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Medical Engineering and Physics 38 (2016) 1505–1512
Contents lists available at ScienceDirect
Medical Engineering and Physics
journal homepage: www.elsevier.com/locate/medengphy
Technical note
Anisotropic abdominal aortic aneurysm replicas with biaxial material
characterization
Sergio Ruiz de Galarreta
a , Raúl Antón
a , ∗, Aitor Cazon
a , Gorka S. Larraona
a , Ender A. Finol b
a Department of Mechanical Engineering, Tecnun, University of Navarra, Paseo Manuel de Lardizabal, 13, 20018 San Sebastián, Spain b Department of Biomedical Engineering, The University of Texas at San Antonio, One UTSA Circle, AET 1.360, San Antonio, TX 78249-0669, United States
a r t i c l e i n f o
Article history:
Received 10 November 2015
Revised 16 August 2016
Accepted 23 September 2016
Keywords:
Anisotropy
AAA phantom
Biaxial
Stereovision
Vacuum casting
Additive manufacturing
a b s t r a c t
An Abdominal Aortic Aneurysm (AAA) is a permanent focal dilatation of the abdominal aorta at least
1.5 times its normal diameter. The criterion of maximum diameter is still used in clinical practice, al-
though numerical studies have demonstrated the importance of other biomechanical factors. Numerical
studies, however, must be validated experimentally before they can be clinically implemented. We have
developed a methodology for manufacturing anisotropic AAA replicas with non-uniform wall thickness.
Different com posites were fabricated and tested, and one was selected in order to manufacture a phan-
tom with the same properties. The composites and the phantom were characterized by biaxial tensile
tests and a material model was fit to the experimental data. The experimental results were compared
with data from the literature, and similar responses were obtained. The anisotropic AAA replicas with
non-uniform wall thickness can be used in benchtop experiments to validate deformations obtained with
numerical simulations or for pre-intervention testing of endovascular grafts. This is a significant step for-
ward considering the importance of anisotropy in numerical simulations.
© 2016 IPEM. Published by Elsevier Ltd. All rights reserved.
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. Introduction
An Abdominal Aortic Aneurysm (AAA) is a permanent dilata-
ion of the abdominal aorta, and its rupture remains a significant
ause of death in developed countries among men aged 65–85 [1] .
n clinical practice, uncertainty still remains about the correct time
o operate, but the criterion of maximum diameter is commonly
ccepted as a rupture prediction factor, meaning that medical doc-
ors recommend surgical interventions when AAA diameters are
reater than 55 mm [2] . When studying AAA rupture, numerical
imulations via Finite Element Analysis (FEA) have proved to be
ery useful in indicating that this rupture criterion needs to be
omplemented with AAA wall stress data [3–5] , meaning that fac-
ors such as geometry and biomechanics must also be considered
6,7] .
In vitro experiments with AAA phantoms have been shown
o be very useful in several applications [8–10] . These phantoms
anged from totally idealized geometries to real AAA geometries
enerally with uniform [11,12] but also with non-uniform thick-
ess [13] . However, the anisotropy found in the AAA tissue [14–
7] is an important property that was not considered in these
rojects. When using inverse analysis the mechanical properties
∗ Corresponding author. Fax: + 34 943 31 14 42.
E-mail address: [email protected] (R. Antón).
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ttp://dx.doi.org/10.1016/j.medengphy.2016.09.010
350-4533/© 2016 IPEM. Published by Elsevier Ltd. All rights reserved.
re not critical [18,19] , but the consideration of the anisotropy pa-
ameter plays an important role in the results of numerical simu-
ations with a forward approach, and it must be taken into account
or further studies [20–22] . Within the context of physical replicas,
his consideration points to a step forward in the manufacturing of
hantoms with arterial anisotropic behavior.
The purpose of the present work is to describe and apply
new methodology for manufacturing AAA replicas that display
nisotropic behavior. To the best of the authors’ knowledge, this is
he first time that a methodology is reported for creating arterial
eplicas with non-uniform wall thickness and defined anisotropy.
. Materials and methods
.1. Uniaxial testing of isotropic specimens
Tensile tests were carried out following ASTM D412 Type B.
pecimens were manufactured via the vacuum casting technique
y using the bi-components PUR SLM 7140, 7160 and 7190 at seven
ifferent mixing ratios.
Tensile tests were performed on the specimens to generate
orce-extension data using an INSTRON MINI 44 (Instrom World-
ide, Norwood, MA) tensile test machine. Each specimen was sub-
ected to a cross-head speed of 3.4 mm/min until failure with
re-conditioning for 10 cycles to 7.5% of the gauge length. The
1506 S. Ruiz de Galarreta et al. / Medical Engineering and Physics 38 (2016) 1505–1512
Fig. 1. Complete process for creating composite specimens: (a) mold, (b) fibers, (c)
fibers placed in the mold before vacuum casting, (d) composite specimen.
Fig. 2. Complete biaxial testing system (left) and stereovision system (right).
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force-extension data from the uniaxial tests were converted to
strain and Cauchy stress ( Eqs. (1 ) and ( 2 )),
ε =
�l
l 0 (1)
σ =
F
A
∗ (2)
where �l is the change in specimen length at any time, l 0 is the
original length, F is the force required and A
∗ is the area at any
instant ( Eq. (3 ) assumes incompressibility).
A
∗ =
A 0 l 0 l 0 + �l
(3)
As the mechanical behavior of the components was quasi-
linear, a trend line for each material was fitted to the data
( R 2 = 0.96 ± 0.04) and stiffness was directly derived.
2.2. Biaxial testing of composite specimens
To reproduce the anisotropic behavior, various composite spec-
imens (3 mm-thick 25 × 25 mm
2 ) were fabricated, altering the fol-
lowing properties:
- f, proportion of fibers i.e. fiber volume/total volume ratio.
- Em , matrix elastic modulus.
- Ef , fiber elastic modulus.
Variations of f , Em and Ef were tested and the composite com-
bination with the properties closest to AAA tissue was selected. As
these properties were compared with Vande Geest’s data [16] , the
specimens’ geometries were the same, i.e. squares.
The procedure for obtaining the specimens with fibers inside
them was as follows. The specimen was modeled using CREO 2.0
and printed with an OBJET EDEN 330 Additive Manufacturing (AM)
printer. The specimen was used as a model for creating a silicone
mold that defined the external shape of the specimen ( Fig. 1 a). The
fibers were then attached to the mold ( Fig. 1 c), the PUR resin was
poured to fill the mold, and when cured the silicone mold was
opened and the specimen removed ( Fig. 1 d).
Tests on each specimen were conducted using a custom-made
planar biaxial testing system ( Fig. 2 ) similar to the one described
y Raghavan et al. [23] . Each sample was loaded with the help of
ixteen hooks (four at each side) that were connected to sixteen
ontainers able to hold weights. The load at each point was con-
rolled by gradually placing weights into the containers. Four load-
ng scenarios (240 g, 480 g, 720 g and 960 g) were considered. The
train measurement was calculated through the binocular stereo-
ision technique. Four small markers forming a 5 x 5 mm
2 were
laced in the center of the testing specimen for optical tracking.
his technique allows the markers’ 3D coordinates to be com-
uted by triangulation from a pair of images, which were cap-
ured by a pair of Logitech QuickCam E3500 webcams (resolution
0.03 mm/pixel) mounted at the top of the device. Before running
he biaxial tensile tests, the cameras were calibrated with an open-
ource code in MATLAB (Mathworks, Natick, MA, USA) language to
nsure accuracy [24] . To inspect the stereovision system’s accuracy,
he same template used for the calibration was recorded in 6 dif-
erent positions. For each position, 8 measurements were taken in
ifferent directions. The known distances were compared to the
nes calculated via stereovision and the errors were calculated as
percentage:
( % ) = | RD − MD | /RD (4)
here e is the error, RD is the real distance between selected
oints and MD is the distance measured by the stereovision sys-
em. The average error (SD) of the stereovision system was equal
o 0.58% (0.37%), i.e. a maximum error of 48 μm for 5 mm lengths
distance between markers).
Sample thickness was measured several times with a digi-
al caliper before testing, and the average thickness was used in
ubsequent stress calculations. The specimen was tested using a
tress-controlled protocol, where the first Piola–Kirchhoff stresses
i.e. the engineering stresses) served as a measure. The non-zero
omponents of the first Piola–Kirchhoff stress tensor P have the
orm:
θθ = f θ /T X L , P LL = f L /T X θ (5)
here f θ and f L denote the forces in each direction, T the thick-
ess in the unloaded configuration, and X θ and X L the dimensions
etween the hooks along the circumferential and longitudinal di-
ections of the square specimen, i.e. 20 mm. Each biaxial specimen
as tested in the following order: P θθ : P LL = 1:1, 0.75:1, 1:0.75,
.5:1, 1:0.5, 1:1, keeping the ratio P θθ : P LL constant for each pro-
ocol. The last equibiaxial tension protocol (i.e., P θθ : P LL = 1:1) was
erformed to confirm that no structural damage occurred in the
pecimen. Each specimen was preconditioned through six loading
nd unloading cycles, and the seventh cycle was used for the sub-
equent analysis.
From the recorded marker positions, deformation gradient ten-
or F was calculated at each measured value of imposed load [25] .
reen strain tensor E was calculated as denoted in Eq. (6 ). The
hear components of deformation gradient tensor F were found to
S. Ruiz de Galarreta et al. / Medical Engineering and Physics 38 (2016) 1505–1512 1507
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e negligible, so the in-plane Green strain tensor components were
etermined with Eq. (7 )
=
1
2
(F T F − 1
)(6)
θθ =
1
2
(λ2
θ − 1
), E LL =
1
2
(λ2
L − 1
)(7)
ith λθ and λL denoting the stretches in both directions.
Constitutive modeling
To model the mechanical response of the materials tested, we
sed strain energy function W ( Eq. (8 )), developed by Choi and
ito [26] and used by Vande Geest et al. [16] for both aneurysmal
nd non-aneurysmal abdominal aortic tissue:
= b 0 (e ( 1 / 2 ) b 1 E
2 θθ + e ( 1 / 2 ) b 2 E
2 LL + e b 3 E θθ E LL − 3
)(8)
here b 0 , b 1 , b 2 and b 3 are the material coefficients to be
etermined.
From Eq. (8 ) the in-plane second Piola–Kirchhoff stresses ( Eqs.
10 ) and ( 11 )) can be determined, applying Eq. (9 ).
=
∂W
∂E
(9)
θθ = b 0 (b 1 E θθ e ( 1 / 2 ) b 1 E
2 θθ + b 3 E LL e
b 3 E θθ E LL
)(10)
LL = b 0 (b 2 E LL e
( 1 / 2 ) b 2 E 2 LL + b 3 E θθ e b 3 E θθ E LL
)(11)
The data from the five biaxial protocols ( P θθ : P LL = 1:1, 0.75:1,
:0.75, 0.5:1, 1:0.5) for each specimen were fitted to this model
nd the material coefficients were derived for individual samples.
n order to derive a single constitutive model, data from each pro-
ocol was averaged to obtain a single dataset of the composite.
Additionally, the anisotropy parameter [16,26] :
I =
√
b 1 / b 2 (12)
as calculated.
.3. Anisotropic AAA phantom
The above process for creating anisotropic specimens was used
o manufacture an anisotropic AAA replica. Once the fiber volume
raction ( f ) is defined, the dimensions of the fibers need to be cal-
ulated. Due to the non-uniformity of the AAA wall thickness, the
ber dimensions should be variable in order to achieve the desired
along the whole AAA phantom. To that end the following pro-
ess was followed using MAGICS v16.02 (Materialise, Leuven, Bel-
ium). The AAA geometry was cut using nine planes normal to the
ongitudinal direction and spaced 10 mm apart. As a result, 8 dif-
erent cylindrical slices with non-uniform thickness were obtained.
hen each slice was divided into eight parts, and for each part the
verage thickness was measured. Finally the appropriate fiber di-
ensions for each slice were designed by considering the selected
alue of f .
Using these fibers, the process for creating the phantom was
ased on a previous work that describes how to build isotropic
AAs from medical images [13] . The patient-specific AAA geom-
try was obtained by segmenting CT images from Allegheny Gen-
ral Hospital (Pittsburgh, PA) using in-house software (AAAVASC,
niversity of Texas at San Antonio, San Antonio, TX) capable of
dentifying the boundaries of the lumen and the inner and outer
all surfaces [27,28] . Based on this virtual geometry, a rigid phys-
cal replica was created using an OBJET EDEN 330 AM printer. The
rinted artery was used as a model to create an outer silicone
old that defined the external shape of the artery and an inner
ax mold that defined its internal geometry. The wax mold was
laced inside the silicone mold, and the PUR 7140, 7160 or 7190
as poured to fill the gap between the inner and outer molds with
he help of a MCP 4/01 vacuum casting machine. When the gap
as filled, the vacuum was released and the mold was placed in
n oven at 45 °C to cure the resin. After 24 h of curing, the oven
emperature was increased to 85 °C to melt the inner wax. At the
nd of the melting process the silicone mold was opened and the
ubber-like artery removed.
To manufacture the anisotropic AAA, the abovementioned pro-
edure was slightly modified to add the fibers into the AAA prior
o the vacuum casting process with PUR resin. Each fiber was at-
ached to the external surface of the wax mold by gluing its two
nds with cyanoacrylate glue, i.e. Superglue 3 (Loctite, Düsseldorf,
ermany). Each fiber was fixed along the circumferential direction
Fig. 3 ). Next, the wax mold with its fibers was placed inside the
ilicone mold and the process mentioned above was subsequently
ollowed.
The AAA phantom was created with the same properties as the
ested composite. From this phantom, two square specimens were
ut for subsequent biaxial analysis. Before the tensile tests, sample
hickness was averaged by measuring it ten times with a digital
aliper. The same protocol was followed for the biaxial analysis.
. Results
.1. Uniaxial testing of isotropic specimens
A total of 42 specimens were uniaxially tested, six per material.
able 1 shows the average stiffness of each component, together
ith 95% confidence intervals. The stress–strain curves and the ul-
imate tensile strength of the materials are included in the Supple-
entary material.
.2. Biaxial testing of composite specimens
Considering the results of the uniaxial and biaxial tests (listed
n the Supplementary material), six composite specimens with the
ollowing parameters were fabricated and biaxially tested: f = 0.15,
m = 0.54 MPa (Component 7) and Ef = 1.64 MPa (Component 5).
he choice of this composite was made by comparing the over-
ll stiffness and the grade of anisotropy between the AAA tissue
16] and the different composites.
For all tested specimens, the results from the first and last
quibiaxial protocol coincided and thus suggested that no struc-
ural damage of the tissue occurred as a result of testing. The cir-
umferential and longitudinal experimental results for all the spec-
mens are shown in Tables 2 and 3 , respectively. The experimental
ata was averaged for this composite, and the representative S –E
lots are displayed in Fig. 4.
In addition, the ratio between the maximum Green strain in the
ongitudinal and the circumferential ( E LL,max /E θθ ,max ) directions was
alculated with an average (SD) equal to 1.77 (0.50). These values
re also shown in Table 4.
.3. Mathematical model of composite and phantom specimens
The material parameters for the composite specimens are re-
orted in Table 4 and shown graphically in Fig. 4 . As Fig. 4 il-
ustrates, the material model fit very well to the average compos-
te data ( R 2 = 0.98); it also fits the individual composite specimens
R 2 = 0.96) and AAA phantom specimens ( R 2 = 0.96) ( Table 4 ).
As a measure of overall stiffness, the strain energy at an equib-
axial nominal stress of 60 kPa was also calculated for each speci-
en and is reported in Table 4 together with the anisotropy factor
I .
The results derived from the AAA phantom specimens were
imilar and the model is shown in Fig. 4 , with the corresponding
1508 S. Ruiz de Galarreta et al. / Medical Engineering and Physics 38 (2016) 1505–1512
Fig. 3. PUR fibers attached to the wax inner mold (left) and as a part of the final anisotropic AAA physical replica (right).
Table 1
Stiffness average and stiffness extreme values (confidence interval 95%) of the seven components.
# Component SLM material A:B mixing ratio Stiffness mean [MPa] (95% C.I. n = 6)
1 7190 100:92 12.37 (11.50, 13.24)
2 7190 100:85 6.65 (5.81, 7.49)
3 7190 100:75 3.37 (2.62, 4.11)
4 7160 100:69 2.87 (2.75, 3.00)
5 7160 100:60 1.64 (1.49, 1.80)
6 7140 100:45 1.11 (1.07, 1.15)
7 7140 100:38 0.54 (0.45, 0.62)
Table 2
Nominal Stress–stretch response of specimens in the circumferential direction.
Protocol Nominal stress (kPa) Stretch Aver. stretch SD
V1 V2 V3 V4 V5 V6
1:1 39 .24 1.014 1.006 1.005 1.008 1.017 1.013 1.010 0.005
78 .48 1.028 1.008 1.014 1.028 1.025 1.023 1.021 0.008
117 .72 1.043 1.018 1.024 1.039 1.043 1.032 1.033 0.011
156 .96 1.059 1.031 1.036 1.053 1.067 1.040 1.048 0.014
1:0.75 39 .24 1.020 1.011 1.008 1.016 1.025 1.016 1.016 0.006
78 .48 1.036 1.018 1.018 1.028 1.044 1.029 1.029 0.010
117 .72 1.057 1.031 1.029 1.054 1.074 1.041 1.048 0.017
156 .96 1.075 1.045 1.043 1.072 1.094 1.061 1.065 0.020
1:0.5 39 .24 1.025 1.019 1.007 1.020 1.026 1.020 1.019 0.007
78 .48 1.051 1.032 1.023 1.045 1.055 1.032 1.040 0.013
117 .72 1.073 1.050 1.036 1.066 1.086 1.057 1.061 0.018
156 .96 1.098 1.067 1.056 1.090 1.110 1.074 1.082 0.020
0.75:1 29 .43 0.990 1.005 0.998 1.007 1.011 1.008 1.003 0.008
58 .86 1.002 1.002 1.002 1.013 1.018 1.010 1.008 0.007
88 .29 1.012 1.009 1.004 1.019 1.029 1.010 1.014 0.009
117 .72 1.020 1.008 1.010 1.025 1.036 1.020 1.020 0.010
0.5:1 19 .62 0.996 1.006 1.0 0 0 0.993 1.0 0 0 1.003 1.0 0 0 0.005
39 .24 0.990 1.0 0 0 0.998 0.998 0.999 1.004 0.998 0.005
58 .86 0.990 0.999 0.999 0.998 1.006 1.003 0.999 0.005
78 .48 0.984 0.992 0.991 1.0 0 0 0.999 0.994 0.993 0.006
S. Ruiz de Galarreta et al. / Medical Engineering and Physics 38 (2016) 1505–1512 1509
Table 3
Nominal Stress–stretch response of specimens in the longitudinal direction.
Protocol Nominal stress (kPa) Stretch Aver. stretch SD
V1 V2 V3 V4 V5 V6
1:1 39 .24 1.021 1.027 1.021 1.011 1.028 1.021 1.021 0.006
78 .48 1.047 1.049 1.041 1.028 1.051 1.031 1.041 0.010
117 .72 1.068 1.064 1.058 1.049 1.075 1.047 1.060 0.011
156 .96 1.101 1.079 1.074 1.065 1.101 1.057 1.080 0.018
1:0.75 29 .43 1.014 1.025 1.018 1.007 1.023 1.014 1.017 0.006
58 .86 1.024 1.034 1.026 1.015 1.037 1.022 1.026 0.008
88 .29 1.045 1.046 1.037 1.022 1.044 1.029 1.037 0.010
117 .72 1.061 1.054 1.041 1.030 1.056 1.033 1.046 0.013
1:0.5 19 .62 1.006 1.018 1.008 1.002 1.007 1.009 1.008 0.005
39 .24 1.005 1.018 1.011 1.004 1.012 1.009 1.010 0.005
58 .86 1.012 1.021 1.011 1.001 1.014 1.008 1.011 0.006
78 .48 1.016 1.021 1.013 1.002 1.019 1.006 1.013 0.008
0.75:1 39 .24 1.033 1.036 1.021 1.020 1.033 1.022 1.028 0.007
78 .48 1.069 1.052 1.049 1.037 1.058 1.040 1.051 0.012
117 .72 1.094 1.073 1.068 1.054 1.090 1.060 1.073 0.016
156 .96 1.127 1.098 1.09 1.082 1.124 1.072 1.100 0.022
0.5:1 39 .24 1.030 1.037 1.036 1.025 1.039 1.028 1.033 0.005
78 .48 1.068 1.060 1.055 1.046 1.071 1.047 1.058 0.010
117 .72 1.105 1.094 1.076 1.074 1.107 1.068 1.087 0.017
156 .96 1.156 1.113 1.105 1.098 1.151 1.093 1.119 0.028
Fig. 4. S –E plots of the composite average (top) and phantom specimen experimental data with the corresponding material model for the circumferential (left) and longitu-
dinal (right) direction.
1510 S. Ruiz de Galarreta et al. / Medical Engineering and Physics 38 (2016) 1505–1512
Table 4
Model parameters for average and individual specimen fits to Eqs. (8) –( 11 ), anisotropy factor ( AI ), peak
Green strain ratio and W values.
Specimen b 0 (kPa) b 1 b 2 b 3 AI E LLmax /E θθmax W 60 (kPa) R 2
V1 2787.43 0.48 0.35 0.24 1.18 1.76 1.91 0.98
V2 2482.89 0.83 0.50 0.34 1.29 2.62 1.48 0.92
V3 1278.27 1.96 1.10 0.74 1.33 2.06 1.31 0.97
V4 3205.90 0.51 0.47 0.27 1.04 1.25 1.41 0.98
V5 2590.61 0.45 0.36 0.21 1.12 1.54 2.13 0.97
V6 3784.70 0.51 0.42 0.25 1.09 1.43 1.29 0.97
Composite average 3716.25 0.45 0.33 0.21 1.17 1.77 1.58 0.98
P1 2696.18 0.45 0.35 0.19 1.14 1.81 2.15 0.96
P2 3906.77 0.33 0.21 0.12 1.24 1.98 2.29 0.96
Table 5
AI , Green strain ratio and W compared against Vande Geest’s data, composite specimens and AAA phantom specimens.
Vande Geest ( n = 26) Composite ( n = 6) AAA phantoms ( n = 2) Differences com posite Vande Geest (%) Differences phantom composite (%)
AI 1.13 ± 0.27 1.19 ± 0.07 1.29 ± 0.09 3 .54 1 .71
E LLmax /E θθmax 1.61 ± 1.08 1.77 ± 0.50 1.89 ± 0.12 9 .94 6 .78
W 60 1.25 ± 0.47 1.58 ± 0.35 2.22 ± 0.09 26 .40 40 .51
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material coefficients illustrated in Table 4 . The differences in AI ,
E LL,max /E θθ ,max and W between the composite specimens and AAA
phantom specimens are not statistically significant, with p -values
equal to 0.871, 0.764 and 0.053, respectively.
4. Discussion and conclusions
The present work describes a methodology for manufacturing
patient-specific replicas of arteries with regionally varying wall
thickness and an overall anisotropic behavior. By varying the com-
posite parameters ( f, E m
and E f ), different mechanical properties
can be obtained for the AAA phantoms. It should be noted that
the fibers employed in this study are not intended to imitate the
wavy and dispersed distribution of collagen fibers nor their thick-
ness (0.8–2.4 μm); that cannot be achieved even with a 3D printer.
The purpose of including the fibers was to provide the AAA phan-
tom with anisotropic behavior at a macro scale, which was verified
by the experiments.
Two simplifications were made in this study. The first one was
the omission of the thrombus and calcifications, present in 75% of
AAAs [29,30] , as they are beyond the scope of this study. How-
ever, the inclusion of the thrombus in idealized AAA replicas was
studied by Corbett et al. [31] and could be implemented in this
methodology in an analogous way. The second simplification was
that we treated AAA tissue behavior as being quasi linear, while
several studies [16,17,32] have revealed the nonlinear behavior of
AAA tissue. However, the stress–strain curves reported in those
studies were obtained under a zero-stress condition, neglecting the
fact that in-vivo tissue is exposed to a stressed configuration [33] .
The residual strain [34] when tissue is excised was not contem-
plated either. Thus considering these factors, we believe that a re-
alistic AAA replica should mimic the second region of the curve,
which is why the material properties of the manufactured phan-
toms were contrasted with that region. Furthermore, small strain
ranges in a hyperelastic material can be considered linear since
the physiological strain ranges due to the pulsatile hemodynam-
ics are small; therefore mimicking the anisotropic range linearly is
a good first approach. In order to compare the grade of anisotropy
in the AAA tissue [16] and our composites, two parameters were
selected: AI and the mean peak Green strain ratio ( E LL,max /E θθ ,max ).
As a measure of overall stiffness, the strain energy at an equibiaxial
nominal stress of 60 kPa ( W 60 ) was also compared. As mentioned
above, in vivo the realistic material properties correspond to the
inear region of the stress–strain curves (we considered it to be
bove 10 kPa). Hence, in order to estimate W 60 from Vande Geest’s
ata, we used Eq. (13 ).
60 = ( W 70 − W 10 ) −∑
i = θ,L
[S i, P I =10
(E i, P i =70 − E i, P i =10
)](13)
The differences between the AAA samples and our data are
mall ( Table 5 ).
Comparing our tested composite with Vande Geest’s data, the
erived differences are relatively low ( Table 5 ) and acceptable,
ith a maximum difference equal to 26.40% in the W 60 parameter.
he differences in the factors measuring the grade of anisotropy
re lower than 10%. As the manufactured AAA phantom has the
atient-specific geometry and non-uniform wall thickness, it is not
urprizing that the difference between the composite and the AAA
eplica specimens exists because the experiments were not carried
ut under the same conditions. That is, while the composite
pecimens were completely planar and had a uniform thickness,
he AAA phantom specimens were not perfectly planar due to
urvature, and the thickness varied due to the patient-specific AAA
eometry. Both factors influenced the experimental results. An-
ther influential factor is the thickness error of the phantom due
o the manufacture process (average dimensional mismatch of 180
icrons, 11.14% [13] ), which changes the proportion of the fibers
nd thus the mechanical properties. This is why the main differ-
nce between the phantom and composite specimens ( p = 0.053)
s factor W 60 , with an average difference equal to 40.51%. However,
I index and Green strain ratio do not differ significantly between
he phantom and composite specimens ( p = 0.871 and p = 0.764),
ith differences lower than 7%. Apart from the composite shown
n this study, other composites have been tested and some of
hem are shown in Fig. 5 with the corresponding composite
arameters, material model, AI and W values. It is worth not-
ng that before the tensile tests, the specimens’ thickness was
easured by a digital caliper due to its simplicity, although it
s not devoid of measuring errors and a thickness gauge would
e a more accurate system [35] . The influence of fiber diameter
as also studied by manufacturing two specimens with the same
omposite parameters but a different number of fibers (2 and
). They were biaxially tested and similar results were obtained,
ndicating that fiber diameter is not as critical as the other com-
osite parameters. The results are included in the Supplementary
aterial.
S. Ruiz de Galarreta et al. / Medical Engineering and Physics 38 (2016) 1505–1512 1511
Fig. 5. S –E plots, AI and W 60 values for the two other tested composites.
t
n
t
p
n
fl
o
a
c
t
f
a
O
i
p
a
a
r
t
a
d
a
s
l
o
p
C
F
E
S
f
0
R
It must be noted that 3D printing is a promising alternative
echnology, and the considerable evolution over the last decade
ow makes multi-material 3D printing possible. However, the ini-
ial investment for 3D printers is high ($120,0 0 0), as is the case for
rinting flexible materials (approximately 300$/kg). Recently, Cloo-
an et al. [36] used this technology to manufacture AAAs with a
exible material. The development of the multi-material technol-
gy may allow anisotropic AAA phantoms to be printed, but to the
uthors’ knowledge it has not been done yet, possibly due to the
onsiderable effort and difficulties required to design the applica-
ion and set process parameters. The limitation in raw materials
or 3D printing is another drawback, which means there is a rel-
tively low variety in the mechanical properties of printed AAAs.
ther factors that may affect mechanical properties of the result-
ng AAAs are layering and multiple interfaces, which can cause im-
erfections in the phantom.
This study represents a step forward in manufacturing more re-
listic AAA models, presenting the development and application of
novel methodology for making AAA replicas with patient-specific,
egionally varying non-uniform wall thickness and anisotropic ma-
erial properties. However, one limitation should be noted: the
nisotropy attained in the AAA models is global, not local. This
rawback could be attenuated by increasing the number of fibers
nd reducing their size, but even using micro fibers would not re-
ult in a complete realistic in-vivo local anisotropy. Despite this
imitation, the global anisotropy yielded a more realistic behavior
f patient-specific AAA phantoms, which will have a positive im-
act on various clinical applications such as:
• The validation of numerical studies, medical image-based mod-
els and inverse characterization methods. • In vitro experiments (as an alternative to computational mod-
eling) for studying overall aneurysm mechanics coupled with
blood flow dynamics or for predicting rupture risk. • Pre-clinical testing of endovascular grafts where more realistic
in vitro models are needed. • Benchtop testing of endovascular grafts for the detection of
type III endoleaks.
• Experimental assessment of new or existing designs of catheter
devices in terms of trackability forces, the rigidity of catheter
guides, and the deployment of stent grafts.
ompeting interests
None declared.
unding
None.
thical approval
Not required.
upplementary materials
Supplementary material associated with this article can be
ound, in the online version, at doi:10.1016/j.medengphy.2016.09.
10 .
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