Fabrication and characterization of fibrous poly(ε-caprolactone)/poly(3-hydroxybutyrate-co-3-hydroxyhexanoate) scaffold for cardiac patch application

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CORSO DI LAUREA MAGISTRALE IN CHIMICA INDUSTRIALE (classe LM-71)

Fabrication and characterization of fibrous

poly(ε-

caprolactone)/poly(3-hydroxybutyrate-co-3-hydroxyhexanoate) scaffold for cardiac patch application

Relatore Correlatore

Professoressa Federica Chiellini Professor Aldo Boccaccini

Controrelatore Professor Roberto Solaro

Candidato Giulia Rella

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Myocardial infarction is one the leading cause of death in western countries. Cardiac tissue engineering aims to replace infarcted myocardium, which has limited capability to regenerate by providing healthy functional cells to the injured region via a carrier substrate, and providing mechanical support, thereby preventing ventricular remodelling. The aim of this thesis was the development and characterization of highly porous scaffolds for cardiac tissue regeneration using poly(ε-caprolactone) (PCL) blended with poly(3-hydroxybutyrate-co-3-hydroxyhexanoate) (PHBHHx) to improve its cytocompatibility and hemo-compatibility. Electrospinning technique was chosen to develop biodegradable cardiac patches. Electrospinning conditions such as polymer concentration, applied voltage and distance from the collector were investigated and optimized. Microstructural analysis revealed that uniform and smooth fibers were obtained for all three fiber mats investigated which include PCL/PHBHHx 100/0, PCL/PHBHHx 70/30 and PCL/PHBHHx 30/70, although, PCL/PHBHHx 30/70 displayed fusion of the fibers junction. Mechanical properties of the developed patches were investigated through stress-strain displacement curves and showed that blending PCL with a flexible polymer like PHBHHx slightly enhanced the mechanical properties of the electrospun PCL/PHBHHx 70/30 fibrous mat proving that the two polymers are compatible although immiscible, as revealed by thermal analysis. In vitro cytocompatibility studies of PCL/PHBHHx fibrous mats demonstrated that the developed scaffolds promote adhesion and proliferation of C2C12 cell.

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Acknowledgement

I would like to express my deepest gratitude to Professor Aldo Boccaccini for the opportunity to pursue my thesis in his research group at the Institute for Biomaterials, Department of Materials Science and Engineering, University of Erlangen-Nuremberg. His countless and extremely valuable comments and advises were an inspiration to me. I would like to thank my supervisor Dr. Ranjana Rai, who helpfully followed me during my internship. I appreciate enormously all her contributions of time and ideas. Furthermore, I would like to thank Dr. Judith Roether, Dr. Menti Goudouri, Dr. Joachim Kaschta, Marwa Tallawi and Alina Grünewald for their technical assistance and fundamental help during the experiments. I am most grateful to Dipl.-Ing Heinz Mahler for his kindness and support with any problem. I would also like to thank all people from the BIOMAT group, especially my colleagues from room 0.211 for the great time we had together.

My special thanks goes to Professor Federica Chiellini for her generous support, patience and help during the writing of my thesis. I would like to acknowledge Professor Roberto Solaro for his constructive feedback and helpful suggestions. In the end I would like to thank my friends and family for their loving support, encouragement, and understanding throughout my entire academic career. Without them all of this would not have been possible.

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C

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1 Introduction

1.1 Cardiovascular diseases

Cardiovascular diseases (CVD) cause the highest number of dead in developed countries [1]–[4] reaching approximately 47% of all deaths in Europe and 40% in the EU [5]. Thus, alone in Europe more than 4 million people die every year of diseases of the heart and circulatory system, while in the EU deaths of CVD amount to 1.9 million people per year [5]. CVD lead to substantial financial burden for the EU economy estimated over € 196 billion annually [5]. Furthermore, changing nutritional habits and lifestyle lead to an increase in CVD also in the developing countries [1] while longer life expectancy and improvement of survival of subjects with coronary heart disease in developed countries [6] results in a rapidly rising number of patients. Coronary heart disease is the most common risk factor for left-side cardiac failure occurring through acute myocardial infarction [2]. Myocardial infarction is caused by the sudden obstruction of one or more coronary arteries [7]. The blood supply is interrupted and no oxygen and nutrients are delivered to the blood-deprived part of the heart muscle tissue [8]. This leads to substantial death of cardiomyocytes (myocyte necrosis). Cardiomyocytes enable the pumping of blood into the body by synchronized mechanical contraction. The cardiomyocytes contract more than three billions times during an average human lifetime, and pump more than 7000 liters of blood every day. Their morphology is elongated rod-shaped, when they are undamaged. Cardiac muscle automaticity controls the contractions of the heart almost entirely, and can be attributed to a groups of specialized cardiomyocytes (pacemakers), which are located in the

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sinoatrial node and enable the depolarization wave [9]. Apart from cardiomyocytes which constitute approximately 30% of the myocardium, the residual part, approximately 70%, are non-myocytes constituents [10]. This part of the myocardium consists of fibroblasts containing a dense supporting vasculature and a collagen-based extracellular matrix and together with the cardiomyocytes they form the native myocardium [9], [11]. The myocardium has not the ability to completely restore after myocardial infarction. This lies in the fact that myocytes have only a restricted capacity to regenerate. Around 1% of cells are regenerated per year at the age of 25 and less than 0.5% at the age of 75, meaning that merely 50% of cardiomyocytes are exchanged during a median person’s life span [12]. Healing mechanism after myocardial infarction is initiated by rigorous inflammatory response causing the migration of macrophages, monocytes and neutrophils into the infracted region [13], [14]. The inflammatory cells have the function to restore the ventricular walls, mediating the readsorption of necrotic tissue, scar formation and development of new blood vessels [15]. However, the failing load bearing ability of the compromised tissue enables wall thinning and ventricular enlargement. Ventricular enlargement can endure up to months until a temporary compensation is achieved [14]. Whereas capillary microcirculation is restored through regeneration of smooth muscle cells and endothelial cells, ventricular myocytes do not regenerate and are replaced by a fibrotic scar [15]. Scar tissue has altered properties and cannot contract contrary to beating myocytes causing a reduction of the heart contractile efficiency [16], [17]. Late ventricular remodelling is induced by hypertrophy to compensate increased wall stresses given by the fibrous scar. However, heart function is impaired enduringly. Finally this can lead to diastolic and systolic dysfunction and congestive heart failure [2], [4].

Although mortality rate from heart disease is declining as a consequence of improved treatment and primary prevention in industrialized countries, life expectancy for heart failure stays poor with a median survival of 2–3 year after

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establishing the diagnosis of heart failure [18]. Advances in medical treatment can improve life quality of patients after myocardial infarction, however they are unable to restore the cardiac function. Surgical treatments such as coronary artery bypass grafting, heart valve replacement, implantable left ventricular assist device and finally heart transplantation are required for patients with marked symptoms and limitation in activity, and with end-stage heart failure [2]. Heart transplantation is probably the last option for end-stage congestive heart failure patients, but due to problems of severe shortage of donors and the need for immune suppressive treatment this option is often impracticable [19].

Therefore constantly new strategies to repair the injured heart and restore cardiac function are investigated [20]. Figure 1.1 displays the different approaches studied for alleviation of myocardial infarction. The ideal strategies for clinical interven-tional therapy would be either to avoid scar formation, or replace formed scar tissue with functioning cardiac muscle tissue. In a primary approach to such therapy, scientist have investigated new cells injections into the injured areas of the cardiac tissue [17]. These studies stared in the mid-1990s using cell transplantation techniques to repair the failing heart [20]. Cells of various origins and developmental stages have been injected into healthy and diseased hearts, including skeletal myoblasts, bone marrow stem cells or adipose stem cells, foetal cardiomyocytes, embryonic stem cell–derived cardiomyocytes and potential cardiac progenitors from peripheral blood, bone marrow or heart [2], [21]. Among these cell sources, embryonic stem cells have proven great prospective due to their ability to be cultured in the undifferentiated state in vitro and subsequently differentiated in vitro into cells of the human myocardium [2]. Most studies have reported that cell engraftment in various animal models of myocardial infarction and various disease models can improve contractile function [22]. The exact mechanism of functional improvement has not been well established yet. It could include cardiogenesis from engrafted cells, paracrine effectors homing of cardiac progenitors [21]. The results of clinical trials have demonstrated the safety of this

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procedure, although the benefits in terms of increased ejection fraction are modest [1]. At present major drawbacks include cell death, exit of cells from the heart, and reduced cellular integration with the receiving heart tissue [17]. To replace approximately 1 billion cardiomyocytes a high cell number is needed as up to 90% of cells are lost in the blood circulation during delivery and of those injected successfully approximately 90% die within a week [7], [23]. Therefore, further improvements have to be achieved concerning scaling up of cells, cell delivery, efficiency of grafting to avoid leakage of the cells from the injured side and immune rejection [2].

Figure 1.1: (a) Classical cell therapy in the heart (b) Tissue engineering approaches with cell sheets,

scaffolds or injectable materials (c) Ventricular restrain device [1]

Cell sheets have been proposed to deliver cells into the infarcted region. This technique has the advantage that cells are not damaged by detachment from the culture surface. Cells are kept together by their own extracellular-matrix developed during culturing allowing them to preserve their electromechanical

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connection. However the number of cell needed and the difficulty of overlapping several layers of cells caused by a lack in vascularization, is a major drawback in their successful use [1].

A repair strategy termed dynamic cardiomyoplasty was developed in 1987 in which the heart is wrapped around with latissimus dorsi muscle that functions as a mechanical support to avoid ventricular dilatation and restore myocardial contractility [1], [24]. In later studies wrapping constructs were realized with biomaterial meshes that constrained the heart by wrapping the biomaterial meshes around. However, clinical trial investigations have raised concern about the technical difficulties and risk introduced by this approach for patient with end-stage congestive failure [25]. Further concerns include the potential of producing significant pathologic diastolic restriction and restrictive pericarditis [23]. Finally this approach can prevent remodeling, but does not repair or regenerate injured myocardium [3].

Alternative approaches assessed the use of biomaterials for injections into the infarcted heart [26]. One group suggested the addition of a heart patch to the border zone. According to the theoretical simulations the biomaterial injected into the border of the infarcted zone with stiffness up to 200% of the average stiffness of the passive myocardium could reduce infarct expansion and global left ventricle remodelling by decreasing heart wall stress. The higher the stiffness of the material the higher was the reduction in wall stress induced by the patch [27]. In vivo simulations confirmed partly these findings [28], [29]. The stiffness of the heart muscle is known to be 10 kPa at the beginning of diastole and reaches 200– 500 kPa at the end of diastole. Early studies on biomaterial injections evaluated natural polymers and showed increase in immunogenicity while studies on synthetic materials showed that they could trigger an inflammatory and foreign body responses [26].

This has motivated researchers to find alternative delivery techniques capable of providing the same advantages of the cells transplantation but also supply

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mechanical support to the infarcted myocardium. Therefore procedures based on combining cells with regenerative capacity, proangiogenic growth factors, biological matrices and biocompatible synthetic polymers have been studied. The combination of these techniques is identified as tissue engineering [30]. The approach helps to deliver and graft cells into the heart by using a tissue engineered synthetic biodegradable patch [22]. This tissue engineered construct aims to support the defected cardiac muscle by providing healthy cardiac cells and provision for left ventricular restrain. The cardiac patch thus repopulates the diseased heart and prevents left ventricular dilation [21]. Scaffolds are generally seeded with cells in vitro and implanted subsequently onto the infarct regions. Biomaterial scaffolds have been increasingly used in targeting heart repair after infarction, as they can provide a proper support for cell survival, proliferation and also stem cell differentiation [4]. Furthermore they provide a guide for three-dimensional tissue reconstruction [31]. Proper myocardium regeneration and function may be largely dependent on the properties of the scaffolds. This approach shows promising features however it has to meet the many requirements of a tissue engineering construct.

1.2 Tissue engineering

Tissue engineering has been defined in 1997 as “an interdisciplinary field that applies the principles of engineering and the life sciences toward the development of biological substitutes that restore, maintain, or improve tissue function” [32]. Thus, tissue engineering wants to provide functional equivalents of native tissues that can be used for in vivo implantation of a tissue or organ. This brings to a variety of challenges, which can be divided into three main categories; science and technology of cells, materials, and cell-material interaction [2]. Normally a highly porous artificial extracellular matrix or scaffold is used to accommodate

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mammalian cells and to grow the cells for tissues regeneration in three dimensions. In the ideal case, cells are retrieved directly from the patient and cultured in vitro to multiply their population. The cells are than seeded on a three dimensional scaffold and grafted back into the same patient. The three dimensional scaffold should provide sufficient mechanical strength to reinforce and support the injured tissue [33]. Furthermore, the scaffold must be able to promote cell adhesion, proliferation and maturation. This means it must be highly porous with desired pore size and interconnectivity able to provide transfer of oxygen and nutrients to the seeded cells. However the material must be also biocompatible and have adequate surface morphology and chemistry. Physical-chemical and mechanical properties of the scaffold should be tailored to match degradation profile with the regeneration rate of the injured tissues [34]. This means the scaffold must be biodegradable with rates compatible with decreased need in mechanical support as the tissue regenerates. Finally non-toxic degradation products must be generated that can be metabolized and expelled from the body. Generally the scaffold is engineered to match the native extracellular matrix as the extracellular matrix plays a significant role in regulating the structural and mechanical properties of the entire tissue as well as providing structural and molecular signaling to guide the assembly of the surrounding cells [35]. Moreover, also ease of fabrication and commercialization are aspects that have to be considered when producing a scaffold for tissue engineering applications.

1.2.1 Cardiac tissue engineering

Tissue engineering strategy for patients with terminal heart failure can be divided in two major classes. The first and most simple approach is to use a cardiac patch of engineered heart muscle that is placed over the ischemic side. The second approach foresees to remove the injured side and replace it with the engineered

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heart muscle scaffold. This approach is certainly more complicated and requires high mechanical stability of the scaffold to endure systolic pressure [36]. Thus, the former approach using a heart patch instead of replacing the injured tissue is currently the most realistic perspective.

Nevertheless, criterions for a tissue engineered construct listed above are still essential requirements in cardiac tissue applications. The scaffold must be biocompatible, biodegradable and biomimetic meaning that the scaffold should replicate the physical and mechanical properties of the extracellular matrix. The microenvironment in which the cells are placed is extremely important in maintaining their basic properties and function. In fact, studies have confirmed that communication between cells depends mainly on their interactions with components of the extracellular matrix [30]. Therefore, the scaffold must provide the correct interaction with the cells and mimic their microenvironment. For clinical utility, cell density and dimensions are essential aspects for an engineered cardiac construct [7]. The human heart is thick and highly cellularized with a thickness of ∼1.5 cm and a cell density of approximately 2×108 cells/cm3 [19]. Consequently, engineered cardiac patches are requested to have a certain thickness, generally of several millimeters and an average area of 10–50 cm2. Major difficulties lie in the ability to engineer viable millimeters thick cardiac constructs capable of integration to the vascular supply of the host [36]. Thus, elevated porosity and interconnectivity of pores is a crucial requirement to achieve high colonization and migration of the cells in addition to vascularization. Porosity promotes angiogenesis within the scaffold, which is essential for maintaining a permanent flow of oxygen and nutrient between the cells. This facilitates efficient migration and survival of the cells in the infarcted tissue. However not only pore distribution is important but also pore size has to be carefully considered. Cardiomyocytes have a dimension that ranges from 10–100 µm while endothelial cells, that are responsible for vascularization, range from 8–10 µm. Pore size that exceeds remarkably the size of endothelial cells prevents them from bridging the

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pores and from migrating into the scaffold. Furthermore it provides insufficient surface area for the cells to attach and proliferate. On the other hand too small pores can cause the obstruction of the scaffold and lead to necrosis in the core of the scaffold. Therefore a cardiac tissue engineered construct should have pores that range maximal from tents to hundreds of micrometers to present best conditions for cell fostering and angiogenesis [7].

From the mechanical point of view, the heart muscle has dynamic functional properties and undergoes cyclic contraction and relaxation. Therefore the governing factors for the choice of a scaffold are not only biodegradability and biocompatibility, but also the mechanical properties of the matrix. The scaffold must be able to endure the contractile and expansive forces of the heart of each cycle. Elasticity and strength of the scaffold are extremely important and must ideally match the ones of the heart. Cardiac muscle shows a Young’s modulus of approximately 10–20 kPa at the beginning of diastole with a strain smaller than 10%. At the end of diastole a healthy heart has a Young’s modulus of approximately 50 kPa while a diseased heart has a Young’s modulus that reaches 200–300 kPa and strain is reported in the order of 15–22% [2]. This is due to the formation of the fibrous scar responsible for changing the stiffness of the infracted region from soft to rigid [36]. Therefore the ideal range reported for the elastic modulus of a heart patch is in the range of 3–15 kPa [2], while that of the ultimate strain lies in the region between 22–90%. The application of a cardiac patch with stiffness from 1–200% of the native myocardium in the border zone of the infracted tissue has demonstrated to be efficient in decreasing heart wall stress, while direct application in the infarcted zone was not able to enhance contractile efficiency of the heart [27]–[29]. Materials with elastomeric properties have the appropriate characteristics to sustain cyclic stress (Table 1.1).

However an additional obstacle to overcome is the nonlinear elasticity of heart muscle together with the anisotropic structure. The three dimensional structure of the scaffold has to provide an adequate surface topography and mechanical

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environment, to ensure that cardiac myocytes can organize into an electrical and functional syncytium. Surface topography is responsible for cardiomyocyte organization in vivo, which in turn affects cardiac function [37]. The anisotropic structure of the myocardium is characterized by fiber alignment along one direction on top of which elongated and aligned cells lay. The alignment of the cardiomyocytes helps to provide the optimal mechanical and electrical coupling between them [38]. Fibrils of a rat myocardium show diameter of approximately ∼100 nm. Mechanical properties of the heart change in dependence of the direction of stress applied. The tissue engineered construct should therefore display the same anisotropicity as the native tissue.

Table 1.1 Investigated or potential biomaterials used in cardiac muscle engineering (adapted from

Chen et al. [2]) Polymer Elastomer (E) or thermo-plastic (T) Young's modulus (or stiffness) (MPa) Tensile strength (MPa) Degradation (month) Synthetic, nondegradable Poly(ethylene terephthalate) (PET) T or E 4000 Nondegradable Poly(ethylene terephthalate)/

dilinoleic acid (PET/DLA) E 2–3 Nondegradable

Synthetic, degradable

Poly(glycolic acid) (PGA) T 7000–10000 70 2–12

Poly(L-lactic acid) (PLLA) or

poly(D,L-lactide) (PDLLA) T 1000–4000 30–80 2–12 Poly(β-hydroxybutyrate) (PHB) E 2000–3000 36 Degradable poly(p-dioxanone) (PPD) E 600 12 6 Polyurethane (PU) E 5–60 20–45 Surface errodible 1,3-trimethylene carbonate (TMC) E 6 12 Degradable 1,3-trimethylene carbonate– poly(D,L-lactide) (TMC– PDLLA) (50:50) E 16 10 Degradable Poly(1,8-octanediol- co-citric

acid) (POC) T 1–16 6.7 Degradable

Poly(glycerol sebacate) (PGS) E 0.04–0.282 0.5 Degradable

Naturally occurring

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(tendon/cartilage/ ligament/bone)

Collagen gel (calf skin) E 0.002–0.022 0.001–0.009 Degradable

Heart muscle

Myocardium of rat E

0.14 (at the end of

diastole) 0.03–0.07 Not Applicable

Myocardium of human E

0.2–0.5 (at the end

of diastole) 0.003–0.015 Not Applicable

In summary, we could list that a fabricated tissue engineering scaffold for cardiac patch applications should be [10]:

• highly porous with interconnected pores (to facilitate mass transport and vascularization),

• biomimetic (to mimic the highly complex architecture of the heart in which cardio-myocytes promote their correct organization that is essential for optimal mechanical and electrical coupling)

• elastic (to enable transmission of contractile forces ) • mechanically stable (to support the injured tissue),

• biodegradable, biocompatible and non-immunogenic (to degrade in the time the extracellular matrix is formed without or minimal inflammatory response)

To fulfill all the requirements described above is a challenging task. Biomaterial as well as fabrication technology of the scaffold must be carefully chosen to create an engineered myocardial patch that should be easy to sterilize and to seed with cells that will proliferate to form a functional myocardium.

1.2.2 Biomaterials for cardiac tissue engineering

A biomaterial can be defined in various ways. The ‘Williams Dictionary of Biomaterials’ defines a biomaterial as “a non-viable material used in a medical device, intended to interact with biological systems” [39]. Another definition of biomaterial is “a substance that has been engineered to take a form which, alone

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or as part of a complex system, is used to direct, by control of interactions with components of living systems, the course of any therapeutic or diagnostic procedure” [40]. Biomaterials are indispensable in most tissue engineering strategies. They can be found in a wide variety of compositions and with many different properties. Therefore the range of different types and classes of biomaterials used in the tissue engineering field is surely enormous. Also for cardiac tissue engineering many different types of materials have been considered [1]–[4]. Arnal-Pastor et al. [1] distinguish three main types according to their origin: natural materials, decellularized tissues and synthetic materials. Natural occurring materials include collagen, gelatin, fibrin, chitosan, silk and alginate as well as polyhydroxyalkanoates (PHAs); and synthetic materials include poly(ε-caprolactone) (PCL), poly(glycerol sebacate) (PGS), polyurethane (PU), poly(lactic acid) (PLA), poly(glycolic acid) (PGA), among others. Polymers, both natural and synthetic, are the most utilized materials in soft tissue engineering [2]. They are also the largest class of engineered biomaterials used today for myocardial tissue reconstruction [3] (Table 1.2).

Table 1.2 Overview of biomaterials used in cardiac tissue engineering [2]

Biomaterials Physical states Tissue engineering approaches

Naturally occurring (also

degradable)

Collagen based Gel Epicardial heart patch

Fibrin glue Injectable gel Endoventricular heart patch Peptide nanofiber Injectable gel Endoventricular heart patch Collagen mesh 3D porous mesh 3D tissue engineering construction Collagen-glycosaminoglycan 3D porous mesh 3D tissue engineering construction Gelatin mesh 3D porous mesh 3D tissue engineering construction Alginate mesh 3D porous mesh 3D tissue engineering construction

Synthetic

Biodegradable

Poly(glycolic acid)(PGA) and copolymer with poly(lactic acid)

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Poly(L-lactic acid) (PLLA) 3D porous mesh 3D tissue engineering construction Poly(ε-caprolactone) (PCL) and

copolymer with poly(lactic acid)

(PLA) 3D porous mesh 3D tissue engineering construction Poly(glycerol sebacate) (PGS)

3D porous foam (sponge)

3D tissue engineering construction and epicardial heart patch

Polyurethane (PU)

3D porous foam (sponge)

2D tissue engineering films that could be used as heart patches Poly(ester urethane) (PEU) [or

poly(ester urethane)urea (PEUU)] Solid sheet Cardiovascular grafting

Nondegradable

Poly(ethylene terephtalate) (PET). The fibers are manufactured under trade names Dacron, Terylene &

Trevira Knitted mesh Left ventricular constrain

Solid sheet Cardiovascular grafting

Polypropylene (PP) Solid Left ventricular constrain Poly(tetrafluoroethylene) (PTFE)

with or without PGA/PLA Solid sheet

Treatment of congenital heart disease

Cardiovascular grafting

Poly(N-isopropylacrylamide)

(PNIPAAm or PIPAAm) Solid sheet Scaffold-free cell sheet

Due to their wide range of properties, that can be customized to match the properties of soft tissues, synthetic polymers have been implanted successfully in many cardiac tissue applications [8]. Thus, their greatest advantages are their variable properties in term of mechanics, chemistry, and degradation as well as their tunable surface properties in terms of morphology and chemistry. Adhesion peptides or biological molecules can be used to bio-functionalize the surface to enhance functional cellular interactions [38]. Moreover, synthetic polymers have reduced or absent immune-responses and high reproducibility, but they can cause inflammatory responses induced by the original material or the degradation products. Furthermore, disadvantages are the fact that they may not exhibit the superior biological response that natural materials usually display [41].

Some of the first synthetic polymeric materials used for heart tissue engineering were based on aliphatic polyesters composed of polylactic acid (PLA), polyglycolic adic (PGA) and their co-polymer polylactic-co-glycolic acid (PLGA) [3], [8]. These polymers are hydrolytically degradable biocompatible polymers and their

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degradation products, glycolic and lactic acid, are completely metabolized by the human body. However it has been reported that the acidic by-products cause an inflammatory response damaging the local tissue [8]. Subsequently many other synthetic polymers have been studied, such as poly(ε-caprolactone) (PCL) [23], [42]–[48], polyurethane (PU) [49], [50], Poly(glycerol sebacate) (PGS) [22], [51]– [53], and acrylate based material [54], [55]. However only poly(ε-caprolactone) (PCL) and polyhydroxyalkanoates (PHAs) will be further discussed in detail in this Chapter.

Natural polymers have attractive properties like their non toxicity, renewability and biodegradability. Generally they induce a more mild inflammatory response. Nevertheless faster degradation and lower mechanical properties in addition to reduced reproducibility are major drawbacks in their applications [2].

Natural polymers used in cardiac tissue applications include naturally occurring extracellular matrix components as well as materials from different natural sources (Table 1.3). The native extracellular matrixes of soft tissues are composed of different types of collagens, a protein that is the main component of connective tissue and the most abundant protein in mammals. Thus, much research has been focused on collagen for tissue engineering application [4]. Already in 1997, Eschenhagen et al. [56] reported an artificial heart tissue of collagen gel, which was combined with embryonic chick cardiomyocytes. They were able to produce a contracting syncytium in vitro on the collagen matrix. Unfortunately, collagen has a great swelling rate and poor mechanical properties in aqueous medium [1]. Presently, other natural polymers such as alginate, gelatin, fibrin and chitosan are under intensive investigation for myocardial tissue engineering as shown in Table 1.3. Moreover, Sodian et al. [57] firstly reported on the natural-origin poly(hydroxyalkanoate) (PHA) for cardiac tissue applications. These microbial aliphatic polyesters circumvent the mechanical issues of natural polymers. A great variety of copolymers of this family can be produced making it possible to tailor their properties and their degradation rate by varying their copolymer ratio.

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Table 1.3 List of naturally occurring polymers, their sources and main application fields [2]

1.2.2.1 Poly(

εεεε

-caprolactone) (PCL)

The synthesis of poly(ε-caprolactone) (PCL) was described in the early 1930s and thus, it can be defined as one of the earliest polymers studied [58]. PCL is a hydrophobic and semi-crystalline thermoplastic polyesters with low melting point (59–64 °C) and excellent compatibly for blending with many polymers. Variations in melting point and the degree of crystallinity are dependent on the molecular weight, that ranges generally between 3000–100000 g mol-1 [59]. Moreover, PCL is soluble in many organic solvents and a cost-effective candidate for many

Polymer Source Main application fields

Collagen Tendons and ligament

Multi-applications, including cardiac tissue engineering

Collagen- glycosaminogly can (alginate) copolymers

Artificial skin grafts for skin replacement

Albumin In blood

Transporting protein, used as coating to form a thromboresistant surface

Hyaluronic acid

In the ECM of all higher animals

An important starting material for preparation of new biocompatible and biodegradable polymers that have applications in drug delivery and tissue engineering, including cardiac tissue engineering

Fibrinogen-fibrin

Purified from plasma in blood

Gelatin

Extracted from the collagen inside animals’ connective tissue

Chitosan Shells of shrimp and crabs

Matrigel™ (gelatinous protein

mixture) Mouse tumor cells Myocardial tissue regeneration Alginate

Abundant in the cell walls of brown algae

Polyhydroxyalk

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applications. Owing to its biodegradability and biocompatibility it has encouraged extensive research into its potential application in the biomedical field. The attention was drawn to this biopolymer due to its interesting mechanical properties, ease of manufacturing, good processability and customizable degradation kinetics. PCL can be prepared in two ways: by ring-opening polymerization starting from of ε-caprolactone via catalytic polymerization or starting from 2-methylene-1-3-dioxepane via free radical polymerization [58]. The superior rheological and viscoelastic characteristics compared to numerous other biodegradable polymers, which contribute to its simple processability, have made it particularly appealing for tissue engineering applications. Hence, PCL was transformed into a great variety of scaffolds with numerous fabrication techniques [59]. Porogen leaching [60], [61], phase separation [61], electrospinning [59], [62], [63], melt spinning [64], gravity spinning [65] and solid freeform techniques such as 3-dimensional printing [66], bioextrusion [67], [68], stereolithography [69], selective laser sintering [23], are some of the most important fabrication techniques used to obtain highly porous three dimensional scaffold of PCL.

Degradation of PCL is quite slow and occurs in two stages: first the molecular weight decreases due to random hydrolytic chain scission of the ester linkages. This stage was reported to be non-enzymatic as no evident differences in degradation behavior was shown in vitro and in vivo [58], [70]. In the second stage of degradation the polymer undergoes intracellular degradation with particle uptake into phagosomes of macrophages and giant cells and within fibroblasts. Total degradation of the polymer was reported to occur in 2–4 years depending obviously on the initial molecular weight [58].

As previously described, PCL is an exceptional candidate for blending or copolymerization. Blendes of PCL with other polymers demonstrated two different blending behaviors; showing only one Tg, when the polymers were completely miscible or distinctive Tgs ascribable to the single components of the blend, when the blends were immiscible. However, in the latter case, in some experiments, the

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blends exhibited superior mechanical properties indicating compatibility between the components, or else they proved to be incompatible exhibiting the enhanced properties of phase-separated material [58].

PCL has been proposed for myocardial regeneration as original material, as blend or as copolymer. Shin et al. [42] reported the formation of an electrospun nanofibrous PCL graft for cardiac tissue applications. The cardiac graft was seeded with cardiomyocytes from neonatal Lewis rats, which started contracting after 3 days. The cardiomyocytes attached well on the PCL fiber mats and expressed cardiac-specific proteins indicating that cardiomyocyte sheets can be matured in vitro [42]. Subsequently Ishii et al. [43] demonstrated the possibility to overlay the electrospun PCL mats cultured with cardiomyocytes to construct a thicker 3– dimentional cardiac graft. The PCL mats were overlaid between days 5 and 7 and exhibited strong synchronized contractions and no ischemia was found in the center of the constructs.

Blends and copolymers of PCL with PLA, PGA and PDLLA have also been investigated by various groups. Poly(glycolide-co-caprolactone) (PGCL) biodegradable scaffolds seeded with bone marrow-derived mononuclear cells have been implanted in rat myocardial infarction model. The implants were able to attenuate left ventricular remodeling and left ventricular systolic dysfunction. Differentiation to cardiomyocytes and migration of the bone marrow-derived mononuclear cells inside the scaffold was observed as well as neovasculature in infarcted areas and in infarct border zones [44]. PCL-co-PLLA cell-seeded patches were compared with gelatin, PCL, PGL cell-seeded patches and their ability to replace a defect in the right ventricular outflow tract of rats was investigated. PCL-co-PLLA showed best cellular ingrowth in vitro than PCL and PGA patches [45]– [47]. Yeong et al. [23] employed selective laser sintering (SLS) with an automated algorithm to fabricate porous PCL scaffold for cardiac tissue engineering. High density of mouse pre‐myoblast cells (C2C12) on the scaffold was recorded after 4 days of culture, while after 6 days, fusion and differentiation of the cells was

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observed in vitro with formation of multinucleated myotubes. Furthermore they demonstrated feasibility of fabricating a scaffold with adequate stiffness for cardiac tissue engineering and ease of tailoring the mechanical property using this technique. Recently Fleischer et al. [48] reported the fabrication of 3-dimensional spring-like fiber scaffold to successfully mimic the coiled perimysial fibers of the heart. They showed that the spring-like PCL fiber scaffolds promoted the formation of a functional cardiac tissue capable of generating a strong contraction force.

1.2.2.2 Polyhydroxyalkanoates (PHAs)

Polyhydroxyalkanoates (PHAs) are a class of optically active biological polyesters produced by various microorganisms, such as soil bacteria, blue-green algae, and genetically modified plants [71]. PHA granules are accumulated in the cytoplasm as intracellular storage of carbon and energy, when bacteria are exposed to unbalanced nutrients supply, with limitation of oxygen, nitrogen, sulphur, magnesium and phosphorous and excess supply of carbon source [72]–[76]. PHA were firstly discovered in the 1920s and gained great interest for biomedical applications in the 1980s because of their biodegradability, biocompatibility and good processing properties [34]. Since then much research has been published on PHAs for tissue engineering applications [73], [77] making them among the most investigated biomaterials for the development of tissue-engineered scaffolds, both for hard and soft tissues, including cardiovascular products [76]. There are more than 150 different types of PHAs offering an extremely wide variety of physical and mechanical properties. The general structure is displayed in Figure 1.2. The polymer backbone can consist of 3, 4, 5, and 6-hydroxyalkanoic acids with tunable length of the side chain. Depending on their side chain length, PHAs are classified as a) short chain length PHAs containing 3–5 carbon atoms in the monomeric unit;

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b) medium chain length PHAs that contain 6–14 carbon atoms in the monomeric unit; c) long chain length PHAs, which contain more than 14 carbon atoms [73].

Figure 1.2: The general structure of polyhydroxyalkanoates. R1/R2 = C1– C13 alkyl groups, x = 1–4,

n = 100–30,000 [72]

PHAs with short side chain tend to be more crystalline and therefore more brittle, as for example poly(3-hydroxybutyrate), P(3HB), poly(4-hydroxybutyrate) P(4HB), whereas PHAs with longer side chains are generally rubber-like, showing typical elastomeric behaviour, as for example poly(3-hydroxyhexanoate), P(3HHx), and poly(3-hydro-xyoctanoate), P(3HO). Properties of these polymer can be easily customized by varying the length of the side chain and by inserting more than one type of a hydroxyalkanoate as monomer unit creating copolymers such as poly(3-hydroxybutyrate-co-3-hydroxyvalerate), P(3HB-co-3HV), poly(3-hydroxyhexanoate-co-3-hydroxyoctanoate), P(3HHx-co-3HO), and poly-3-hydroxybutyrate-co-3-hydroxyhexanoate, P(3HB-co-3HHx). The type of monomers incorporated in the polymer chain are determined by the organism used, the culture conditions and the carbon source provided [72]. P(3HB) is the most studied member of this family and exhibits high crystallinity, being organized in a compact right handed helix stabilised by carbon–methyl group interaction. This makes it a brittle and stiff material. The introduction of long side chains brings to disturbances in the regularity of crystalline lattice leading to an amorphous polymer structure. The incorporation of copolymers like P(HHx) to P(HB) does not break the helical conformation but introduces irregularities that lead to amorphous regions within the lattice [72].

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PHAs, being biological polymers, have excellent biocompatibility. P(3HB) has been found to naturally occur in the cellular membrane of animals and in human blood, as low molecular weight polymer [34]. The degradation products of PHA released during breakdown in vivo are the hydroxyl acids of the polymer backbone. These degradation acids have proven to cause a significantly lower inflammatory response due to a lower acidity compared to numerous synthetic biodegradable polymers used in tissue engineering applications, as for example PLA or PGA [34], [78]. They can be eliminated from the body by the Krebs cycle in merely 35 minutes [34]. Furthermore 3-hydroxybutanoic acid (3HB) is transformed in a natural human ketone body occurring in different parts of the human body like brain, heart, lung, liver, kidney and muscle. It is found in relatively high concentrations also in the blood of healthy adults (3–10 mg per 100 ml) [78]. Although P(3HB) exhibits high brittleness, poor processability and slow degradation, it has been investigated for various applications in the repair of bone, nerve, the cardiovascular system, or urinary and gastrointestinal tract defects [34]. Table 1.4 shows medical applications of PHAs. Copolymerization with 3HV, 4HB, or 3HH has shown great efficiency to overcome the inferior mechanical properties of P(3HB) [79] (Table 1.5).

Table 1.4 PHAs used in biomedical applications [73]. Modified/unmodified

PHAa Physical properties Medical application

P3(HB) Subcutanous patches

PHBV Subcutanous patches

PHBV

Design of a 3D microfibrous material-formed by the blend and electrospun

into fiber materials Myocardial patch

P3(HB) Peripheral nerve guide

PHBHHx Flexible

Vessel stent, hemocompatibility and cytocompatibility

PHBHHx Porous tube form, flexible Nerve conduit Fibronectin and alginate

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PHBV-PLGA

Turning a porous micropatterned film into a tube wrapping aligned

electrospun fibers Nerve guide

PHB Asymetric patches Stomach wall patch

PHBHHx Scaffold, flexible Bone regeneration

PHBV

Bone regeneration, better performances on attachment, proliferation, and differentiation of osteoblasts

HA reinforced PHB Bone regeneration

HA reinforced PHBV Bone regeneration

PHBHHx

Porous three dimensional scaffolds,

flexible Cartilage proliferation

PHBV Chondrocyte seeded Cartilage proliferation

PHB/PHBHHx

Human adipose tissue embedded

scaffolds Cartilage proliferation

PHB microspheres Solvent evaporation technique

Drug delivery, chemoembolization PHBV and

P(3HB-co-4HB) Semicrystalline Drug delivery

PHB homopolymers

Unable to entrap the drug because of its high melting temperature and

rapid crystallization rate Not available for drug delivery PHO

Autologous ovine endothelial cell

seeded Pulmonary valve

PHO

Autologous ovine endothelial cell

seeded porous PHO patches Pulmonary heart valve

a P3(HB) or PHB, poly(3-hydroxybutyrate); PHBV, poly(3-hydroxybutyrate-co-3-hydroxyvalerate);

P4(HB), poly(4-hydroxybutyrate); PHBHHx, poly(3-hydroxy;butyrate-co-3-hydroxyhexanaote); PHO, poly(3-hydroxyoctanoate); PLGA, poly(lactic-co-glycolic acid); HA, hydroxyapatite.

Table 1.5 Thermal and mechanical properties of PHAs, their copolymers and blends [34].

Polymera Mw Tg (°C) Tm (°C) Young's modulus (GPa) Ultimate tensile strength (MPa) Elongation at break (%) P3HB 9 175 3.08 45 4 4 180 43 5 4 185 62 58 370000 2.05 36 2.05 10 178 2.07 36 2 P3HB-co-7%3HV 450000 1.04 24 2.08 P3HB-co-11%3HV 529000 1.01 20 17 P3HB-co-11%3HV 2 157 3.07 38 5 P3HB-co-19%3HV 1.05 18 25 P3HB-co-22%3HV 227000 0.62 16 36 P3HB-co-20%3HV –5 114 1.09 26 27

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–1 145 20 50 P3HB-co-28%3HV –8 102 1.05 21 700 P3HB-co-34%3HV –9 97 1.02 18 970 P3HB-18%4HB – 22 137 0.32 12 1120 P3HB-co-16%4HB –7 150 26 444 P4HB – 47 61 0.23 36 1140 P4HB 0.05 52 1000 P3HB-co-2.5%3HH 0.62 26 7 P3HB-co-7%3HH 0.29 17 24 P3HB-co-9.5%3HH 0.155 8.8 43 P3HB-co-10%3HH –1 127 21 400 P3HB-co-6%3HD –8 130 17 680 P3HO-3HH 0.010 45 198

a P3HB, poly(3-hydroxybutyrate; P3HB-co-3HV, poly(3-hydroxybutyrate-co-3-hydroxyvalerate);

P(3HB-co-3HH), poly(3-hydroxybutyrate-co-3-hydroxyhexanoate); P(3HB-co-4HB), butyrate-co-4-hydroxybutyrate); P4HB, poly(4-hydroxybutyrate); P3HB-co-3HD, poly(3-hydroxy-butyrate-co-3-hydroxydecanoate); P3HO-3HH, poly(3-hydroxyoctanoate-co-3-hydroxyhexanoate);

1.2.2.2.1 Poly-3-hydroxybutyrate-co-3-hydroxyhexanoate (PHBHHx)

As mentioned above PHBHHx is the copolymer of 3HB and 3HHx monomer units showing improved physical and chemical properties compared with PHB and PHBV [80]. The “x” in PHBHHx represents the amount of hydroxyl hexanoate present in the copolymer. PHBHHx showed excellent elastic flexibility with HHx content higher than 5–10% [34] due to the inability of the HHx side chains to fit into the crystalline lattice of PHB, creating amorphous regions between the crystalline regions. Furthermore PHBHHx has demonstrated improved hemocompatibility and cytocompatibility in vitro [81]. Qu et al. [81] suggested it potential as a blood-contact material with less platelet adhesion, reduced erythrocyte blood-contact and hemolysis reactivity compared with PHB and PHBV films. Furthermore the same group investigated the possibility to develop tissue engineered blood vessel of PHBHHx and found that PHBHHx films containing high HHx content promoted smooth muscle cells’ differentiation [82] making them suitable for blood vessel

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tissue engineering. Finally, in vivo studies demonstrated a particularly mild tissue response of PHBHHx during 6 months subcutaneous implantations in rabbits while PHB and PLA showed a relative acute immunological reactions [83].

PHAs have been widely used in cardiovascular tissue engineering, for example artery augments, cardiologic stents, vascular grafts, heart valves, pericardial patches, implants, dressing tablets and microparticulate carriers [73], [84], [85]. Kenar et al. [86] reported the fabrication of a myocardial patch to replace myocardial infarcts via electrospinning of poly(3-hydroxybutyrate-co-3-hydroxyvalerate) (PHBV), poly(L-D,L-lactic acid) (P(L-D,L) LA) and poly(glycerol sebacate) (PGS). The electrospun aligned fiber mats seeded with mouse mesenchymal stem cells (MSCs) were effectively stacked to form a 3D microfibrous construct and a thick myocardial patch with structure similar to the native tissue was obtain.

Yang et al. [87] developed a cardiovascular scaffolds modified by silk protein. They obtained a porous implant with less inflammation reactions and a more hydrophilic surface. The silk-modified PHBHHx scaffolds exhibited better cell adhesion and proliferation compared with PHBHHx with better biocompatibility for human fibroblasts, smooth muscle cells, human umbilical vascular endothelial cells, and endothelial- like cell line-ECV304.

Song et al. [88] investigated intravascular biocompatibility of PHBHHx in vivo using the polymer as coating for a decellularized xenogenic vascular matrix scaffolds. The hybrid patches were implanted into the abdominal aorta of New Zealand rabbits and removed after 1, 4 and 12 weeks after implantation. The hybrid patches showed complete histologic restitution of vessel tissue, as well as a confluent endothelial surface after 12 weeks of implantation. PHBHHx showed remarkable intravascular biocompatibility and was suggested as candidate for lumen in cardiovascular tissue engineering.

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1.2.3 Electrospinning

Electrospinning has gained enormous attention in tissue engineering particularly for its ability to spin fibrous matrices that mimic the highly fibrous nature of some extracellular matrices [89]–[92]. Electrospinning allows the formation of non-woven mats of ultrafine fibers of a number of different materials. The nanometer-size diameter of the electrospun fibers resembles that of the various proteins of the fibrous extracellular matrix in the body, in particular collagen microfibrils. Matrices with extensive surface to volume ratio can be obtained having consequently substantial porosity and enhanced mechanical performances [90]. The electrospun structure increases cell attachment and penetration due to the high porosity and the above mentioned similitude to fibrous extracellular matrices. Furthermore electrospinning has been shown to be a quite simple and cost-effective technique with the possibility of large scale production and the availability to produce different morphological structure due to the existence of various fabrication modes such as coaxial electrospinning [91], [93].

Electrospinning process consists in spinning a polymer solutions (or melt) under a high electrical field applying an electrostatic potential between a spinneret and a collector. The charge induced on the polymer results in charge repulsion within the solution and the droplet exiting the tip is deformed by this electrostatic force and the surface tension forming a cone (known as Taylor cone). Ultimately, the electrostatic force overcomes the surface tension and an electrically charged fiber jet is initiated which is drawn towards the collector. The jet undergoes a stretching and whipping process and evaporation of the solvent resulting in a randomly oriented, non-woven fiber mat on the collector. However, if the polymeric fluid has insufficient macromolecular entanglements it undergoes Rayleigh instability which break the jet down into droplets [89], [91], [92].

Extensive literature is available on electrospinning of PCL, which is surely one of the most commonly used polymers in electrospinning applications in the past

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decade [58], [59]. On the other hand, electrospinning of PHBHHx was first reported by Cheng et al. [94] only rather recently. They demonstrated that electrospinning of PHBHHx was possible and by varying electrospinning parameters such as polymer solution concentration, co-solvent weight ratio, and applied voltage they were able to customize fiber diameters. Shortly after Ying et al. [78] reported successful electrospinning of various PHA copolymers like poly((R)-3-hydroxybutyrate-co-5mol%-(R)-3-hydroxyhexanoate), co-7mol%-4-hydroxybutyrate) and poly((R)-3-hydroxybutyrate-co-97mol%-4-hydroxybutyrate). They investigated biocompatibility and bioabsorption of the fiber mats by subcutaneous implantation in rats. The implants were well tolerated in vivo with only mild tissue response during the 4 and 12 weeks of implantation while bioabsorption rate was highly dependent on the monomer content. A later study by Cheng et al. [95] reported the successful fabrication of electrospun fibrous blends of PHBHHx/PDLLA with tunable mechanical strength and degradation rate showing the possibility of blending PHBHHx with other polymers to modify each other’s properties.

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1.3 Research objectives

The aim of this thesis was to investigate and develop PCL/PHBHHx scaffolds for cardiac tissue engineering applications using electrospinning as fabrication technique with the aim of enhancing the mechanical properties and ameliorating the cytocompatibility.

The specific aims included:

• determination of the electrospinning parameters to produce smooth and uniform fibers of the PCL/PHBHHx blends

• investigation of the miscibility of PCL/PHBHHx blends

• investigation of the mechanical properties of various electrospun PCL/PHBHHx scaffold.

• investigation of the degradation behavior of the electrospun fiber mats in saline buffer solutions

• investigation the cell-material interaction of the various electrospun PCL/PHBHHx scaffolds.

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C

HAPTER

2

2 Materials and Methods

All chemicals were obtained from Sigma-Aldrich, Germany unless otherwise stated. Poly(ε-caprolactone) (PCL) of Mn = 70.000–90.000 was obtained from Sigma-Aldrich, Germany. Poly(3-hydroxybutyrate-co-3-hydroxyhexanoate) (PHB-HHx) (86% 3-hydroxybutyrate content, and 14% 3-hydroxyhexanoate content, Mw =102.000 g) was received as a gift from Dr. José M. García-García (Consejo Superior de Investigaciones Cientíﬁcas (ICTP-CSIC), Madrid, Spain). Dichloro-methane (DCM) and methanol were used as the solvent system for the polymers.

2.1 Sample fabrication

2.1.1 Fabrication of PCL/PHBHHx films

2% w/v solutions of varying blend proportions were prepared by dissolving a combined mass of 50 mg of PCL and PHBHHx in 3 ml dimethyl carbonate at room temperature. The polymer solution was stirred overnight and films of the blends were obtained by solvent casting into petri dishes. The petri dishes were covered with perforated aluminum foil and placed under a fume hood at room temperature to allow uniform evaporation of the solvent. The blends prepared in this way ranged from 100/0 to 20/80 weight ratio of PCL/PHBHHx.

2.1.2 Fabrication of PCL/PHBHHx electrospun fibers

PCL/PHBHHx fibers were prepared by electrospinning technique. PCL and PHBHHx were mixed in weight ratios of 100/0, 70/30 and 30/70 of varying concentrations

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ranging from 5 to 15 wt % and dissolved in a mixture of D

blends were stirred overnight to ensure a homogenous solution. The blend solutions were fed into a 10 ml glass syringe and fitted with a metal needle used as noozle. During the electrospinning process an electrical syringe pump (capa from 0.25 to 143 cc/hr) regulated the flow rate of the solution. A high dc potential difference was applied to the needle

power generator HVG-P60-R

were collected onto a sheet of aluminum foil wrapped around a fixed L collector at defined distance. The solutions were electrospun co

approximately 1.5 hours. The process was performed at room temperature and in air without humidity control. The tip to collector distance, the voltage and the flow rate were adjusted until a stable electrospinning process and bead

obtained (Table 2.1). The typical distance ranged voltage (15 and 20 kV) while the

shows the employed electrospinning setup

Figure

wt % and dissolved in a mixture of DCM/methanol 70/30. The blends were stirred overnight to ensure a homogenous solution. The blend ml glass syringe and fitted with a metal needle used as noozle. During the electrospinning process an electrical syringe pump (capa

cc/hr) regulated the flow rate of the solution. A high dc potential difference was applied to the needle using a high voltage generator (High Voltage R-EU; Linari Engeneering s.r.l.). The obtained fibers ollected onto a sheet of aluminum foil wrapped around a fixed L

collector at defined distance. The solutions were electrospun consecutively for . The process was performed at room temperature and in . The tip to collector distance, the voltage and the flow rate were adjusted until a stable electrospinning process and bead free fibers were (Table 2.1). The typical distance ranged between 15 and 20 cm as well as

while the flow rate was between 1-5 cm3/h. Figure spinning setup.

Figure 2.1: Electrospinning device

CM/methanol 70/30. The blends were stirred overnight to ensure a homogenous solution. The blend ml glass syringe and fitted with a metal needle used as noozle. During the electrospinning process an electrical syringe pump (capacity cc/hr) regulated the flow rate of the solution. A high dc potential a high voltage generator (High Voltage EU; Linari Engeneering s.r.l.). The obtained fibers ollected onto a sheet of aluminum foil wrapped around a fixed L-shape nsecutively for . The process was performed at room temperature and in . The tip to collector distance, the voltage and the flow fibers were cm as well as Figure 2.1

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Table 2.1: Electrospinning parameters investigated for the fabrication of PCL/PHBHHx fibers

2.2 Characterization studies

2.2.1 Morphology studies

Initial observations of the collected fiber mats were carried out using an optical microscope (LEICA; Wetzlar, Germany) to assess its morphological aspect such as bead formation. Detailed study of the morphology and aspects of surface topographies were analyzed using a scanning electron microscope (LEO 435VP SEM, Germany). The fibrous samples were fixed on a sample holder with an electrically conductive adhesive and sputtered with graphite prior to observation. All samples were observed at magnifications between 500x and 5000x.

2.2.2 Fiber diameter and porosity

The average fiber diameter and pore size was also calculated from the SEM images using a computer image analyzer (ImageJ). Approximately more than 100 random fibers and 100 random pores at various points were selected in order to determine fiber diameter distribution and pore size distribution of the fiber mats. A SEM image with magnification 4000x was used for each specimen.

The porosity (theoretical value)of the samples was calculated according to the equation:

Polymer blend Blend proportions Voltage (kV) Distance (cm) Flow rate (cm3/h) Concentration (% w/v) PCL/PHBHHx 70/30 70/30 12–20 12–20 1–7 5–15 PCL/PHBHHx 30/70 30/70 12–20 12–20 1–5 5–15 PCL/PHBHHx 100/0 100/0 12–20 12–20 1–5 5–15

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where is the density of the bulk material

material. Bulk density was calculated by cutting the samples into 4 weighing them with an analytica

measured by a digital micrometer for determining the bulk volume.

2.2.3 Water contact angle Surface wettability was studied (DSA30, Kruess GmbH Germany)

The sessile drop method was applied to determine the static water contact angle. Ten droplets of 10 µl of distilled water

positions of each sample (Figure 2.2). Av

sample. The contact angle was directly calculated from the projected image of the camera to the attached computer.

Figure 2.2: Drop deposition on a electrospun fibrous sample 1

100%

is the density of the bulk material and is the true density of the . Bulk density was calculated by cutting the samples into 4 x 0.5

analytical balance. The thickness of the samples was measured by a digital micrometer for determining the bulk volume.

Water contact angle

urface wettability was studied by means of a contact angle measuring instrument Germany) at room temperature and in an air atmosphere The sessile drop method was applied to determine the static water contact angle.

µl of distilled water were randomly deposited at dif positions of each sample (Figure 2.2). Average values are reported for each sample. The contact angle was directly calculated from the projected image of the camera to the attached computer.

Drop deposition on a electrospun fibrous sample

true density of the .5 cm2 and l balance. The thickness of the samples was

instrument at room temperature and in an air atmosphere. The sessile drop method was applied to determine the static water contact angle. at different erage values are reported for each sample. The contact angle was directly calculated from the projected image of the

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2.2.4 Fourier transform infrared attenuated total reflectance (FTIR-ATR) spectroscopy

FTIR-ATR spectra were obtained using an FTIR spectrometer (Nicolet, USA) with 4 cm-1 resolution in attenuated total reflectance mode. Spectra were collected in the mid-infrared (MIR) region which includes a wavenumber range between 4000 to 530 cm-1. Due to inhomogeneous sample thickness no quantitative analysis was performed.

2.2.5 Mechanical properties

2.2.5.1 Static mechanical characterization

Statistical mechanical properties were evaluated by tensile testing using Zwick Z050 tensile materials testing machine (Germany) (Figure 2.3). Four individual samples were assessed for each blend. The samples were cut into rectangular specimen of approximately 40 x 5 mm. Thickness was measured for each specimen separately with a digital thickness gauge in 4 different points as it varied from 3.5 to 5.8 mm.

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All experiments were carried out at room temperature. The tensile testing machine was equipped with a load cell of 50 N. A pretension of 0.1 N and a stretching speed of 10 mm/min were applied. A stress-strain curve was obtained from the load and displacement measurements. The tensile modulus (Young’s modulus) was evaluated as the slope of the stress-strain curve in the initial, linear portion of the curve. Furthermore the tensile strength and the elongation to break were determined.

2.2.5.2 Dynamic mechanical thermal analysis (DMTA)

Viscoelastic measurements were performed in tensile mode with a dynamic mechanical thermal analysis instrument (Rheometric Scientific DMTA IV). All measurements were carried out at room temperature and in air. Rectangular fiber mats of 20 x 5 mm size were assessed. Measurements were performed at changing frequency and changes in storage modulus (E’), loss modulus (E”) and tanδ were recorded. Frequency was varied between 0.1 and 30 Hz while a constant strain of 15 g was applied. Results presented are average values of 3 repeated measurements for each sample.

2.2.6 Thermal properties

Thermal analysis was carried out using Differential Scanning Calorimetry (DSC Q2000 V24.10 Build 122). Small samples were cut from the fiber mats to obtain specimen with an approximate weight of 6 mg. Samples were heated from -90 °C to 150 °C with a heating rate of 10 K/min and heating-cooling-heating scans were performed. Melting temperature as well as glass transition temperature and enthalpy of fusion were obtained. Thermo gravimetric analysis (TGA) was carried out using a thermoanalyzer (2950 TGA V5.4A) operated in air at a heating rate of 10 °C/min. The samples were heated from 30 °C to 600 °C.

Fabrication and characterization of fibrous poly(ε-caprolactone)/poly(3-hydroxybutyrate-co-3-hydroxyhexanoate) scaffold for cardiac patch application