et al.,
2007), optical coherence tomography (OCT) (Hendriks et al.,
2006; Li et al., 2012; Liang and Boppart, 2010), magnetic
resonance imaging (MRI) (Tran et al., 2008) and ultrasound
(Diridollou, 2001; Diridollou et al., 1998; Gahagnon et al.,
2009; Gennisson et al., 2004; Moran et al., 1995; Pan et al.,
1998; Piotrzkowska et al., 2009; Zheng and Mak, 1996).
Commonly used in vivo mechanical methods were developed
in terms of measurement of suction (Barel et al., 1998; Bunegin
and Moore, 2006; Couturaud et al., 1995; Cua et al., 1990;
Diridollou et al., 2000; Grahame and Holt, 1969; Hendriks et
al., 2003, 2004; Khatyr et al., 2006; Malm et al., 1995), torsion
(Leveque et al., 1980; Escoffier et al., 1989; Sanders, 1973),
tension (Lapeer et al., 2010), indentation (Bader and Bowker,
1993; Delalleau et al., 2006; Pailler-Mattei and Zahouani,
2004; Pailler-Mattei et al., 2008; Tran et al., 2008; Zheng and
Mak, 1996), extension (Alexander and Cook, 1977; Khatyr et
al., 2004; Lim et al., 2008; Quan et al., 1997) and traction
(Wang and Hayward, 2007). Agache et al. (1980) obtained
values for Young’s modulus ranging from 0.42 MPa to 0.85
MPa with torsion tests. Manschot and Brakkee (1986) used
tensile tests and found values for Young’s modulus of the
human skin between 4.6 kPa and 20 kpa. For the suction tests,
Diridollou et al. (2000) measured Young’s modulus between
217 kPa and 341 kPa. Hendriks et al. (2003) found 10, skin
be 9.4 ± 3.6 kPa and 11, skin
C to be 82 ± 60 kPa for a finite
element model exhibiting Mooney material behavior to account
for the non-linear stress-strain relationship. It is obvious that
results from different researchers are different. The main reason
might be that they modified the human skin’s natural state of
stress because the experimental devices had to be fixed to the
skin throughout the testing period. It is very difficult to control
the prestress introduced by the mechanical device. Therefore,
the measured values of mechanical properties might be
affected. Although not commonly used, the indentation method
can be used to measure the skin mechanical properties without
prestressing the skin specimen before the test (Bader and
Bowker, 1983). Furthermore, the indentation method can
determine the surface properties of the cutaneous surface,
which reflect the physic-chemical properties of the skin/
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indenter interface (Pailler-Mattei and Zahouani, 2004).
Pailler-Mattei et al. (2008) conducted experiments with an
original indentation device and obtained an average value of
Young’s modulus for the skin between 4.5 kPa and 8 kPa,
which are quite close to the results reported by Bader and
Bowker (1983) and Pailler-Mattei and Zahouani (2004). The
mechanical methods and imaging methods were combined by
most researchers to measure the skin mechanical properties in
literature. For example, ultrasound combined with indentation
(Zheng and Mak, 1996), MRI combined with indentation (Tran
et al., 2008), and suction combined with ultrasound and OCT
(Hendriks et al., 2003, 2006) have all been used. More recently,
some researchers have used airflow to measure the skin
mechanical properties (Boyer et al., 2012; Fleury et al., 2010;
Fugimura et al., 2008). The mechanical properties of skin
vary with the skin locations on body (Fung, 1993; Wilkes et al.,
1973). Staloff et al. (2008) claimed that the results of tests on
other parts of the human body such as the forearm cannot be
used to replace the results gained from facial skin because the
facial skin has a very complex underlying connective structure.
Thus, the studies on facial skin provide more accurate data on
facial skin mechanical properties (Barbarino et al., 2011; Barel
et al., 1998; Cua et al., 1990; Couturaud et al., 1995; Malm et
al., 1995; Ohshima et al., 2011; Tsukahara et al., 2004).
In literature, researchers have used linear and nonlinear skin
mechanical properties in skin simulations. Many skin models
claim skin as isotropic, linear elastic material (Diridollou et al.,
2000; Pailler-Mattei et al., 2008; Khatyr et al., 2006). Young’s
modulus and Poisson’s ratio can be used to define the linear
skin model. However, the linear elastic material models are
difficult to model the nonlinear properties of the skin in some
cases. Mooney-Rivlin strain energy function was employed by
Hendriks et al. (2006) to model skin.
The strain energy density function for an incompressible
Mooney-Rivlin material is:
C and 2
C are empirically determined material
constants, and 1
are the first and the second
invariant of the unimodular component of the left Cauchy-
; 22 22 22
where F is the deformation gradient.
Evans (2009) used Odgen strain energy function with a pre-
strain parameter to simulate in-plane deformations of skin. The
Ogden strain energy for incompressible rubber-like material is
expressed in terms of the principal stretches i
are constant shear moduli and
dimensionless constants. The principal stresses are found from:
Arruda-Boyce strain energy function was employed by
Bischoff et al. (2000) to model the nonlinear skin. The strain
energy density function for the incompressible Arruda-Boyce
model is given by: sinh
WNk n n
where n is the number of chain segments,
k is the
is the temperature in Kelvin, Nis the
number of chains in the network of a cross-linked polymer,
I is the first invariant of the left Cauchy-Green
deformation tensor, and
is the inverse Langevin
Devices measuring facial skin mechanical properties
Boyer et al. (2012) developed an air flow device called
Tonoderm which consists of an air compressor and an Air Mass
Flow Controller, 5851S (Brooks Instrument, Hatfield, PA,
USA). Fleury et al. (2010) developed an air puff that can
deform the skin surface and measure the mechanical properties.
The air puff developed by Fleury et al. (2010) mainly consists
of an air port (Porter Instruments, Hatfield, PA, USA), a
pressure gauge (Kelatron, Coporation, Ogden, UT, USA), and a
flux meter , GPE meterate 314 (GPE Scientific Ltd,
Bedfordshire, England). Fujimura et al. (2008) developed an air
flow system to measure the skin mechanical properties which
consists of an air reservoir and a displacement sensor. The
displacement of the skin was recorded by Fujimura et al.
(2008) using a displacement sensor, Z4W-V25R (Omron,
Kyoto, Japan). Staloff et al. (2007) used digital image speckle
correlation (DISC) to measure the skin mechanical properties.
Staloff et al. (2007) took images in rapid succession using a
five megapixel camera, Canon D60 and Canon Powershot Pro1
(Lake Success, NY, USA). Mahmud et al. (2010) used a 3-
camera optial motion analysis system, Qualisys Proflex-
MCU1000 (Qualisys, Gothenburg, Sweden) with 50mm focal
length lenses. Hendriks et al. (2004) performed suction
measurements and obtained elevation and thickness of the skin
with OCT. The OCT device and suction device used by
Hendriks et al. (2004) were customized. Li et al. (2012) used a
home-made surface wave generator with an OCT system, SD-
OCT (DenseLight Semiconductors Ltd, Singapore) to observe
the deformation and mechanical properties of the skin. Liang
and Boppart (2010) used an OCT system and observed the skin
deformation with a mechanical wave driver, SF-9324 (PASCO
scientific, Roseville, CA). Liang and Boppart (2010) also
Copyright © 2015 by ASME
verified their results for mechanical properties of the skin with
a commercial instrument, Custometer MPA 580 (Courage
Khazaka Electronic, Koln, Germany). Gahagnon et al. (2009)
used a high frequency ultrasound scanner (Dermcup 2020, Atys
Medical, Soucieu en Jarrest, France) and an extensiometer
developed in LMARC laboratory and P. Fabre Research
Institute in France to develop skin imaging. Gennisson et al.
(2004) used a 50 MHz ultrasound probe (Panametrics, La
Garenne Colombe, France) with a vibrator, Brüel & Kjær 4810
(Brüel & Kjær Sound & Vibration, Nærum, Denmark). Lapeer
et al. (2010) used the Stable Micro System TA.XT2 texture
analysis machine (Stable Micro Systems Ltd., Godalming, UK)
to measure the human skin mechanical properties in vitro.
Khatyr et al. (2004) developed an extensometer which provides
a non-invasive measurement of the skin mechanical properties
on different parts of the body and in different directions at each
point. Lim et al. (2008) designed an extensometer to measure
the mechanical properties of skin. Wang and Hayward (2007)
developed an apparatus that can deform the skin in the large
deformation range with traction and measure the skin
properties. Pailler-Mattei and Zahouani (2004; 2008)
developed an experimental device which can study the
adhesion forces and mechanical properties of human skin in
vivo with an indentation test. Diridollou et al. (2000) developed
a suction test device which consists of a pressure control
system that can apply increasing and decreasing suction to the
skin, an electronic circuit that detects the displacement of the
skin and an ultrasound device. Hendriks et al. (2003) attached
an ultrasound system (DUB 20, Taberna Pro Medicum,
Germany) to a self-developed suction device to perform the
suction test in vivo and measured the mechanical properties of
the skin. Agache et al. (1980) introduced a torsional device
which contains a torque-motor and a rotational detector which
can provide voltage proportional to the rotational angle.
Facial muscle thickness
As mentioned in the Human Anatomy section, the major facial
muscles include masseter, levator labii superioris and
zygomaticus major muscles. The major facial muscles
thicknesses are most commonly measured by several in vivo
imaging techniques: ultrasound scanning, computerized
tomography (CT) and magnetic resonance imaging (MRI). CT
was used to measure the facial muscle thickness by several
researchers (Chabanas et al., 2003; Weijs and Hillen, 1984; Van
Spronsen et al., 1989; Katsumata et al., 2004). The results
provided by CT are considered to be reliable (Satiroglu et al.,
2005). However, the CT can unavoidably bring radiographic
exposure to the subjects (Satiroglu et al., 2005), which is a
known cumulative biological effect. MRI is a method that
received relatively little attention by researchers in literature
(Farrugia et al., 2007; Fischbein et al., 2001; Kaylie et al.,
2003; Von Sponsen et al., 1991; Hannam and Wood, 1989).
MRI was found to be accurate (Hannam and Wood, 1989);
however, the relatively high cost of MRI reduces its clinical
availability. In literature, the ultrasonography method has been
used by most of the researchers for measuring the facial muscle
thickness (Satiroglu et al., 2005; Castelo et al., 2007;
Gergiakaki et al., 2007; Kiliaridis and Kalebo, 1991; Kiliaridis
et al., 2003, 2007; Kubota et al., 1998; Muller et al., 2012;
Palinkas et al., 2010; Raadsheer et al., 1996). The
ultrasonography has been proved to be a simple, inexpensive
and accurate method to measure the muscle thickness
(Satiroglu et al., 2005) and it is applicable for large scale
studies (Radsheer et al., 1994). Also, ultrasonography does not
have known cumulative biological effects. Satiroglu et al.
(2005) used the ultrasonography method to measure the
thickness of the major facial muscles in adults, and results are
shown in Table 2. The influence of age and gender on the
thickness of human facial muscles was studied by Palinkas et
al. (2010), and the results are shown in Table 3.
Table 2 Mean and standard deviation of muscle thickness on
different facial sites for adults (Satiroglu et al., 2005)
Muscles Me an and SD (mm)
Masseter (relaxed) Males 14.92 (1.59)
Females 12.74 (1.69)
Combined 13.56 (1.95)
Masseter (clenching) Males 15.92 (1.89)
Females 13.76 (1.24)
Combined 14.57 (1.83)
Levator labii superioris Males 3.60 (0.49)
Females 3.41 (0.44)
Combined 3.48 (0.46)
Zygomaticus major Males 3.62 (0.38)
Females 3.33 (0.39)
Combined 3.44 (0.40)
Devices measuring muscle thickness
CT machines used by Katsumata et al. (2004) were the high
speed advantage (General Electronic, New York, NY, USA)
and the X-Vision (Toshiba, Otawara, Japan). Farrugia et al.
(2007) gained MRI with a 1.5T Siemens Sonata (Siemens
Medical Solutions, Erlangen, Germany). Castelo et al. (2007)
used the ultrasound system, Just Vision Toshiba (Toshiba,
Otawara, Japan) with a 10-MHz linear transducer. Gergiakaki
et al. (2007), Kiliaridis et al. (2003, 2007) and Raadsheer et al.
(1994) used real time ultrasound scanner, Pie Medical Scanner
480 (Pie Medical, Maastricht, The Netherlands) with a 7.5-
MHz linear array transducer. Kubota et al. (1998) used
ultrasonographic scanning system, SSD-500 (Aloka Co., Ltd,
Tokyo, Japan) with a 7.5 MHz linear type scanning probe.
Muller et al. (2012) used the real time ultrasound scanners,
FALCO I00, (Pie Medical, Maastricht, The Netherlands) and
MyLab25 (Esaote, Genova, Italy) with linear array transducers
(6-8 MHz, 4-13 MHz). Palinkas et al. (2010) used SonoSite
Titan ultrasound tool (Duluth, Georgia, USA) with a 56mm/10
MHz linear-array transducer. Satiroglu et al. (2005) used a real-
Copyright © 2015 by ASME
time ultrasound scanner, Siemens Elegra (Siemens, Erlangen,
Germany) with a 7.5-9 MHz broadhand transducer.
Table 3 Means and standard deviations of the thickness (cm) of
the right masseter (RM), left masseter (LM), right temporal
(RT), and left temporal (LT) in the five age groups in relation to
gender-male (M) and female (F) –in the clinical condition of
rest (Palinkas et al., 2010).
Age Groups Gender Muscles
RM LM RT LT
07-12 M 0.78 (0.04) 0.77 (0.03) 0.51 (0.03) 0.49 (0.03)
F 0.73 (0.04) 0.73 (0.03) 0.44 (0.03) 0.43 (0.03)
13-20 M 0.94 (0.04) 0.95 (0.03) 0.62 (0.03) 0.63 (0.03)
F 0.92 (0.04) 0.92 (0.04) 0.65 (0.03) 0.63 (0.03)
21-40 M 1.02 (0.04) 1.06 (0.04) 0.67 (0.03) 0.68 (0.03)
F 0.82 (0.04) 0.84 (0.04) 0.56 (0.03) 0.59 (0.03)
41-60 M 1.16 (0.04) 1.14 (0.03) 0.74 (0.03) 0.71 (0.03)
F 0.93 (0.04) 0.94 (0.03) 0.59 (0.03) 0.59 (0.03)
61-80 M 1.10 (0.06) 1.07 (0.06) 0.65 (0.05) 0.66 (0.05)
F 0.90 (0.06) 0.89 (0.05) 0.47 (0.04) 0.46 (0.04)
Facial Muscle Mechanical Properties
In the literature, there is no direct experimental data available
for human facial muscle mechanical properties. However, the
mechanical properties of human muscle on other sites of human
bodies and on animal (Linder-Ganz and Gefen, 2003) have
been studied by previous researchers. Several testing
techniques have been used to measure the mechanical behavior
of the human muscles. Compression tests were used to invest
the mechanical properties of skeletal muscles (Aritan et al.,
2008; Bosboom et al., 2001; Linder-Ganz and Gefen, 2004;
Livarinen et al., 2011; Van Loocke et al., 2006; Van Loocke et
al., 2008). Some researchers performed indentation tests on
skeletal muscles (Gefen et al., 2005; Livarinen, et al., 2014;
Palevski et al., 2006; Zheng et al., 1999). Some other
researchers use ultrasound (Blackburn et al., 2009; Gennisson
et al., 2010, 2005; Nordez et al., 2008; Nordez and Hug, 2010)
to measure the muscle mechanical properties. Some researchers
(Chabanas and Payan, 2000; Chabanas et al., 2003; Gladilin et
al., 2004) used the Young’s modulus values reported by Duck
(1990) in their linear elastic FEM facial muscle simulation
models. Duck (1990) reported 6.2 kPa for a human muscle in
its rest position and 110 kPa for the same muscle, which are
close with the measurements reported by Ohayon et al. (1999)
on cardiac muscles. Some researchers developed non-linear
models for muscle tissues. Hyperelastic muscle models were
performed by Cox et al. (2007), Jia et al. (2006) and Linder-
Ganz and Gefen (2004). Viscoelastic muscle models were used
by Aritan et al. (2008), Best et al. (1994), Bosboom et al.
(2001) and Van Loocke et al. (2008). The detailed parameters
for the mechanical properties of muscle in those models are
measured and reported.
Devices measuring muscle mechanical properties
Aritan et al. (2008) designed a creep testing machine to
quantify the non-linear viscoelastic behavior of muscular bulk
tissue, which consist of a main frame, a squeezing chamber, a
tensional load cell and a data collection unit. Bosboom et al.
(2001) performed compression test with Zwick 1445 machine
(Zwick Roell, Ulm, Germany). Linder-Ganz and Gefen (2004)
used compression test with a testing machine, INSTRON 5544
(Instron, Norwood, Massachusetts, USA). Van Loocke et al.
(2006, 2008) performed compression tests with a uniaxial test
machine, Zwick Z005 (Zwick Roell, Ulm, Germany). Nordez
and Hug (2010) used an AixPlorer ultrasonic scanner
(Supersonic Imagine, Aix en Provence, France), coupled with a
linear transducer array (4-15 MHz, SuperLinear 15-4, Vermon,
Tours, France) in their test. Gefen et al. (2005) developed an
electromechanical, computer-controlled and precalibrated
indentor with a miniature linear stepper motor, force transducer
and a linear variable displacement transducer to measure the
indentation force and material displacement. Livarinen et al.
(2014) developed a hand-held indentation device to measure
the muscle mechanical properties, which consist of two load
cells: 10 kg Honeywell Sensotec load cell 31/1430-04 Mid
(Honeywell Sensotec, Columbus, OH, USA) and 1000 g
Honeywell Sentec load cell 31/ 1426-02 Mid (Honeywell
Sensotec, Columbus, OH, USA). Palevski et al. (2006)
developed a pneumatic indentation testing device which can
have an indentation speed of 2000 mm/s during the ramp
phase. Gennisson et al. (2005) developed an ultrasound
indentation vibrator with a measuring excitor, Brüel & Kjær
4810 (Brüel & Kjær Sound & Vibration, Nærum, Denmark) to
measure the hardness of the muscle in vivo. Nordez et al.
(2008) used a shear elasticity probe which is composed of an
ultrasonic 5 MHz transducer and a vibrator, Brüel & Kjær 4810
(Brüel & Kjær Sound & Vibration, Nærum, Denmark) to
measure the hardness of the muscle in vivo. Blackburn et al.
(2009) used B-mode ultrasound device, LOGIQe (General
Electronic, Milwaukee, WI, USA) with a 7 MHz linear-array
transducer. Livarinen et al. (2013) used a suction device made
of thermoplastic elastomer SEBS, Thermolast (Kraiburg TPE
GmbH & Co. KG, Waldkraiburg, Germany) to measure the
forearm muscle mechanical properties.
SUBCUTANEOUS ADIPOSE TISSUE
Adipose tissue thickness
In literature, there is no direct data about the subcutaneous
layer thickness on healthy human face. However,
subcutaneous tissue thickness on other sites of human body
have been measured by researchers using computed
tomography (CT) (Burbridge et al., 2007) and ultrasound
method (Gibney et al., 2010; Guaraldi et al., 2005; Nordander
et al., 2003; Presti et al., 2012; Rolfe et al., 2010; Smith et al.,
1991). As mentioned in the former chapters in this paper, CT
has potential biomedical hazard because of the radiation
exposure. Most of the researchers in literature chose ultrasound
to measure the subcutaneous tissue thickness under skin
Copyright © 2015 by ASME
because of the effectiveness and accuracy of ultrasound.
Guaraldi et al. (2005) measured the facial subcutaneous
thickness change on HIV patients using ultrasound device.
Presti et al. (2012) performed ultrasound measurement of
subcutaneous thickness at the insulin injection sites on
Devices measuring subcutaneous tissue thickness
Gibney et al. (2010) used high frequency Cortex Dermascan C
ultrasound unit with a 20 MHz probe and a GE LOGIQ e
ultrasound unit (General Electric Healthcare, Little Chalfont,
Buckinghamshire, United Kingdom) with a 4.5-13.0 MHz
probe. Nordander et al. (2003) used ultrasound equipment with
a linear array probe and a frequency of 12 MHz (Logiq 700,
General Electric OEC Medical Systems, Salt Lake City, Utah,
USA) to measure the distance between skin surface and muscle
belly. The thickness of the subcutaneous tissue was measureb
by Nordander et al. (2003) using a skinfold caliper (Harpenden,
British Indicators, West Sussex, UK). Presti et al. (2012)
measured the thickness of subcutaneous tissue using the
MyLabTouch portable US unit with a 33-mm, 13.6-MHz
transducer/probe (ESAOTE Biomedica Deutschland GmbH,
Koln, Germany). Rolfe et al. (2010) measured the visceral and
subcutaneous abdominal fat thickness with a Logic Book XP
ultrasound device and a 3C MHz-RS abdominal curved array
transducer (General Electric Healthcare, Little Chalfont,
Buckinghamshire, United Kingdom). Smith et al. (1991) used a
Harpenden stadiometer (Harpenden, British Indicators, West
Sussex, UK) to measure the thickness of subcutaneous tissue.
Smith et al. (1991) measured the distances from skin to muscle
fascia using high resolution real time ultrasonography with an
electronically focused 5 MHz linear array transducer.
Subcutaneous tissue mechanical properties
No facial subcutaneous tissue mechanical properties were
reported in literature. However, the mechanical properties of
subcutaneous tissue tested on other sites of human body, like
breast (Chung et al., 2007; Samani et al., 2001; Samani and
Plewes, 2004; Wang et al., 2009), heel (Miller-Young et al.,
2002), fingertip (Wu et al., 2003), porcine tissue (Zheng and
Mak, 1996) or animals (Wu et al., 2004, 2007) were published.
The subcutaneous mechanical properties were tested either in
vitro or in vivo. Wu et al. (2004, 2007) conducted in vitro tests
on pig subcutaneous tissues and created nonlinear-elasticity
models to describe the mechanical behavior of the tissues.
Miller-Young et al. (2002) used specimen from human
cadaveric feet to test the fat pad on heel. The in vivo methods
include compression method (Wang et al., 2009; Wu et al.,
2007), tensile method (Chung et al., 2008), indentation method
(Samani and Plewes, 2004; Zheng and Mak, 1996) ultrasound
(Zheng and Mak, 1996) and MRI (Samani et al., 2001). The
subcutaneous tissue is believed to behave nonlinear elastic
mechanical properties (Wu et al., 2007). Several non-linear
models were introduced in literature. Miller –Young introduced
a nonlinear viscoelastic model for the fat pad on human heel
and specified parameters for the Mooney-Rivlin material
equations. Chung et al. (2008) specified the parameters for neo-
Hookean equation used for a non-linear elastic model of human
breast. Samani et al. (2001) and Samani and Plewes (2004)
introduced hyperelastic models for the human breast and
provided the parameters.
Devices measuring subcutaneous tissue mechanical
Chung et al., (2008) developed a tensile testing device which
consisted of a force transducer coupled to a displacement
actuator. Samani and Plewes developed an indentation
measuring system which consists of a programmable high
resolution linear servo actuator LAL30 (SMAC, Carlsbad, CA,
USA). Samani et al. (2001) used GE SIGNA 1.5-T scanner (GE
Medical Systems, Milwakee, WI) for the MRI. Wu et al.(2007)
performed compression tests with a universal micro-mechanical
testing mashine (Model Mach-I, Biosyntech, Montreal,
Canada). Zheng and Mak (1996) used Panametrics model
5052UA ultrasound (General Electric Healthcare, Little
Chalfont, Buckinghamshire, UK) pulser/receiver as both an
ultrasound source and indenter.
A detailed overview of various study of the thickness and
mechanical properties of human facial soft tissues and the
models for the facial soft tissues are presented in this paper.
The human facial soft tissues are believed to have non-linear
elastic, anisotropic and multi-layer structures. The measuring
methods and devices for the thickness and mechanical
properties of different layers of the human soft tissues were
overviewed in this paper. The measuring methods are either in
vivo or in vitro. Several finite element models were created for
different layers for the facial soft tissues. More detailed and
accurate measurement for the thickness and mechanical
properties of each layers of human facial soft tissues are still
needed in the future. Future work is needed to create a non-
linear elastic finite element model that can precisely describe
the mechanical behavior of each layer of human facial soft
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