OF PULMONARY ARTERY FOR PULMONARY HYPERTENSION By Yuheng Wang A THESIS Submitted to Michigan State University in partial fulfillment of the requirements for the degree of Mechanica l Engineering Master of Science 2019 ABSTRACT OF PULMONARY ARTERY FOR PULMONARY HYPERTENSION By Yuheng Wang Pulmonary hypertension (PH) is associated with elevated pulmonary arterial pressure. PH prognosis remains poor with 15% mortality rate within 1 year even with modern clinical managements. Previous clini cal studies proposed the wall shear stress (WSS) to be an important hemodynamic factor for affecting cell mechanotransduction and growth and r emodeling in the disease progress. However, a typical range of WSS in vivo is at most 2.5 Pa and a doubt has been casted whether WSS alone can influence the disease pro gress. Furthermore, our current understanding of PH pathology has largely been obtained through small animals and there has been seldom reports of caliber enlargement in the PH animal models. Therefore, a large - animal experiment on pulmonary arteries (PAs) is needed for validating whether an increased pressure can induce an enlargement of pul monary caliber. In this study, we use an inflation testing device to characterize the mechanical behavior, both nonlinear elastic behavior and irreversible damage of po rcine arteries. The parameters of elastic behavior are estimated from the inflation test at a low - pressure range first and then are compared with those from a high - pressure range, which tests if those behavior are significantly different. At the end of mec hanical tests, histological images are qualitatively examined for medial and adventitial layers. This study, therefore, sheds light on the relevance of pressure - induced damage mechanism in human PH iii ACKNOWLEDGEMENTS During my M.S. program, I met many peopl e who helped and encouraged me. First, I would like to thank my primary advisor Dr. Sara Roccabianca who gives me a great opportunity to do the research in her group as master student. I am very grateful to her encouragement, inspiration and knowledge supp ort through my entire master program. I would also like to thank my committee member Dr. Seungik Baek and Dr. Lik Chuan Lee for their constructive guidance and valuable feedback. I also appreciate all the members from our research group Dr. Sheng Chen, Dr . Hamidreza Gharahi , Marissa Rae Grobbel , Mayank Sinha , Tyler Tuttle and Laura Alison Ny e . They have contributed a lot to my research. As co - workers, they provide d a lot of technic supports and suggestions while doing the experiment. Finally, special thanks should belong to my lovely family and friends for their unconditional supports and enc ouragemen t iv TABLE OF CONTENTS LIST OF TABLES ................................ ................................ ................................ .......................... v LIST OF FIGURES ................................ ................................ ................................ ....................... vi CHAPTER 1: Int roduction and liter ature review ................................ ................................ ........... 1 1.1 Overview ................................ ................................ ................................ ................................ ... 1 1.2 Arterial mechanics and its role in the progression of cardiovascular pathology ...................... 2 1.3 Pulmonary artery microstructural organization in healthy and PH ................................ .......... 6 1.4 Pulmonary artery constitutive model ................................ ................................ ...................... 10 1.5 Mechanical test of elastic vessels in vivo and in vitro ................................ ............................ 15 C HAPTER 2 : Mechanical characterization of damage behavior in the pulmonary artery .......... 17 2.1 Introduction ................................ ................................ ................................ ............................. 17 2.2 Methods ................................ ................................ ................................ ................................ ... 17 2.2.1 Specimen preparation ................................ ................................ ................................ ... 17 2.2.2 Mechanical testing ................................ ................................ ................................ ....... 18 2.2.3 Histological analysis ................................ ................................ ................................ .... 19 2.2.4 Mechanical model ................................ ................................ ................................ ........ 19 2.3 Results ................................ ................................ ................................ ................................ ..... 21 2.4 Discussion ................................ ................................ ................................ ............................... 27 2.5 Conclusion ................................ ................................ ................................ .............................. 30 B IBLIOGRAPHY ................................ ................................ ................................ ......................... 32 v LIST OF TABL ES Table 1 . Best - fit material parameters for the 2 - fiber family model and estimated axial stretch .. 27 vi LIST OF FIGURES Figure 1. Sche matic of the layered microstructure of the arterial wall. (G. A. Holzapfel, Gasser, & Ogden, 2000) ................................ ................................ ................................ .............................. 6 Figure 2. Contour plot of the convex potential (1) with material parameters = 44.2 kPa a nd = 16.7 (see (D elfino et al., 1997) ) ................................ ................................ ................................ .... 11 Figure 3 . Contour plots of the potential (3) with (a) material parameters , in (Chuong & Fung, 1983) , and (b) a set of parameters chosen to illustrate non - convexity ........................... 13 Figure 4 . Inflation - extension testing device with specimen mounted in between ....................... 18 Figure 5 . Luminal pressure and outer diameter raw data for a representative sample, as collected throughout the mech anical test. Specifically, part 1 (dark gray line) describes the mechanical behavior of the PA before the damage, part 2 (light gray line) describes the response to over - pressurization which generates damage within the wall, and part 3 (dark gray line) des cribes the mechanical behavior of the PA after damage. Also shown, the preconditioning protocol (dotted light gray line). ................................ ................................ ................................ .............................. 22 Figure 6 . Average pressure vs. normalized diameter curve for all specimens. and correspond to before and after damage d ia meters, respectively. The before damage behaviors have been collected during the first loading of part 1 of the test, while the after damage behavior have been collected during part 3 of the test. The asterisk indicates that the normalized diameters were s ignificantly different, at p < 0:05, when comparing before and after damage behaviors .. 24 Figure 7 . Change in diameter during different loading curves. Black circles represent that diameter difference between the first and second loading curve of part 2 of the test, gray squares represent the difference in diameter between the second and third loading curve. The * indicates that the diameters were significantly different at p < 0:05. ................................ .......................... 25 Figure 8 . Histology images of tissue samples from the PA. Top row: Pi crosirius red under polarized light (collagen fibers in red, yellow and green); bottom row: VVG (elastin in black, nuclei in purple, and cytoplasm in pink). Left: before damage samples; right: after damage ................................ ................................ ............ 26 1 C HAPTER 1: Introduction and literature review 1.1 Overview Pulmonary hypertension (PH) is a type of high blood pressure that a ffects the right ventricle and the arteries in the lungs . PH can develop without a known cause (primary pulmonary hypertension) o r as a result of other disease s (secondary pulmonary hypertension). Secondary pulmonary hypertension is often associated with collagen vascular disease, chronic thromboembolism, human immunodeficiency virus, and other diseases. Prognosis is severe in PH an d when untr eated it is a potentially fatal disease that rapidly leads to disability and premature mortality (Stepnowska, - . Since 1999, the number of deaths and hospitalizations as well as death rates and hospitalization rates for PH have increased. Despite improvement s in the diagnosis and management of PH over the past 2 decades , with the introduction of targeted medical therapies leading to improved survival, the disease c ontinues to have a poor long - term prognosis (Mehari, Valle, & Gillum, 2014) . Further, the reversibility of PH is dependent on the relative contribution of vaso constriction (reversible) and structural changes in the pulmonary vessels (mostly irreversible) ( Rounds & Hill, 1984) . Furthermore, w hen patients have symptoms or signs of PH, the disease is usually at an advanced stage, at which the pulmonary vascula tu r e already shows remodel ing in terms of the wall microstructur e . Therefore, it is vital to determin e the bio - chemo - mechanical process es involved with the initiation and progression of microstructural remodeling associated with PH , in order to prevent irrevers ible outcomes . Furthermo re, a better understanding of the mechanisms that contributes to wall re modeling in PH could lead to innovative therapeutic targets. 2 1.2 Arterial mechanics and its role in the progression of cardiovascular pathology Arteries have, primarily, a mechanical function, which is influenced heavily mechanical characteris tics. Furthermore, it has been hypothesized that perturbations of the mechanical homeostatic state in arteries could be a driving force of disease pr ogression. For example, data suggests that pressure - induced stress concentrations may play an important rol e in the formation and progression of atherosclerotic plaques (Thubrikar & Robicsek, 1 995) . In addition, it is widely accepted that arterial stiffening plays an important role in hypertensio n progression (Mitchell et al., 2010). A large longitudinal study recently established that arterial stiff e n ing precedes an increase in systolic blood pressure , while initial blood pressure was not independently predictive of subsequent aortic stiffening ( Kaess et al., 2012). These studies highl ighted the importance of understanding the mechanism s underlying the initiation and progression of arterial stiffening in relation to hypertension, and of determining the temporal relationship between the development of arterial stiffness , high blood press ure, and cardiovascular disease (Weisbrod Rober t M. et al., 2013). Finally, many researchers suggested that arterial wall mechanical dysfunction is one of the crucial mechanisms in the formation of a bdominal aortic a neurysms (AAA). AAA are associated with atherosclerosis, aging, smoking, and hypertension, but the mechanism underlying the development of atherosclerotic aneurysms is not well understood (Hammond & Garfinkel, 1969). Previous studies observed that the form ation of at herosclerotic aneurysm is associated with dramatic changes in the extracellular matrix composition and organization of the arterial wall . Collagen concentration has been found to be deficient (R. J. Rizzo et al., 1989), unchanged (Dubick, Hunter , Perez - Liz ano, Mar, & Geokas, 1988), or increased (Menashi, Campa, Greenhalgh, & Powell, 1987) in atherosclerotic abdominal aortic aneurysms, whereas elastin concentration has been found to be consistently deficient in 3 aneurysms (Campa, Greenhalgh, & Powe ll, 1987; D ubick et al., 1988; Robert J Rizzo et al., 1989). From these studies it appears that an imbalance between synthesis and degradation of arterial connective tissue (collagen and elastin) may be responsible for the development of atherosclerotic an eurysms. P ulmonary hypertension: the biomechanical hypothesis. Several hypotheses have been made to explain the mechanisms that lead to arterial mechanical dysfunction that aggravate , or possibly trigger , PH . These are complicated by the multifactorial na ture of the disease , involving : (1) maladaptive pulmonary wall remodeling that leads to media hypertrophy , adventitial thickening and neointimal lesions (Botney, 1 999) ; ( 2) d egradation of molecular tissue components linked with age - related changes (Akhtar, Sherratt, Cruickshank, & Derby, 2011) ; and (3) mechanical damage mechanism triggered by abnormally large stress es exerted by hemodynamic force s (Zambrano et al., 2016 , Humphrey & Tellide s, 2018) . Although the initiating mechanism of PH is still unclear , i t is commonly accepted that hemodynamic conditions, especially wall shear stress (WSS) , influence cardiovascular disease by affecting cell mechano transduction and microstructural remodel ing. T here is increasing evidence , in the systemic vasculature , that low WSS is a promoter of increased wall stiffness and atherogenic vascular states, and could be an independent predictor of cardiovascular mortality (Cu nningham & Gotlieb, 2005) . Furthermore, recent studies propose d that WSS is the primary mechanical force affecting cell mechanotrasnduction in PH (Bürk et al., 2012) , causing inflammatory response and change in cell expression from contractile to proliferative (Li, Scott, Shandas, Stenmark, & Tan, 2009; Li, Stenmark, Shandas, & Tan, 2009) . N umerical studies suggest ed that a 30% increase in flow r ate produce s a 9% increase in arterial diameter (Valentín A, Cardamone L, Baek S, & Humphrey J.D, 2009) , however, it has been showed experimentally that the increase in caliber of human pulmonary artery (PA) in PH patients 4 is significantly larger than that . Specifically, p eople affected by PH present more than 50% larger pulmonary diameters when compared to healthy subjects in young patients (Truong et al. , 2013) , and 30 % in adult s patients (Edwards, Bull, & Coulden, 1998) . Clearly, d iameter enlargement is one of the most prominent feature of PH and the distribution of WSS was found to have substantial in fluence on both the diameter and the shape of the vessels in the lungs (Vorp, Wang, Webster, & Federspiel, 1998) , yet it s eems improbable that low WSS alone could be responsible for the signifi cant increase in arterial caliper in PH. Furthermore, previous studies also suggested that changes in the mechanical characteristics of the arterial wall could be involved in aggravatin g PH pathology. For example, Sanz and colleagues observed that , in exercise - induced PH, the pulmonary arter stiffness increase preceded the observation of an increase in both pressure and luminal diameter (Sanz et al., 2009) . The Authors of that study suggested that an interplay between variations of wall stiffness, luminal diameter, and wall shear stress , as opposed to one isolated cause, could be at the root of PH progression. In this study, we hypothesize that a damage mechanism could a lso contribute to irreversible diameter enlargement in PH . T o test this hypothesis, we combined in vitro experiments and theoretical modeling to investigate the role of damage in the mechanics of pulmonary artery . Damage mechanisms in elastic arteries. Lit tle is known about the response to damage of pulmonary arteries, however some studies focused on investigating damage mechanisms in other elastic arteries, both theoretically and experimentally . In 1987, one of the first attempts to model arterial wall dam age at large strains was pursed by Simo (Simo, 1987) . The author proposed a nonlinear viscoelastic constitutive model , capable of accommodating general anisotropic response and relaxation functions. T his viscoelastic model successfully predicted the progressiv e l oss of stiffness and increasing dissipation with increasing maximum amplitude of strain energy , which is 5 in agreement with the so - called Experimentally , Sommer and colleagues (Sommer, Regitnig, Költringer, & Holzapfel, 2009) performed quasi - static extension - inflation tests at human common carotid arteries to understand the mechanical behavior of a vessel subject to loads beyond the physiological domain. Th e results if this study showed that t he burst pressure of ~60 kPa (~450 mmHg) may lead to damage or rupture of human carotid media - intima. A s imilar study , performed on human left anterior descending coronary arteries (Gerhar d A. Holzapfel, Sommer, Gasser, & Regitnig, 2005) , identified the jacket like behavior of the adventitia l layer at higher pressure that prevent artery f rom overstretch and rupture. Finally, several p revious experimental and numerical studies have demons trated that central arteries could sustain blood pressure values > 200 mmHg, without sustaining damage (Ferrara & Pandolfi, 2008a; Martin, Sun, Pham , & Elefteriades, 2013) , w hich is well over the physiologic range (Ferrara & Pandolfi, 2008b) . While pulmonary arterial pressure is approximately one sixth that of the systemic pressure (Lammers et al., 2012) , we do not know if the mechanical strength of the pulm onary arterial wall is comparable to other elastic arteries. T here are important aspect s of the mechani cs of pulmonary arteries , which could be contributing to PH initiation and progression, which have yet to be addressed . The goal of this dissertation is to investigate the irreversible mechanical damage that occurs in pulmonary arteries when subjected to s upra - physiological loading conditions. We aim to achieve this goal by completing the following objectives 1) To c haract eriz e experimentally the elastic and inelastic (irreversibly damaged) mechanical behavior of pulmonary arter ies ; 2) To identify change s in the arterial wall microstructure associated with damage, using modeling and histological analysis . 6 Figure 1 . Schematic of the layered microstructure of the arterial wall. (G. A. Holzapfel, Gasser, & Ogden, 2000) . 1.3 Pulmonary artery microstructural organization in healthy and PH Arterial wall microstruc ture. The wall of elastic arteries, including the pulmonary artery, is made of three layers. The layers are identif ied as intima, media and adventitia , form the innermost (luminal) to the outermost, see Figure 1 . The intimal layer, or tunica intima , consists of a thin layer of endothelial cells that are in direct contact with the blood flow. The medial layer, or tunica media , is composed of a series of concentric layers of elasti c l amina alternated with layers of v ascular smooth muscle cells (i.e., cells that have the ability of synthetizing proteins and of generating force by contracting) . The elastic laminae are made of tightly packed elastin fibers and microfiblis. The smooth muscle cells and the elastin lamina a re connec ted by radially oriented 7 fibers that allow the cells to sense their mechanical environment. The adventitial layer, or tunica adventitia , is composed of a three - dimensional network of connective tissue , mostly collagen type I and III, and fibroblas ts (i.e., cells that have the ability of synthetizing structural proteins such us collagen fibers). The collagen fibers in the adventitia are somewhat organiz ed along two preferential directions, which are at an angle (~ 45 ° ) with respect to the vertical a xis. This results in a three - dimensional spiral - shaped organization of bundles fibers , which can be referred to as fiber The collagen bundle has a coiled structure that will increasingly straighten with increased loading. Structural function of the wall layers . The intimal layer is thought to not have a considerable contribution to the mechanical behavior of elastic arteries. The medial layer gives the artery its strength and its elastic behavior, it is significantly engaged in the ho meostatic range of loading, and it offers re sistance to both longitudinal and circumferential loads. These behaviors are due to the presence of elastin fibers within the media. The adventitia is mostly engaged at higher values of stresses, it is thought to act as a protective sheath against over - pressurization (Bellini, Ferruzzi, Roccabianca, Di Martino, & Humphrey, 2014) . This behavior is mostly associated with the structural recruitment of the embedded wavy collagen bundles, and it is generally thoug h to lead to the characteristic anisotropic mechanical behavior of arteries (G. A. Holzapfel et al., 2000) . Because of this micr ostructural organization, h ealthy elastic arteries behave as deformable composite struc tures , show ing a highly nonlinear response and a stiffening effect at higher pressures. Furthermore, not just the organization of each constituent separately, but the ba lance of different constituents within the wall, has been thought to contribute significantly to the mechanical behaviors of healthy arteries (Armentano et al., 1991) . However, in diseased arteries, this microstructure could be disrupted, leading to the ina bility of arteries to efficiently perform 8 their mechanical function . Therefore, to understand arterial dysfunction, due to long term remodeling of the wall or due to damage mechanisms, it is crucial to investigate and quantify changes in the microstructure of each constituent. Structural function of extracellular matrix fibers. Collagen and elastin are the major load bearing constituent s of the arterial wall . S pecifically, the collagen fibers network has a stiff non - linear behavior, and the elastin fiber ne twork shows a more compliant linear elastic behavior (Roach & Burton, 1957) . In their seminal paper, Roach and Burton selective digestion to understand the role played by the collagen and elastin fiber network in defining the overall arterial behavior by isolating the ir contribution . Specifically, employing chemical degradation of collagen fibers and elastin fibers in iliac arterie s followed by mechanical testing, they concluded that most of the mechanical behavior of arteries at lower values of pressure is ascribed to elastin, while most of the mechanical behavior at high values of pressure is influenced by collagen fiber network. Another study by Wagenseil and colleagues (Wagenseil & Mecham, 2012) reviewed the mechanical properties and contribution of elastic fibers to arterial stiffness , using genetically modified mice. First , they recognized that elastin fibers played a crucial role in influencing not just the mechanics of the wall but also cell behavior. They observed that elastin knockout (Eln - / - ) mice die within a few days after birth because of the remarkably increase of cell number with in the arteri al wall , which obstructed the arteries . They also observed significant differences in incremental arterial stiffness between elastin knockout and wild type animals, despite the differences of aortic diameter at systole are minor in all genotypes, but the change in diameter between systole and diastole is approximately three times less in elastin knockout mice. Then , m ice ha ve been generated that express human elastin in a bacterial artificia l chromosome (BAC - ELN) . These animals are characterized by an increase in elastin amount from 30% to 120% 9 of normal levels. The results of this study suggested that elastin density is inversely proportional to arterial stiffness and blood pressure. All t he se studies taken together explained the mechanical role of the most relevant extracellular matrix components in the arterial wall , collagen and elastin. Damage modeling in soft tissues typically tends to be phenomenological, which requires the related para meters to be adjusted to specific experiments. Here, we propose to employ a micromechanically motivated approach to describe vessels before and after damage, to provide some physical interpretability of damage mechanism (Schriefl, Schmidt, Balzani, Sommer, & Holzapfel, 2015) . Changes of microstructural organization in PH . Previous pathological studies mainly established the relationship be tween hemodynamic factor s and the change in the mechanical properties of tissues (Humphrey, Baek , & Niklason, 2007; Sheidae i, Hunley, Raguin, & Baek, 2009; Sho et al., 2004; Zambrano et al., 2016) . O nly few studies that relate pressure - induced microstructural change to the changes in the mechanics of the arterial wall. Until recent years, Bloksgaard claimed the first study that quantitatively relate pressure - induced microstructural changes in resistance arteries to the mechanics of their wall. Both modeling and imaging data suggested that the acute pressure - induced structural changes of hu man pericar dial resistance arteries (hPRA) are small (Bloksgaard et al., 2017) . In addition, Sanz et al. (Sanz et al., 2009) found that in exercise - induced PH patients, the pulmonary artery stiffness increased without significant change of luminal diameter. A n accurate representation o f the effect of damage on each microstructural components is important to investigate PH , for the following reasons: (1) Different mechanical functions of artery . Arter ies is a multi - layered material consists of numbers of different constituents. Esp ecially collagen fibers in artery, which is predominantly present in the adventitia as a dense network and is believed to have a load - bearing function. 10 The collagen configuration evolves continuously until rupture occurs (Schrauwen et al., 2012) . Hence the pressure induced change of collagen network maybe monitored in over pressurizing procedure ; (2) Different mechanical response of su bject . Current understanding of PH pathology has largely been obtained through small animal models in which there has been seldom reports of caliber enlargement (Nickel et al., 2015) . Furthermore, permanent damage in healthy subject is not very common, the vessel wall damage has been re ported in in vitro test (Wulandana & Robertson, 2005) and pathological conditions. Thus, experiments on pulmonary arteries in large animals are need for validation whether pressure alon e can induce an enlargement caliber. 1.4 Pulmonary artery constitutive model Const itutive models describin g the mechanical properties of elastic arteries have been used to predict or investigate changes associated with disease initiation of progression . For example , models can predict blood flow and pressure, wall distension, normal and shear stresses, and ene rgy requirements in elastic arteries, including the pulmonary vasculature . The choice of an appropriate constitutive model, however, is complicated by the complex wall structure of elastic arteries, that dictates passive mechanical behaviors of these vesse ls (Hunter, Lammers, & Shandas, 2011) . The constitutive models can be categorized into two typ es: phenomenological and structural models. Phenomenological models are mathematical expression s that describe the behavior of vessels independent ly of any particular anatomic al or physiologic al parameters. Delfino et al. (Delfino, Stergiopulos, Moore, & Meister, 1997) proposed an isotropic rubber - like strain energy potential to describe carotid arteries . This strain energy function , which is able to model the typical stiffening at high pressure s, has the form 11 , (1) w here is a stress - li ke material parameter , is a non - dimensional parameter , and t he first invariant of the modified Cauchy - Green deformation tensor is defined as . Figure 2 . Contour plot of the convex potential (1) with material parameters = 44.2 kPa and = 16.7 (see (Delfino et al., 1997) ) A nother phenomenological strain energy function form have been introduced by Humphrey (Humphrey, 1995) , and has the form , (2) where is a material parameter (dimension of a stress) and is given by (3) 12 wh ere , are non - dimensional material parameters, while , for are the components of the modified Green - Lagrange strain tensor referred to cylindrical polar coordinates . The mod ified Green - Lagrange strain tensor can be written as , ( 4 ) where denotes the second - order identity tensor, and denotes the right Cauchy - Green tensor, defined as , ( 5 ) where represents the modified deformation gradient , defined to satisfy the relation ( represent the deformation gradient in large deformation). Compa red to the work of Delfino, this strain energy func tion has no a prior i restriction on the material parameters presented by assuming the artery is homogenous and incompressible. Figure 3 . Contour plots of the potential (3) with (a) material parameters , in (Chuong & Fung, 1983) , and (b) a set of parameters chosen to illustrate non - convexity . 13 Finally , another well - known form of strai n energy function for arteries is the one proposed by Takamizawa and Hayashi (Takamizawa & Hayashi, 1987) it has the logarithmic form , ( 6 ) where is a stress - like material parameter and the function is given in the form , ( 7 ) where are non - dimensional material parameters and , are the componen ts of the modified Green - Lagrange tensor in the circ umferential and axial directions, as defined in Eq. (4). D ue to the logarithmic form, in the particular condition , the value is infinite. Additionally, would lead to an undefined function of . Therefore, t his type of strain - energy function is only applicable for a limited range of states of deformation. Due to the strong influence of residual stresses, if (4) is used within a (displacement - driven) fin ite element formulation, may lead to numerical difficulties (G. A. Holzapfel et al., 2000) . The strain energy functions discussed above have the benefit of describing accurately the mechanical behavior of arteries. D ue to their phenomenological nature, however, these descriptions are lacking connection with anatomic and physiologic quantities, significantl y reducing the opportunity for independent validation of parameters (Hunter et al., 2011) . In the last 20 years, howev er, there has been an effort to incorporate microstructural information in the mechanical modeling of arteries. H olzapfel and Gasser first introduced these concepts, inspired by the mathematical methods used to describe fiber - reinforce and multi - laminated composites (G. A. Holzapfel et al. , 2000; Hunter et al., 2011) . Briefly, t he basic idea is to include histological information with in th e constitutive model so that the material parameters could be connected with 14 . For the first time, th e Authors introduced the idea of modeling the arterial wall as a bilayered structure, where each layer i s made of fiber - reinforced elastic materials with different characteristics. For example, the medial layer, which is mostly comprised of elastin, is des cribed by an isotropic material description, while the adventitial layer, which is mostly made of collag en fibers, is endowed anisotropic mechanical behaviors . Finally , the anisotropic and isotropic parts are summed together using the concepts of the tradi tional mixture theory, . ( 8 ) Specifically, in (G. A. Holzapfel et al., 2000; Hunter et al., 2011) the isotropic portion is described employing a neo - Hooken model , ( 9 ) where is a stress - like material parameter. While the collagen fibers are described by an exponential strain energy function , to represent the stiffen ing behavior at higher pressure, as follows ( 10 ) where is a stress - like material parameter and is a dimensionle ss parameter. Note that and are the squares of the stretches in the direction of and , that characterize the direction of two (reference) fiber families. Therefore, the reduced from is given by . (11) Based on the strain - energy function proposed by Holza p fel and Gasser (G. A. Holzapfel et al., 2000) , this d issertation will focus on determine the constituent - wise indication of the damage 15 of the pulmonary arterial wall. Collagen fiber stiffness para meter and fiber direction will also be estimated from the experimental data and histological images. 1.5 Mechanical test of elastic vessels in vivo and in vitro Cyclic inflation test has been employed in several previous studies to identify and investigate th e biaxial mechanical behavior of elastic and muscular arteries (Saravanan, Baek, Rajagopal, & Humphrey, 2006) . T hat is because through inflation test one could investigate physiologically meaningful mechanical conditions by preserving the native geometry of the vessel and mimicking the in vivo loading conditions during testing (Macrae, Miller, & Doyle, 2016) . Comparably, b i - axial extension test is better able to characterize the anisotropic behavior of an artery, since both circumferential and lo ngi tudinal directions are loaded simultaneously (Humphrey, 1995; Sacks & Sun, 2003) . Also, the applied force and the amount of stretch in circumferential and longitudinal direction can be controlled, allowing the in vivo state simulation unde r p hysiological condition (Tian & Chester, 2012) . Geneti c m odification and constituent purification combined with extension test is another approach that focus on analyzing the mechanics and microstructure of the isolated tissue. In 1998, Lillie and Gosline (Lillie, David, & Gosline, 1998) compared the behavior of purified elastin with and without its microf ibr ils by autoclaving procedures, in order to determine the contribution of the microfibrils to the performance of entire elastin meshwork. In addition to pressure - diameter and stress - strain relationship obtained via in vitro measurements, hemodynamic factor s could also be monitored through in vivo measurements. Many clinical studies expressed interests on determining the pulse wave velocity and pulse pressure associated aortic stiffening due to microstructural changes (Mitchell, 2009) . However, the current pulse wave velocity estimation is only capable of representing an average global 16 relationship between two remote measurement location due to lacking of the exact arterial geometry (Vappou, Luo, & Konofagou, 2010) . 17 CHAPTER 2 : Mechanical characterization of damage behavior in the pulmonary artery 2.1 Introduction We perform an i n vitro cyclic inflation tests to characterize the elastic mechanical behavior of porcine pulmonary arteries, and the effect of irreversible damage. Specifically, we analyze the pressure - diameter relationshi p before and after an over - pressurization to unde rstand the onset of the mechanical damage. Moreover, we use an elastic constitutive model of the arterial wall, which consists of two collagen fiber families and an isotropic elastin matrix (Bellini et al., 2014; Gerhard A. Holzapfel, Gasser, & Ogden, 2000) , to quantitatively compare the mech anical properties control and damaged specimens. Finally, the pressure - induced changes in micro - structural configuration of the pulmonary arteries are studied using histological images 2.2 Methods 2.2.1 Specimen preparation Chest cavities (i.e., heart, lungs, trach ea, esophagus ) are obtained from six adu lt, male pigs from the MEAT laboratory at Michigan State University, and then stored at - Prior to testing, samples are defrosted at room temperature for 24 hours, then the PA is separated from the right ventricle and from the aort a by removing the connective tissue. We then isolated the PA from the lungs, up to the second bifurcation, which allowed us to identify a right and left branch. One of the branches was selected for mechanical testing, randomized between left and right, whi le a ring was cut from the other branch for histology. Before mechanical tests, small branches are sutured to allow pressurization of the sample, up to obtain a length of ~ 10cm for the overall sam olution (HBSS) in the fridge . 18 Figure 4 . I nflation - extension testing device with specimen mounted in between 2.2.2 Mechanical testing Before mounting the samples for pressurization, geometrical characteristics have been recorded. Specifically, we recorde , along the circumferential direction , and total axial length , employing a Vernier caliper. Mechanical tests were carr ied out in a custom - built inflation - extension testing device, as previously published (Kim & Baek, 2011) . The sys tem has the capability to apply, simultaneously, axial pre - stretch via a linear motor, and luminal pressure via a remote controlled syringe pump. The samples' diameter was then recorder throughout the test using a CCD camera (Hitachi KP - M2A), while the pre ssure was measured using a pressure transducer (Honeywell FP2000). The fl uid used for pressurization was NaCl solution (9%). The main branch of the PA specimen was sec ured to a cannula on one end, to allow pressurization, and to a vertically placed suppor t on the other end. The axial pre - stretch was adjusted to ~ 10% of the original length. Using a custom LabVIEW program, we subjected each sample to the biaxial testing protocols consisting of three parts. After preconditioning, which consisted in 5 cycles of pressurization from 0 to 30 mmHg, each vessel was pressurized for three 19 sets of 10 loading unloading cycles, as follows: fi rst, from 0 to 50 mmHg (part 1) ; second, from 0 to 100 mmHg, to induce damage (part 2); and third, for 0 to 50 mmHg (part 3). Fi gure 5 shows the pressure diameter data over the course of the mechanical test for one representative specimen. The protocol also included a 1 - minute recovery period between each two successive parts of the test . We used the pressure - diameter curves coll ected during part 1 of the test to identify mechanical behavior of the vessels before damage, and the curves collected during part 3 of the test to identify the mechanical behavior of the vessels after damage. After testing, we repeated the measure of thic kness and axial length. 2.2.3 Histological analysis Sections from samples collected before damage (from the untested PA branch) and after damage (from the PA te sted branch, after completion of the test) were processed for histological analysis. Briefly, th e samples were fixed in a 10% formalin solution for an hour and then stored at room temperature in 30% ethanol before being embedded in paraffin and sectioned. Sectioning and staining were carried out by the MSU Histopathology Lab. Histological analysis we re focused on determining changes in the collagen fiber's and elastin's structure, by comparing samples collected before and after damage. To analyze collagen f using picrosirius red (PSR) and we imaged them with pol arized light. Finally, to investigate the elastin's structure we employed the Verhoeff - van Gieson (VVG) stain. 2.2.4 Mechanical model The structural properties of the control and damaged PA specimens are characterized using a non - linear hyperelastic model to determine the constituent - wise indications of the damage. Assuming an incompressible material, the Cauchy stress can be computed as (1 2 ) 20 where is the Lagrange multiplier enforcing incompressibility, and and are the deformation gradient and right Cauchy - Green tensor, respectively. In addition, W is the stored elastic strain en ergy in the material. Histological analyses on the arteries showed that the arterial wall is comprised of layers of elastin and collagen fibers. Mechanical response of the elastin content of the arterial wall is predominant ly is otropic (Gundiah, Ratcliffe, & Pruitt , 2009) , and thus is modeled as a neo - Hookean material. Alternatively, the contribution of collagen fibers to the anisotropic pa rt of the mechanical response is modeled as an exponential function, proposed by (Gerhard A. Holzapfel et al., 2000) . The total strain energy function c an be written as a summation of two contributions ( 1 3 ) where t he superscript denotes the - th fiber family ; , , and are material parameters (in this study and ) ; is the first invariant of the right Cauchy - Green tensor (i.e. Tr C ). Finally, the stretch ratio of the - th fiber family is defined as , where is the angle between each fiber family direction and the axial direction (here ). In this study, we aim to quantify the mechanical changes associated with damage due to supra - physiological pressure. Therefore, we decided to use the same constitutive fo rm to describe samples before and after damage. B est - fit values for the 4 model pa rameters ( , , , ) have been determined separately for each specimen in the before and after damage conditions specimen . Specifically, using a fminsearch funct ion ( MATLAB ) , we minimized the objective function (Baek, Gleason, Rajagopal, & Humphrey, 2007) . 21 ( 14 ) where and are the measured and computed intramural pressure, respectively. Moreover, the a xial stretch in the experiments are fixed to be ~ 10% with respect to the unloaded control specimen ( ). Therefore, this const raint is added to the objective function with a Lagrangian multiplier . It is worthy to note that for each specimen, th e first loading curve from part 1 (control) and 3 (after damage) are used for parameter estimation. 2.3 Results Of the six samples tested, only five were considered for our analysis; one sample was discarded due to excessive noise in the data. Pre - test me asurements of samples showed initial length of 117.90 ± 13.8 mm, wall thickness of 0.8 9 ± 0.063 mm, and initial outer diamete r of 13.71 ± 2.42 mm. Post - test measurements concluded that there was no significant change in the overall length of the specimens, whereas, the average wall thickness decreased to 0. 79 ±0.044 mm. 22 Figure 5 . Luminal pressure and outer diameter raw data for a representative sample, as collected throughout the mechanical test. Specifically, part 1 (dark gray line) describes the mechanica l behavior of the PA before the damage, part 2 (light gray line) describes the response to over - pressurization which generat es damage within the wall, and part 3 (dark gray line) describes the mechanical behavior of the PA after damage. Also shown, the pre conditioning protocol (dotted light gray line). In Figure 5 we show a representative set of raw pressure - diameter data as co llected throughout the mechanical test. During the initial mechanical test (part 1), the vessel exhibited quasi - elastic characterist ics, showing repeatable behavior for consecutive loading cycles. In the following, we will refer to the fi rst loading curve of this portion of the test as the before - damage behavior of the P A s. During over - pressurization ( p art 2), conversely, we observed a pronounced rightward shift of each consecutive loading cycle, which suggests the development of mechanical damage wit hin the arterial wall. Finally, in part 3 of the test, the vessel displayed some level of recovery of the elastic behavior, shown by the f act that consecutive loading cycles generate repeatable curves. All 23 specimens, however, showed qualitatively a signi fi cant increase in diameter for each value of pressure, when comparing curves from part 1 to curves from part 3 of the test. This seems to c on fi rm the hypothesis that the application of supra - physiological pressures can damage the arterial wall in a potentia lly permanent way. In the following, we will refer to the fi rst loading curve of part 3 of the test as the after - damage behavior of the PA s. 24 Figure 6 . Average pressure vs. normalized diameter curve for all specimens . and correspond to before and after damage diameters, respectively. The before damage behaviors have been collected during the first loading of part 1 of the test, while the after damage behavior have been collected during part 3 of the test . The asterisk indi cates that the normalized diameters were significantly different , at p < 0:05 , when comparing before and after damage behaviors . We then quantified the diameter increase associated with over - pressurization by comparing diameter s before an d after damage , fo r each value of pressure. First , we normalized the diameter values recorded throughout the test by the diameter value at zero pressure after preconditioning , for each sample . Then, we performed an across - sample average of the normalized d iameter before dam age (i.e., ) and after damage (i.e., ) for values of pressures included between 0 and 50 mmHg, shown in Figure 6 . Statistical analysis con fi rms that the normalized diameter increase observed when comp aring before and after damage behavi ors is signi fi cant (p < 0. 05 for 25 each value of pressure). The damaged specimens showed an enlargement amounting to ~20% of the diameter. Figure 7 . Change in diameter during different loading curves. Black circles repre sent that diameter difference between the first and second loading curve of part 2 of the test, gray squares represent the difference in diameter between the second and third loading curve. The * indicates that the diameters were significantly different at p < 0:05. In an ef fort to identify the pressure for which the damage start occurring, we compared the first three consecutive loading curves for part 2 of the test . Specifically, the difference between diameters of the first and second loading curves, and the difference bet ween diameters of the second and third loading curves are compared in Figure 7 . A significant softening behavior can be observed during the first over - pressurization to 100 mmHg. Arteries are enlarged after the first loading cycle, while the change in size is significantly smaller over the next cycles. This result confirms that the irreversible damage of mechanical behavior was caused by a high pressure. 26 Particularly, a pressure of 50 - 60 mmHg, where the diameter difference is the largest, seems to be a reas onable candidate to quantify the onset of damage. However, it is not clear if mechanical damage were present during part 1 of the test. Figure 8 . 4 Histology images of tissue samples from the PA. Top row: Picrosirius red under polarize d light (collagen fibers in red, yellow and green); bottom row: VVG (elastin in black, nuclei in purple, and cytoplasm in pink). Left: before damage samples; right: after damage samples. Bar in each image represents 500 Figure 8 shows histological imag es of a representative specimen s , before and after damage. Polarized PSR images show a color shift from bright red (larger diameters fibers) to yellow - green fibers (smaller diameter fibers) , which could indicate damaged in collagen fibers. On the other han d, although the elastin sheets a ppear to be more dispersed in the VVG stained image of the damaged specimen, there are no clear qualitative indications of increased damage in the elastin fiber network . 27 Finally, t he model describes the results accurately , according to the computed , for both the before and after damage specimens, as shown in Table 1 . c1,2 2 , - pressurization of arteries which is consistent with the observation in histology images. The results indicate a approximately 62% decrease in the dim ensionless material tion. It is ally rved. Table 1 . Best - fit material parameters for the 2 - fiber family model and estimated axial stretch. 2.4 D iscussion Characterizing dissipative behavior of vasculature, such as softening and damage, is an emerging area of research, in predicting potentia l risk of diseases and elucidating me chanisms of disease progression. Previous studies have investigated the mechanical behavior of arteries in pressures 28 higher than physiological range (e.g., over 150 mmHg) mostly in the systemic circulation systems where as lumen enlargement is not a common feature of the hypertensive patients. On the other hand, a larger calibe r of the vessels in the lungs is a characteristic feature of PH patients when compared to healthy individuals , e.g., 20% larger art erial diameter i n adult PH patients and 30% l arger arterial diameter in pediatric PH patients. T he central hypothesis that we sought to test in this study is that over - pressurization could lead to a permanent increase in the luminal diameter of proximal pulmonary arteries. To this end, we designed a mechanical testing protocol to characterizing the change in mechanical behaviors of porcine PA in response to damage. Then , we employed a micros tructurally inspired constitutive law to interpret the results and make hypothesize which tissue components could be more affected by the damage process. Finally, we employed histological ima ges to qualitatively support the proposed damage mechanisms . Whi le previous studies have investigated the irreversible mechanical response of the arteries both in physiological and pathological conditions ( Scott, Ferguson, & Roach, 1972) , to the arteries. Large pulmonary arteries are main conduits in a low - pressure system , in comparison to the systemic circulation , and this physiological function could result in a significantly structure when compared to other elastic vessels . For example , in cerebral arterial tissues, the mechanical response to cyclic loading to a maximum pressure of 100 mmHg was shown to have meter curves (D. Li & Robertson, 2009 . In this study we showed that in PAs , however, a pressure of 100 mmHg appeared to induce an irreversible effect on the mechanical behavior of the wall ( Figure 5 ). Furthermore, the mechanical test we performed showed th at p reconditioning to a pressure of 30 mmHg did not change the mechanical behavior of the artery significantly, indicating that no damage incurred for this pressure level . In addition, a 29 relatively sma ll softening behavior is observed in part 1 of the test s , where the highest pressure is 50 mmHg (data not shown, not significant) . However, our results indicated that a more pronounced damage across all 5 specimens was associated with part 2 of the test ( F ig ure 6 ), where all of the vessels have demonstrated that after the over - pressurization, the arterial wall is more compliant, yet the cyclic pressure - diameter curves exhi bit an elastic behavior. (Schriefl et al., 2015) , observed similar softer but elastic behavior when the arterial collagen was enzymatically removed in human abdominal aorta samples. The com s of the materi al parameters in Table 1 , where the parameter , when compare before and after damage samples . The parameters and represent, in the model employed here, the co nstitutive properties of the collagen in th at (Roach & Burton, 1957) - second and second - thi rd consecutive loading curve s during part 2 of the test demonstrate that a pressure higher than 50 mmHg induces sudden irreversible changes in the structure of the pulmonary arteries , whereas the pressure below 50 mmHg induced relatively smaller mechanical damage ( Figure 7 ) lly, the abrupt c hange in the slope of the curve suggest that the damaging pressure is ~ 60 mmHg. The continuous softening behavior throughout part 2 may i mply that the damag e threshold in each following sets of loading curves may be decreasing . However, t his behavior ma y be due to a constant softening in the arterial structure. Table 1 suggests also could , the angle between ry , show a n average 5 ° , which could suggest that the circumferentially ori ented collagen 30 bers are damaged. Similarly, Converse et al. (Converse et al., 2018) observed that the direct ion of the damage conforms to the direction of the over - stretch using a collagen hybridizing peptide in the ovine midd le cerebral artery specimens. Furthermore, our histology images also suggested that t could be significantly damaged by over - pressurization , reinfor cing our modeling prediction. This study has some limitations. First, the tests were p erformed in open air and room temperature, while other studies perform the tensile te sts in a saline solution (Schrauwen et al., 2012) . Second, we did not account for mechanical re sponse (active or passive) of smooth muscle cells in our study. Third, while we kept t he specimens axially stretched at 10%, we did not measure the axial force induced in the specimens during the pressuri zation. Fourth, we assumed that the damaged specimen s have are fully elastic, i.e., we did not use a damage model in our constitutive rela a result of over - pres surization in the pulmonary arteries. 2.5 Conclusion P H is a complex and multifaceted disease in which the structure of the proximal PAs c hanges as a result of a variety of biomechanical and biochemical factors. In this study, we aimed to analyze only the mechanical response of the proximal PAs under elevat ed pressure. The results presented here suggest that mechanical damage of the arterial wall, associated with significantly increased blood pressure, could be contributing to the pathology of PH. Furthermo re, the combination of model and histological images , seem to suggest that the damage is localized in the collagen fibers network in the a dventitial layer 31 BIBLIOGRAPHY 32 B IBLIOGRAPHY 1. Akhtar, R., Sherratt , M. J., Cruickshank, J. K., & Derby, B. (2011). Char acterizing the elastic properties of tissues. Materials Today , 14 (3), 96 105. https://doi.org/10.1016/S1369 - 7021(11)70059 - 1 2. Armentano , R. L., Levenson, J., Barra, J. G., Fisher, E. I., Breitbart, G. J., Pichel, R. H., & Simon, A. (1991). Assessment of elas tin and collagen contribution to aortic elasticity in conscious dogs. The American Journal of Physiology , 260 , H1870 - 7. 3. Botney, M. D. (199 9). Role of Hemodynamics in Pulmonary Vascular Remodeling. American Journal of Respiratory and Critical Care Medicine , 159 (2), 361 364. https://doi.org/10.1164/ajrccm.159.2.9805075 4. Chuong, C. J., & Fung, Y. C. (1983). Three - Dimensional Stress Distribu tion in Arteries. Journal of Biomechanical Engineering , 105 (3), 268 274. https://doi.org/10.1115/1.3138417 5. Delfino, A., Ste rgiopulos, N., Moore, J. E., & Meister, J. - J. (1997). Residual strain effects on the stress field in a thick wall finite element mode l of the human carotid bifurcation. Journal of Biomechanics , 30 (8), 777 786. https://doi.org/10.1016/S0021 - 9290(97)00025 - 0 6. Edwards, P. D., Bull, R. K., & Coulden, R. (1998). CT measurement of main pulmonary artery diameter. The British Journal of Radiology , 71 ( 850), 1018 1020. https://doi.org/10.1259/bjr.71.850.10211060 7. Faury, G., Pezet, M., Knutsen, R. H., Boyle, W. A., Hexim Mecham, R. P. (2003). Developmental adaptation of the mouse cardiovascular system to elastin haploinsuffi cienc y. The Journal of Clinical Investigation , 112 (9), 1419 1428. https://doi.org/10.1172/JCI19028 8. Ferrara, A., & Pandolfi, A. (2008). Numerical modelling of fracture in human arteries. Computer Methods in Biomechanics and Biomedical Engineering , 11 , 553 5 67. 9. H olzapfel, G. A., Gasser, T. G., & Ogden, R. W. (2000). A new constitutive framework for arterial wall mechanics and a comparative study of material models. Journal of Elasticity , 61 , 1 48. 10. Humphrey, J. D. (1995). Mechanics of the arterial wall: review and directions. Critical Reviews in Biomedical Engineering , 23 (1 2), 1 162. 11. Humphrey, J. D., Baek, S., & Niklason, L. E. ( 2007). Biochemomechanics of cerebral vasospasm and its resolution: I. A new hypothesis and theoretical framework. Annals of Biomedica l Eng ineering , 35 , 1485 1497. 12. Humphrey, J. D., & Tellides, G. (2018). Central artery stiffness and thoracic aortopathy. Ame rican Journal of Physiology - Heart and Circulatory Physiology , 316 (1), H169 H182. https://doi.org/10.1152/ajpheart.00205.2018 33 13. Hunter, K. S. , Lammers, S. R., & Shandas, R. (2011). Pulmonary Vascular Stiffness: Measurement, Modeling, and Implications in Norma l and Hypertensive Pulmonary Circulations. Comprehensive Physiology , 1 (3), 1413 1435. https://doi.org/10.1002/cphy.c100005 14. Kim, J., & Baek , S. (2011). Circumferential variations of mechanical behavior of the porcine thoracic aorta during the inflation test . Journal of Biomechanics , 44 (10). 15. Lammers, S., Scott, D., Hunter, K., Tan, W., Shandas, R., & Stenmark, K. R. (2012). Mechanics and Funct ion of the Pulmonary Vasculature: Implications for Pulmonary Vascular Disease and Right Ventricular Function. In Compr ehensive Physiology (pp. 295 319). https://doi.org/10.1002/cphy.c100070 16. Li Zhihe, Froehlich Jeffrey, Galis Zorina S., & Lakatta Edwar d G. (1999). Increased Expression of Matrix Metalloproteinase - 2 in the Thickened Intima of Aged Rats. Hypertension , 33 (1), 116 123. https://doi.org/10.1161/01.HYP.33.1.116 17. Macrae, R. A., Miller, K., & Doyle, B. J. (2016). Methods in Mechanical Testing of A rteri al Tissue: A Review. Strain , 52 (5), 380 399. https://doi.org/10.1111/str.12183 18. Mehari, A., Valle, O., & Gillum, R. F. (2014). Trends in Pulmonary Hypertension Mortality and Morbidity. Pulmonary Medicine , 2014 . https://doi.org/10.1155/2014/105864 19. Nicke l, N. Rabinovitch, M. (2015). Elafin Reverses Pulmonary Hypertension via Caveolin - 1 - Dependent Bone Morphogenetic Protein Signaling. American Journal of Respiratory and Critical Care Medici ne , 1 91 (11), 1273 1286. https://doi.org/10.1164/rccm.201412 - 2291OC 20. Roach, M. R., & Burton, A. C. (1957). THE REASON FOR THE SHAPE OF THE DISTENSIBILITY CURVES OF ARTERIES. Canadian Journal of Biochemistry and Physiology , 35 (8), 681 690. https://doi.org/10. 1139/ o57 - 080 21. Rounds, S., & Hill, N. S. (1984). Pulmonary Hypertensive Diseases. Chest , 85 (3), 397 405. https://doi.org/10.1 378/chest.85.3.397 22. Sanz, J., Kariisa, M., Dellegrottaglie, S., Prat - González, S., Garcia, M. J., Fuster, V., & Rajagopalan, S. (2009) . Eva luation of Pulmonary Artery Stiffness in Pulmonary Hypertension With Cardiac Magnetic Resonance. JACC: Cardiovascular Imaging , 2 (3), 286 295. https://doi.org/10.1016/j.jcmg.2008.08.007 23. Saravanan, U., Baek, S., Rajagopal, K. R., & Humphrey, J. D. (2006 ). On the Deformation of the Circumflex Coronary Artery During Inflation Tests at Constant Length. Experimental Mechanics , 46 (5), 647 656. https://doi.org/10.1007/s11340 - 006 - 9036 - 2 24. Schrauwen, J. T. C., Vilanova, A., Rezakhaniha, R., Stergiopulos, N., van de Vos se, F. N., & Bovendeerd, P. H. M. (2012). A method for the quantification of the pressure dependent 3D collagen config uration in the arterial adventitia. Journal of Structural Biology , 180 (2), 335 342. https://doi.org/10.1016/J.JSB.2012.06.007 34 25. Schriefl, A. J., Schmidt, T., Balzani, D., Sommer, G., & Holzapfel, G. A. (2015). Selective enzymatic removal of elastin and colla gen from human abdominal aortas: Uniaxial mechanical response and constitutive model ing. Acta Biomaterialia , 17 , 125 136. https://doi.org/ 10.1016/J.ACTBIO.2015.01.003 26. Sheidaei, A., Hunley, S. C., Raguin, L. G., & Baek, S. (2009). Simulation of abdominal ao rtic aneurysm growth with updating hemodynamic loads using a realistic geometry. ASM E J. Biomech. Eng. 27. Sho, E., Sho, M., Hoshima, K., Kimu ra, H., Nakahashi, T. K., & Dalman, R. L. (2004). Hemodynamic forces regulate mural macrophage infiltration in experim ental aortic aneurysms. Experimental and Molecular Pathology , 76 , 108 116. 28. Stepnowsk - & Raczak, G. (2017). Prognostic factors in pulmonary arterial hypertension: Literature review. Advances in Clinical a nd Experimental Medicine , 26 (3), 549 553. https://doi.org/10.17219/acem/61855 29. Takami zawa, K., & Hayashi, K. (1987). Strain energy density function and uniform strain hypothesis for arterial mechanics. Journal of Biomechanics , 20 (1), 7 17. https://doi.org/1 0.1016/0021 - 9290(87)90262 - 4 30. Thubrikar, M. J., & Robicsek, F. (1995). Pressure - induce d arterial wall stress and atherosclerosis. The Annals of Thoracic Surgery , 59 (6), 1594 1603. https://doi.org/10.1016/0003 - 4975(94)01037 - D 31. Truong, U., Fonseca, B., Dunning, J. (2013). Wall shear stress measured by phase contrast cardiovascular magnetic re sonance in children and adolescents with pulmonary arterial hypertension. Journal of Cardiovascular Magnetic Resonance , 15 (1), 81. https://doi.org/10.1186/1532 - 429X - 15 - 81 32. Tuder, R. M., Marecki, J. C., R ichter, A., Fijalkowska, I., & Flores, S. (2007). Path ology of Pulmonary Hypertension. Clinics in Chest Medicine , 28 (1), 23 vii. https://doi.org/10.1016/j.ccm.2006.11.010 33. V alentín A, Cardamone L, Baek S, & Humphrey J.D. (2009). Complementary vasoactivity a nd matrix remodelling in arterial adaptations to alter ed flow and pressure. Journal of The Royal Society Interface , 6 (32), 293 306. https://doi.org/10.1098/rsif.2008.0254 34. V orp, D. A., Wang, D. H. J., Webster, M. W., & Federspiel, W. J. (1998). Effect of in traluminal thrombus thickness and bulge diameter on th e oxygen diffusion in abdominal aortic aneurysm. Journal of Biomechanical Engineering , 120 , 579 583. 35. Wagenseil, J. E., & Mecham, R. P. (2012). Elastin in large artery stiffness and hypertension. Journal of Cardiovascular Translational Research , 5 , 264 273. 35 36. Wagenseil, Jessica E., Ciliberto, C. H., Knutsen, R. H., Levy, M. A., Kovacs, A., & Mecham, R. P. (2010). The import ance of elastin to aortic development in mice. American Journal of Physiology - Heart and Circulatory Physiology , 299 (2), H257 H264. https:/ /doi.org/10.1152/ajpheart.00194.2010 37. Wagenseil Jessica E., Ciliberto Chris H., Knutsen Russell H., Levy Marilyn A., Ko vacs Attila, & Mecham Robert P. (2009). Reduced Vessel Elasticity Alters Cardiovascu lar Structure and Function in Newborn Mice. Circulatio n Research , 104 (10), 1217 1224. https://doi.org/10.1161/CIRCRESAHA.108.192054 38. Wagenseil, Jessica E., Nerurkar, N. L., Knutsen, R. H., Okamoto, R. J., Li, D. Y., & Mecham, R. P. (2005). Effects of elasti n haploinsufficiency on the mechanical behavior of mou se arteries. American Journal of Physiology - Heart and Circulatory Physiology , 289 (3), H1209 H1217. https://doi.org/10. 1152/ajpheart.00046.2005 39. Wulandana, R., & Robertson, A. M. (2005). An inelastic mult i - mechanism constitutive equation for cerebral arteria l tissue. Biomech. Model Mechanobiol. , 4 , 235 248. 40. Zambrano, B. A., Gharahi, H., Lim, C. Y., Jaberi, F. A., Choi, J., Lee, W., & Baek, S. (2016). Association of Intraluminal Thrombus, Hemodynamic Forces , and Abdominal Aortic Aneurysm Expansion Using Longit udinal CT Images. Annals of Biomedical Engineering ,