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Failure of vascular tissue with applications tothe aneurysm wall, carotid plaque and
myocardial tissue
Caroline Forsell
Doctoral thesis no. 83, 2013KTH School of Engineering Sciences
Department of Solid MechanicsRoyal Institute of TechnologySE-100 44 Stockholm Sweden
TRITA HFL-0545
ISSN 1104-6813
ISRN KTH/HFL/R–13/13–SE
ISBN 978-91-7501-798-3
Akademisk avhandling som med tillstand av Kungliga Tekniska Hogskolan i Stockholmframlagges till offentlig granskning for avlaggande av teknisk doktorsexamen fredag den 7Juni kl. 9.00 i D2, Kungliga Tekniska Hogskolan, Lindstedsvagen 5, Stockholm.
Somewhere, something incredible is waiting to be known.
Carl Sagan
i
Abstract
Cardiovascular disease is the leading cause of death in the modern world. Examples are tho-
racic aortic aneurysm (TAA), abdominal aortic aneurysm (AAA) and stroke due to plaque
rupture. Failure in soft tissues caused by medical devices is also a medical challenge. In
all these cardiovascular events a better prediction of failure of the tissue and a better under-
standing about the tissue properties will help in predicament and treatment. For example the
diameter-based indication for surgical repair of AAA and TAAs is not sufficient and refined
methods are needed. In this thesis failures of some soft vascular tissues, was studied. Ex-
periments have been combined with numerical modeling to understand the elastic and failure
properties of AAA, TAA and plaque tissue as well as the ventricular wall. Vascular tissue
is anisotropic, time-dependent, nonlinear and shows large deformations. Among others this
thesis showed the importance of viscoelasticity which motivates to develop a new continuum
mechanical framework. In addition a large part of this thesis dealt with anisotropy of vascu-
lar tissue. For the first time the collagen orientation distribution in the AAA wall has been
identified. Collagen and its distribution orientation is also an important feature of this tissue.
There was a correlation between the strength and stiffness of the AAA samples with the de-
creasing wall thickness. Increased stiffness was found in the aortic wall of patients with chronic
obstructive pulmonary disease (COPD) compared to patients that did not have COPD. As
well as difference in stiffness of TAA tissue, in patients with non-pathologic and pathologic
aortic valves. Some of the findings in this thesis could have a long-term consequence for
management of risk of rupture in AAA, TAA and plaque.
ii
Sammanfattning
Hjart- och karlsjukdomar ar den vanligaste dodsorsaken i det moderna samhallet. Ett par
exempel pa orsaker ar toracala aneurysm (TAA), buk-aneurysm (AAA) eller stroke pa grund
av plackruptur. Skador orsakade av medicintekniska produkter ar ocksa en medicinsk utman-
ing. Vid alla dessa kardiovaskulara sjukdomstillstand ar en battre forstaelse av skador pa
vavnaden och dess matrial egenskaper viktig for att hjalpa till vid diagnos och behandling.
Till exempel ar den numera diametern-baserade indikationen som anvans vid kirurgisks in-
grepp vid AAA och TAA inte tillracklig och battre metoder behovs. I denna avhandling
har skador pa nagra mjuka vavnader studerats. Experiment har kombinerats med numerisk
modellering for att forsta den elastiska och inelastiska responsen av AAA, TAA, plackvavnad
samt den ventrikulara hjartvaggen. Vaskular vavnad ar anisotrop, tidsberoende, icke-linjar
och karikteriseras av stora deformationer. I denna avhandling pavisades markant inverkan
av viskoelasticitet, som motiverade att en ny kontinuum mekanisk modell utvecklades. Aven
vikten av anisotropi undersoktes. For forsta gangen identifierades fordelningen av kollagenfi-
brer i AAA-vaggen. Vi fann att kollagenfibrerna och dess orientering ar en viktig egenskap
i denna vavnad. Det fanns en korrelation mellan styrkan och styvheten hos AAA-vavnaden,
med avseende pa minskad vaggtjocklek. Dessutom fann vi okad styvhet i AAA vaggen hos
patienter med kronisk obstruktiv lungsjukdom (KOL) jamfort med patienter som inte hade
KOL. Forutom detta fann vi en skillnaden i styvhet i TAA-vavnad, hos patienter med icke-
patologiska och patologiska aortaklaffar. Vissa resultat som pavisats i denna avhandling kan
fa en langsiktig relevans for hanteringen av risken for ruptur av AAA, TAA och plack.
iii
Preface
The work presented in this thesis has been carried out at the department of Solid Mechan-
ics, Royal Institute of Technology (KTH), Stockholm between 2008 and 2013. The work has
been financially supported by the Project Grants No. 2007− 4514 and 2010− 4446 from the
Swedish Research Council. The author is grateful for their financial support. First of all I
would like to express my sincere gratitude to my supervisor T. Christian Gasser for giving
me the opportunity to be a part of this project. I am thankful for all his support and encour-
agement and guidance.
During this project I have also have the opportunity to discuss and collaborate with sev-
eral inspiring people. Special thanks too Vincent M. Heiland, Xiao Xing, Joy Roy, Jesper
Swedenborg, Ulf Hedin, Sara Gallinetti, Per Eriksson, Anders Franco-Cereceda, Stanizlav
Polzer, Martin Oberg, Per Berg, Kurt Lindquist, Therese Olsson, Karin Lundstrommer and
Giampaolo Martufi.
I would also like to thank my colleagues and friends at KTH solid Mechanics for making this
a great place to work. In particular I would like to thank my room-mates Irene and Mats for
many interesting discussions. Last but not least I would like to thank my family, friends and
boyfriend Chris for their encouragement and love. Especially thanks to my mom Yvonne for
inspiring me to start with research.
Stockholm, May 2013
iv
List of appended papers
Paper A: Numerical simulation of the failure of ventricular tissue due to deep penetration:The impact of constitutive properties.C. Forsell and T.C. Gasser,Journal of Biomechanics 44 (2011) 45–51.
Paper B: The numerical implementation of invariant-based viscoelastic formulations at finitestrains. An anisotropic model for the passive myocardium.T.C Gasser and C. Forsell,Computer Methods in Applied Mechanics and Engineering 200 (2011) 3637–3645.
Paper C: Spatial orientation of collagen fibers in the Abdominal Aortic Aneurysm’s walland its relation to wall mechanics.T.C. Gasser, S. Gallinetti, X. Xing, C. Forsell, J. Swedenborg, and J. Roy,Acta Biomaterialia Issue 8 (2012) 3091–3103.
Paper D: The quasi-static failure properties of the Abdominal Aortic Aneurysm wall esti-mated by a mixed experimental-numerical approach.C. Forsell, J. Swedenborg, J. Roy and T.C Gasser,Annals of Biomedical Engineering, published online (2012).
Paper E: Identification of carotid plaque tissue properties using an experimental-numericalapproach.V.M. Heiland, C. Forsell, J. Roy, U. Hedin and T.C. Gasser,Journal of the Mechanical Behavior of Biomedical Materials, accepted for publication (2013)
Paper F: Failure properties for the thoracic aneurysm wall; Differences between BicuspidAortic Valve (BAV) and Tricuspid Aortic Valve (TAV) patients.C. Forsell, P. Eriksson, A. Franco-Cereceda and T.C. Gasser,Report 544, Department of Solid Mechanics, Royal Institute of Technology, Stockholm,To be submitted
v
In addition to the appended papers, the work has resulted in the following papers:
Automatic identification and validation of planar collagen organization in vascularwall from histological stains with an application to abdominal aortic aneurysmwall.S. Polzer, T. C. Gasser, C. Forsell, H. Druckmllerova, M. Tichy, R. Staffa, R. Vlachovsky,and J. Bursa,Submitted to Microscopy and Microanalysis.
In addition to the appended papers, the work has resulted in the following conference contri-butions:
Numerical Simulation of Deep Penetration of Ventricular Tissue.C. Forsell, T.C. Gasser, P. Gudmundson and G. Dohr,WCCM8 and ECCOMAS,Venice, Italy, June 30 - July 5, 2008.
Modeling of vascular failure with application to myocardial failure due to deeppenetration.C. Forsell and T.C. Gasser,10th US National Congress on Computational Mechanics, Columbus, Ohio, US, July 16-19,2009.
Modeling of myocardial splitting due to deep penetration.C. Forsell and T.C. Gasser,ECCMR 2009 - 6th European Conference on Constitutive Models for Rubber, Dresden, Ger-many, 7 - 10 September 2009.
Experimental and modeling of myocardial splitting.C. Forsell and T.C. Gasser,ECCM 2010 - IV European Conference on Computational Mechanics, Paris, France, May16-21, 2010.
The impact of constitutive properties on myocardial tissue perforation.C. Forsell and T.C. Gasser,WCB 2010 - 6th World Congress on Biomechanics, Singapore, Singapore, August 1-6, 2010.
Myocardial tissue deformation due to pacemaker lead contact-The impact of ma-terial anisotropy.C. Forsell and T.C. Gasser,CMBE2011-2nd International Conference on Mathematical Computational BiomechanicalEngineering, Washington D.C, USA, March 30-April 1, 2011.
Impact of material anisotropy on deformation of myocardial tissue due to pace-maker electrodes.C. Forsell and T.C. Gasser,SBC2011-Summer Bioengineering Conference, Farmington, Pennsylvania, USA, June 22-25,2011.
vi
A mixed experimental-numerical approach to identify the failure of the Abdom-inal Aortic Aneurysm wall.C. Forsell and T.C. Gasser,Euromech colloquium 534 Advanced experimental approaches and inverse problems in tissuebiomechanics, Saint- tienne, France May 29-31, 2012.
Mixed experimental numerical approach to identify properties of vascular tissue.C. Forsell and T.C. Gasser,WCCM 2012 - 10th Word Congress on Computational Mechanics, Sao Paulo, Brazil, July8-13, 2012.
vii
viii
Contents
Abstract ii
Sammanfattning iii
Preface iv
List of appended papers v
Introduction 1
Some cardiovascular challenges . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1
Risk assessment . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4
Aneurysm rupture . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4
Plaque rupture . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4
Pacemaker leads perforation . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4
Histology of vascular tissue and pathophysiology . . . . . . . . . . . . . . . . . . . 5
Aorta and collagen . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5
Plaque . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6
Ventricular wall . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6
Principles of vascular testing . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7
Experiments . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7
Constitutive models for vascular tissue . . . . . . . . . . . . . . . . . . . . . . . . . 10
Time-dependent properties . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 10
Failure through penetration . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 10
Prediction of failure and damage models . . . . . . . . . . . . . . . . . . . . . . . . 11
Aim 13
Method 15
Experimental study design . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 15
Numerical modeling . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 15
ix
Key Results 17
Discussion/Conclusion 21
Summary of appended papers 23
Bibliography 25
x
Introduction
Some cardiovascular challenges
Cardiovascular disease is the leading cause of death responsible for 17.3 million deaths world-
wide in 2008, of those around 7.3 million from coronary heart disease and 6.2 million died from
stroke [Mendis et al., 2011]. Specially, thoracic aortic aneurysm (TAA) and abdominal aortic
aneurysm (AAA) are diseases that require clinical attention. An aneurysm is a dilatation of
an artery and, due to the risk of rupture, it can become a life-threatening condition. Aortic
aneurysm rupture is the 16th leading cause of death for persons aged over 55 years and
ranked 19th for all age groups in US [CDC, 2010]. Another severe cardiovascular condition is
plaque rupture, especially if the rupture occurs in one of the major arteries [Richardson, 2002].
Carotid plaques (stenosis degree larger than 50%) are also, present in 29.2% of persons effected
by stroke [Kuroda et al., 2000].
Plaque consist of different components, where the proportion of these components influence
the properties of plaque, see the schematic picture Figure 1. Many plaques consist of a lipid-
rich necrotic core that is surrounded by an extracellular matrix (ECM) and is separated from
the blood stream by a fibrous cap. Plaque parts can travel in the arteries to the heart or the
brain leading to a stroke or a heart attack. Failure of soft tissues caused by medical devices
is another medical challenge. Specifically lead perforation is a rare but serious complication
of pulse generator implantation.
A pulse generator (defibrillator or pacemaker) is a device that is used to regulate cardiac
arrhythmias [Medical, 2006]. Figure 2 shows how the lead is placed in the heart. Prevalence
for lead perforation is about, 4.8% [Matsuura et al., 1990] for pacemaker leads and 5.2%
for [Molina, 1996] implantable defibrillator leads. Depending on the time from the operation
to the event, it can be acute, sub-acute, or delayed [Alla et al., 2010].
1
Extracellular Matrix
Fibrous capLipid core
Lumen
Artery
Figure 1: A cross section of an idealized plaque, illustrating the different (pseudo) tissues.
Figure 2: The heart consists of four chambers, and the figure illustrates how the pacemaker leads are placedin the heart [St.Jude, 2008].
2
Table 1: Abbreviations and expressions used in this thesis
AA Abdominal AortaAAA Abdominal Aortic Aneurysm
Anisotropy/ Orthotropic Property that depends on the direction in alldirections/some orthogonal directions
BAV Bicuspid Aortic ValveCEA Carotid Endarterectomy
Constitutive model Relation between stress and strainCOPD Chronic Obstructive Pulmonary DiseaseECM Extracellular Matrix
Hyperelastic The stress-strain relationship can be described witha strain energy function
Incompressible No volume change during deformationNon-linear material A material with non-linear stress-strain relationship
Stiffness The resistance offered by an elastic body to deformationStrain energy function The potential energy stored in a body due to a
deformation. It is equal to the work that must be performedto produce this deformation.
Strength Materials ability to withstand stress without failureTA Thoracic Aorta
TAA Thoracic Aortic AneurysmTAV Tricuspid aortic valve
Toughness Material ability to absorb energy and plastically deform withoutfracturing
UTS Ultimate Tensile StressWSS Wall Shear Stress
3
Risk assessment
Risk assessment is a critical step in the clinical decision making process, in determining the
patients treatment. Typically a clinical intervention is indicated if the risk from the clinical
condition exceeds those from the treatment.
Aneurysm rupture
Risk of aortic aneurysm rupture is related to its maximum diameter. Different diameter
thresholds are used depending where along the aorta the aneurysm is located, see Fig-
ure 3, for definitions of aortic segments. TAA repair is indicated if an aneurysm exceeds
5.5 cm for the ascending aorta, and 6.5 cm for the descending [Elefteriades, 2002]. For
AAA a large clinical study suggested that repair is indicated for a diameter that exceeds
5.5 cm [The UK Small Aneurysm Trial Participants, 1998]. Rupture risk is also related to the
growth rate, which depends on the diameter and other factors. Small aneurysms (less than 5.5
cm) growth in average by 4.4%, whereas the maximum growth reads 16% [Martufi et al., 2013].
Plaque rupture
Carotid endarterectomy (CEA) is a surgical procedure to prevent strokes. This procedure is
indicated for patients with carotid territory transient, due to rupture of vulnerable plaque,
ischemic attacks or a minor stroke with stenosis of 70 − 99%. For patients with 50 − 69%
stenosis a modest favoring for surgery has been found. For patents with less than 50% stenosis
anti-aggregants are recommended. [Biller and Thies, 2000]. The different compositions of
plaque affect the risk of rupture. For example a thin fibrous cap has been more commonly
observed in patients that died from acute myocardial infarction [Virmani et al., 2006].
Pacemaker leads perforation
Different factors have been reported to influence lead perforation [Carlson et al., 2008]. Con-
comitant therapies such as steroids and anticoagulants, implant techniques, and design char-
acteristics of the lead seem to have an influence [Carlson et al., 2008]. However currently it
is unclear why some leads perforate whereas others do not. To prevent lead perforation a
perforation-resistant lead design would be preferable, which in turn requires an understanding
of the mechanical failure mechanisms of the ventricular wall.
4
Figure 3: Aorta and its different parts [Isselbacher, 2005]
Histology of vascular tissue and pathophysiology
Aorta and collagen
Blood vessels consist of three layers, intima, media and adventitia. The intima contains en-
dothelial cells and its structural contribution in healthy subjects seems negligible. The media
is an active vascular layer that contains smooth muscle cells, collagen, elastin and ground
substance. From a mechanical perspective the adventitia mainly contains collagen fibers.
For the thoracic aorta collagen contents of 36.8% and 77.7% for media and adventitia, are
reported [Fung, 1981]. While the amount of collagen critically influences the mechanical be-
havior, the orientation of the collagen is also an important structural wall property. Previous
studies found that in large AAAs the collagen content increase by about 50%, while elastin
almost disappears when compared to the normal aorta [Rizzo et al., 2011]. Collagen is bire-
fringent, a property that can be used to measure the collagen orientation in the wall by
5
polarized light microscopy. The organization of collagen has been reported for intracranial
aneurysms [Canham et al., 1999] coronary arteries [Canham et al., 1989] and AAAs
[Gasser et al., 2012]. The AAA wall collagen fiber orientation was found to be orthotropic and
no difference between the medial and the adventitial layer was seen. A larger orientation vari-
ance has been found in the tangential plane than in the cross-sectional plane [Gasser et al., 2012].
In coronary arteries a main direction, of the fibers, have been found in the circumferential
direction [Canham et al., 1989]. Other methods like using small-angle light scattering have
also been suggested to study the collagen orientation in pericardial tissue [Sacks et al., 1997].
Unfortunately not many studies combine tensile testing and histology measurements. One
rare example is the study that used multi-photon microscope combined with tensile test-
ing [Hill et al., 2012].
Plaque
Plaque consist of different components, where the proportion of these components change in
different plaques [Kim et al., 2008]. Stiffness can vary significantly for the different (pseudo)
tissues. For fibrous cap tissue from human plaques a Young’s modulus of 500 kPa and for lipid
core of 5 kPa was reported [Ohayon et al., 2007]. In addition residual strain in the fibrous
cap seem to increase with increased size of the lipid core [Ohayon et al., 2007]. Another
approach reports cross-plaque properties [Loree et al., 1994] and divides it into hard, mixed
and soft plaques. Based on a Yeoh model [Yeoh, 1993], different material parameters have
been reported for these different groups of plaques [Lawlor et al., 2011]. Calcification can
occur in late stages of plaques, creating a more complex structure with altered mechanical
properties [Kim et al., 2008]. The effect of calcification on plaques instability is not yet clearly
understood [Virmani et al., 2006].
Ventricular wall
The ventricular wall consists of three layers; endocardium, myocardium and epicardium
[Bloom and Fawcett, 1970], from the inside to the outside. The thickest layer is the my-
ocardium that consist primary of fibroblasts and myocytes. The myocyte-fiber direction
changes across the thickness of the myocardial wall [Okamoto et al., 2000].
6
Principles of vascular testing
Several different test methods have been suggested to investigate mechanical properties of
vascular tissue in the literature. A brief summary of the tests methods is discussed below.
Common test methods are uniaxial tensile testing (Figure 4(b)) and uniaxial compression
testing (Figure 4(a)). To study shear deformation in biological soft tissue function, shear
testing can be performed (Figure 4(c)). To better mimic in-vitro loading, biaxial testing
(Figure 4(l)) is recommended. In-vivo vascular tissue is inflated such that inflation testing
(Figure 4(f), Figure 4(g)) is motivated. Splitting tests (Figure 4(h)) is also preformed where
a tear causing a cavity of fluid is formed in the tissue by injecting saline solution. In order
to study aortic dissection, tearing tests (Figure 4(d)) and pealing tests (Figure 4(e)) have
been used. In the literature how the tissue respond to penetration or bending has been
investigated by penetration testing (Figure 4(k)), probing tests (Figure 4(i)) and bending
tests (Figure 4(j)).
Experiments
Uniaxial tensile tests of aortic tissue have been performed in axial, circumferential
[Mohan and Melvin, 1982, Vorp et al., 1996, Vorp et al., 2003] and in radial
[MacLean et al., 1999, Sommer et al., 2008] directions. Results from these tests are repre-
sented for the human and porcine Thoracic Aorta (TA), TAA, Abdominal Aorta (AA) and
AAA. This data suggest a strength for aneurysm tissue (AAA and TAA) in axial direction
in the range of 0.65 − 1.21 MPa and circumferential direction 0.68 − 1.47 MPa. Compared
to normal aortic (TA and AA) with tissue strength of 1.47 − 1.71 MPa, 1.72 − 1.80 MPa,
0.061− 0.140 MPa in axial, circumferential and radial directions respectively. Stiffness have
been reported in a range of 3.51−4.48 MPa in axial direction and 4.76−5.81 MPa in circumfer-
ential direction, in aneurysm tissue (AAA and TAA). For non-aneurysmal tissue (TA and AA)
in the range of 2.61 − 5.69 MPa for longitudinal and circumferential direction have been re-
ported. Planar biaxial tensile testing has been used to determine parameters for a constitutive
relation of ventricular wall [Humphrey et al., 1990b, Humphrey et al., 1990a]. Shear testing
of the passive left ventricular myocardium suggest that the response to simple shear defor-
mation is dominated by the tensile material properties [Dokos et al., 2002]. Splitting tests of
the porcine tissue [Carson and Roach, 1990, Roach and Song, 1994, Roach et al., 1999] and
human tissue [Tiessen and Roach, 1993] is performed in thoracic and abdominal aorta respec-
tively. The splitting energy was found to be in a range of 1.88−15.9 mJ/cm2. Peeling testing
in AA [Sommer et al., 2008] found a dissection energy of 5.1 mJ/cm2. Tearing tests showed
that the fracture toughness of the aorta increases the further away the tissue was from the
heart [Purslow, 1983]. The mean breaking stress found in porcine TA was 2.19 − 3.64 MPa,
0.18 − 0.87 MPa in circumferential and axial directions, respectively. Finally inflation tests
7
done in TA [Mohan and Melvin, 1983, Groenink et al., 1999] showing a ultimate stress of
0.114− 2.7 MPa can be found in literature.
Uniaxial tensile testing of plaque tissue in circumferential direction
[Lawlor et al., 2011, Maher et al., 2009] and compression test in radial direction by
[Maher et al., 2009] can be found in the literature. The results of the Ultimate Tensile
Stress (UTS) varied between 0.131 − 0.779 MPa at a UTS strain of 0.299 − 0.588. Uni-
axial tensile testing has also been used to investigate viscous properties in human plaque
by relaxation testing [Salunke et al., 2001]. The small dimensions and irregularities in the
shape of plaque samples motivates probing at different sites to investigate compression re-
sponse [Tracqui et al., 2011]. Penetration tests have been performed in myocardial tissue
[Gasser et al., 2009]. The penetration pressure (penetration force/punch cross section area)
decrease slightly with larger punch diameter and has been found to be 1.76 MPa and 2.27 MPa
for diameters of 1.32 mm and 2.30 mm.
8
(a) (b) (c) (d)
(e)
P
(f)
P
(g)
P
(h) (i) (j)
(k) (l)
Figure 4: Testing principles to explore mechanical properties of vascular tissue: a)Uniaxial compression b)Uniaxial tensile c) Shear d) tearing e) Peeling f) Inflation g) Inflation h) splitting i) Penetration j)Bending k) Probing l) Biaxial
9
Constitutive models for vascular tissue
The literature on constitutive modeling of vascular tissue is rich [Fung, 1981]. Soft biological
tissue is non-linear elastic, anisotropic and incompressible material, which often is described
by hyperelastic theory. The incompressibility under many loading conditions is a consequence
of a large amount of non-mobile fluid, while macroscopic anisotropy is a consequence of the
organized structure of vascular tissue. Non-linearity arises from the engagement of fibrous
compounds under loading.
Originally, isotropic non-linear constitutive models have been developed for rubber mate-
rials. The simplest model is the NeoHookean [Mooney, 1940]. Others are the Mooney -
Rivling [Rivlin, 1948], Ogden, [Ogden, 1972], Yeoh [Yeoh, 1993] and others found in litera-
ture [Demiray, 1972]. However these models are not describing the anisotropy of the tissue
and hence have limited applicability in vascular mechanics. Anisotropic formulation for vas-
cular tissue are found as phenomenological [Humphrey, 1995] and histo-mechanical
[Gasser et al., 2006, Lanir, 1983, Holzapfel et al., 2000, Martufi and Gasser, 2011] approaches.
Time-dependent properties
Predicting the failure of soft biological tissue requires inelastic constitutive descriptions, where
specifically viscoelastic dissipative effects are important. Commonly used models to describe
viscoelasticity are the Voigt, Maxwell, and the standard linear solid models. For vascular
tissue there is a difference in stress for loading and unloading. It is however reasonable to
assume linear theory for oscillations of small amplitude about an equilibrium state. Nonlinear
generalizations of the standard linear solid model can be found in literature [Viidik, 1968].
A nonlinear differential equation can be used to describe the damper in the Maxwell model
or a quasi-linear solid can be considered where the non-linear viscoelasticity is modeled by
a series of nonlinear elements i.e. Maxwell-elements in parallel. This to capture a large
range of frequencies [Holzapfel et al., 2002, Gasser and Forsell, 2011, Puso and Weiss, 1998,
Idesman et al., 2001, Wu et al., 2006].
Failure through penetration
The application of an isotropic material assumption to tissue penetration found that a flat-
bottom punch leads to a mode-II ring while using the sharp tip formed a planar mode I
crack [Shergold and Fleck, 2004]. In contrast in-vitro experiments on heart tissue, which are
anisotropic, showed that even a flat bottom punch as well as a sharp tip punch created a
mode I failure [Gasser et al., 2009].
10
Prediction of failure and damage models
Failure in vascular tissues is defined by setting up a failure criteria [Volokh, 2011]. Examples
of failure criteria are maximum Von Mises stress, the maximum principal stress, the maximum
strain energy, etc... In the literature damage parameter are commonly used
[Hokanson and Yazdani, 1997, Rodrıguez et al., 2006, Balzani et al., 2006]. It can be imple-
mented on a mechanical level, a finite element level, or a continuum level.
11
12
Aim
This work aimed to advance our understanding in failure of vascular tissues, specifically the
AAA, TAA, the heart wall, and carotid plaques were studied. Constitutive material pa-
rameters were identified using numerical methods in combination with tailored experimental
testing.
Specific aims were:
• To estimate elastic and inelastic properties of the AAA wall.
• To investigate the difference in structural and mechanical properties of the TAA wall from
patients with normal and pathological aortic valves.
• To measure the three-dimensional structure of collagen in AAA and integrate it in a histo-
mechanical constitutive model.
• To identify tissue properties from carotid plaque, using bending testing.
• To study numerically pacemaker lead perforation of myocardium.
• To implement an improved viscoelastic formulation for a class of invariant-based constitu-
tive formulations.
13
14
Method
Experimental study design
Experimental design and results are detailed in the appended papers. In summary, uniaxial
failure tests have been performed with AAA and TAA tissue. The three-dimensional struc-
tural arrangement of collagen in the AAA and TAA wall, was measured from histological
slices using polarized light microscopy.
Dynamic bending testing with plaque samples was used to investigate the material properties
of fibrous tissue components. All tests where displacement controlled and performed in 37 ◦C
salt solution. The test specimens where directly collected from surgery at Karolinska Hospital
and testing was and preformed within one day.
Numerical modeling
The FE packages FEAP (University of California at Berkley) and ABAQUS (Dassault Sys-
tems) were used for modeling. The aneurysm wall was represented by the elasto-plastic
damage model, which allowed models to include the measured collagen orientation distribu-
tion. For the plaque tissue the thrombus and the lipid region were described by a NeoHooken
model with parameters from literature. For the fibrous plaque tissue, the isotropic version
of the GOH-model [Gasser et al., 2006] was used. Viscoelastic effects were captured by five
Maxwell devises. Finally, to study the stress field from contact with the pacemaker lead a
visco-elastic model for myocardial tissue was implemented in ABAQUS.
15
16
Key Results
One key finding was that rate-dependent effects of the bulk material strongly determine the
failure process, whereas dissipative effects directly related to failure zone did not have signif-
icant impact on the simulation results. Compare Figure 5a and Figure 5b, which illustrate
the limited influence of the fracture energy when contributed to changing viscosity-related
parameters.
−1 0 1 2 3 40
0.05
0.1
0.15
0.2
0.25
Displacement [mm]
Fo
rce
[N
]
Fracture Set I
Fracture Set II
−1 0 1 2 3 40
0.05
0.1
0.15
0.2
0.25
0.3
0.35
Displacement [mm]
Fo
rce
[N
]
Prony Set I
Prony Set II
−1 0 1 2 3 40
0.05
0.1
0.15
0.2
0.25
0.3
0.35
Displacement [mm]
Fo
rce
[N
]
Elastic Set I
Elastic Set II
0 0.2 0.4 0.6 0.8 10
0.1
0.2
0.3
0.4
0.5
Relative Thickness
Fo
rce
[N
]
Simulation
Experiment Interventricular Septum
Elastic Set IProny Set I
Elastic Set IFracture Set II
(b)(a)
(c) (d)
Prony Set IIFracture Set II
Figure 5: Penetration force displacement response of myocardium against deep penetration. Impact of thefracture energy (a), the visco-elasticity (b) and the elastic properties (c) on the simulated results.(d) Data from experimental bovine myocardium penetration tests. [Forsell and Gasser, 2011]
17
Figure 6 illustrates the elastic response of a punch pushing down on a disk of porcine
myocardial tissue. The difference in the punch force, for isotropic and anisotropic formula-
tions, can clearly be seen. These results underline the importance of anisotropic constitutive
modeling in biomechanics. The softer cross fiber direction accumulates more strain then the
fiber direction. This deformation mechanism is not possible an the isotropic material, which
might be a reason for this observation.
1 1.05 1.1 1.15 1.2 1.250
0.005
0.01
0.015
0.02
0.025
0.03
0.035
Stretch
Pun
ch F
orce
[N]
Anisotropic modelIsotropic model
Figure 6: The elastic response of a penetrator, pushing down on a disk of porcine myocardial tissue. Punchforce displacement response using the anisotropic (solid line) and isotropic (dashed line) constitutivedescriptions for myocardial tissue. [Gasser and Forsell, 2011]
The collagen dispersion was much larger in the tangential plane than in the cross-sectional
plane for the AAA wall. The adventitial and medial layers where found to have no significant
different in collagen orientation (see Figure 7) for polarized light microscopy images. This is
to the author´s best knowledge the first time that a constitutive model based on structural
properties predict the mechanical properties of AAA wall. So far only theoretical assumptions
are represented in the literature without experimental validation.
Another key result was that the strength and stiffness of the AAA wall increased with de-
creasing wall thickness, but no correlation between strength and aneurysm diameter was seen.
We also found that the AAA wall stiffness and strength increased in patients with chronic
obstructive pulmonary disease (COPD) compared to patients that did not have COPD. The
difference in strength for the two patient groups was statistical significant (p = 0.0290) and
can be seen in Figure 8. To our best knowledge this influence on strength on COPD has
never been reported. The collagen fibers of the COPD patients are thicker and that might be
one of the explanations for the difference among these patients. Alternatively an increase in
collagen content in the AAA tissue, for the patients with COPD, could explain this difference
in strength, in the AAA wall.
18
Figure 7: Polarized Light Microscopy (PLM) images taken from the media (left) and the adventitia (right) ofthree Abdominal Aortic Aneurysm (AAA) wall samples (a)-(c). The horizontal sides of the imagesdenote the circumferential direction. Picrosirius red was used as a birefringent enhancement stainand the images were taken at crossed polar on the microscope. Quantitative collagen orientationdata of the presented samples is given in Figure 4. (a) Typically observed collagen organizationsin the AAA wall, where the media and adventitia shows a very mixed bag of azimuthal alignment.(b) Rather circumferential alignment of collagen fibers that was seen in a few AAA wall samples.(c) Collagen organization that remembers on a normal aorta and seen in two AAA wall samples,i.e. where the adventitia shows two almost perpendicular families of collagen. [Gasser et al., 2012]
19
0.2
0.4
0.6
0.8
1.0
1.2
without CPOD Hn=10L with CPOD Hn=5L
Wa
llstr
en
gth
MP
aFigure 8: Strength of Abdominal Aortic Aneurysm (AAA) wall specimens, that were taken from patients
with and without COPD. The number of specimens is denoted by n, and the difference betweenboth groups is statistical significant (p=0.0290). The thick solid lines are the medians. The barsand boxes denote the standard deviations and the 50% quartile, respectively. [Forsell et al., 2012]
Similarly a difference between the stiffness and strength of TAA tissue, for patients with
normal and pathologic aortic valve was found. In addition to these key findings material
behavior was identified for the different vascular tissues and novel in-vitro mechanical testing
protocols was developed, see the appended papers for details.
20
Discussion/Conclusion
Computer simulations can help with a detailed patient specific analysis and help to predict
and diagnose diseases. Today we are close to a bioengineering knowledge when computational
simulations could be of big help for the clinicians. Due to variability in input parameters such
as loading conditions and constitutive properties models need further development. To get a
better understanding this diversity a mixed numerical approach was used to study the elastic
an inelastic properties of AAA, TAA, heart wall and carotid plaque tissue. This allowed
uncovering the mechanical properties of some vascular tissues. Some of these functions can
have long-term consequences for patient management.
Other approaches for the work in this thesis would be to use a more clinical statistical anal-
ysis and to follow up patients. The author thinks it is important to use new aproaches for
this problem to be able to get more knowledge in this important field. When doing me-
chanical testing in soft tissues it is preferable to mimic the in-vitro conditions. This is a
big challenge and we usually do not know the original stress state. To get an in-vivo load-
ing condition of the anisotropic vascular wall, and to characterize its inherent pseudo-elastic
properties [Humphrey, 2002] planar biaxial testing would be needed. Since we were primar-
ily interested in failure this would require a large number of specimens too get a reasonable
range of biaxial loading states. Consequently we used uniaxial tensile tests in this thesis. Me-
chanical testing has a large variability, and due to the limited number of samples, there are
statistical uncertainties in the conclusions of this thesis. One of our constitutive assumptions
was that collagen fibers are the main load carrier in TAA wall and AAA wall. This is found
to be reasonable for larger AAAs [Rizzo et al., 2011].
21
22
Summary of appended papers
Paper A: Numerical simulation of the failure of ventricular tissue due to deep penetration:
The impact of constitutive properties In this paper the myocardial failure due to deep pene-
tration was investigating. The impact of different constitutive parameters on the simulated
results was controlled. A phenomenological model described the failure with a traction sepa-
ration law and the finite element model considered the non-linear, isotropic and visco-elastic
properties of the myocardium. A non-linear FE modeling integrated the information from
tensile testing in cross-fibre with constitutive data from literature. The result showed the
importance of a viscoelastic formulation.
Paper B: On the numerical implementation of invariant-based viscoelastic formulations at
finite strains. A model for the passive myocardium. A framework for finite strain viscoelas-
ticity was developed and implemented. A superposition of an Elastic Body and Maxwell
Body where the reference configuration of the Maxwell Body moves in space and stores the
history of the deformation with a rate equation in strain space. The current configuration of
the continuum coincides with the reference configuration of the Maxwell body at the thermo
dynamical equilibrium. Two independent strain variables are describing the Helmholtz free
energy. An update equation of the over-stress allows variable time step to be used. The
model was implemented and the material behavior due to a rigid punch into myocardium
was studied. Specially the importance of anisotropy was looked at and the result shows the
importance of an anisotropic model.
Paper C: Spatial orientation of collagen fibers in the abdominal aortic aneurysm’s wall and
its relation to wall mechanics Collagen and its distribution orientation is important for the
mechanical properties of the wall. In this paper samples from AAA was collected at elective
AAA repair from Karolinska hospital. Tissue samples where fixated, embedded, sectioned,
stained and investigated by polarized light microscopy. The 3D orientation of the collagen
was measured and captured by a Bingham distribution function. This was performed to in-
tegrate the identified structural information in the AAA wall. No significant difference of the
orientation of collagen, between media and adventitia lager of the wall were found and the
dispersion was found lager in the tangential plane than in the cross-sectional plane.
23
Paper D: The quasi-static failure properties of the Abdominal Aortic Aneurysm wall esti-
mated by a mixed experimental-numerical approach. The present study estimated the elastic
and inelastic properties of the AAA wall. A mixed experimental-numerical approach was
used. Data from collagen orientation and results from failure tests of AAA wall was com-
bined into finite elements modeling. A histo-mechanical constitutive model was used were the
weakening of the fiber and remaining deformation were described by plastic fibril sliding. The
finite element models simulated the tensile tests, in longitudinal direction, for 16 specimens.
The stiffness and strength of the AAA tissue were found to be higher for patients with COPD
then for non COPD patients. No correlation with gender or, smoking were found to effect
the parameters related to irreversible deformation response.
Paper E: Identification of carotid plaque tissue properties using an experimental-numerical
approach The present study introduces an inverse parameter estimation approach to extract
tissue properties from fibrous cap of carotid plaque. Samples from Carotid Endarterectomy
(CEA) were used for in-vitro force displacement testing. Histology images were used to de-
velop Finite element- (FE-) models. An optimization procedure was then used to identify
material parameters and solve the inverse problem presented in the study.
Paper F: Failure properties for the thoracic aneurysm wall; Differences between Bicuspid
Aortic Valve (BAV) and Tricuspid Aortic Valve (TAV) patients. Aorta is predisposed for
hemodynamic events. Specially patients with aortic valve pathogeneses have an increased
risk of TAA. This study is investigating the mechanical properties of the TAA wall using
uniaxial failure testing and measurement of collagen orientation. The 3D orientation of col-
lagen is measured with help of polarized light microscopy. A FEM models integrate the 3D
collagen information with the tensile properties to identify material parameters for elasto-
plastic damage model. Especially the mechanical differences for the aneurysm tissue between
BAV and TAV patients were looked at. A statistical significant difference where found for
the TAA tissue in the ultimate tensile stress σult, between patients with TAV and BAV
(1.496SD0.845 MPa, 0.723SD0.336 MPa,p = 0.02). Also the collagen fiber stiffness-related
parameter (p = 0.01) had a statistical significant difference.
24
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