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UNIVERSITY OF MINHO DEPARTMENT OF TEXTILE ENGINEERING

Development of Hybrid Braided Vascular Prostheses

Master of Science in Textile Engineering

by AYLIN TEKIN July 2009

Supervisor PROF RAUL FANGUEIRO

Association

of

Universities for Textiles

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ACKNOWLEDGEMENTS

I would like to express my gratitude to my supervisor Prof. Raul Fangueiro for offering me an

opportunity to work in his lab and for his support and guidance in my research.

My special thanks to Antonio Souza Freitas for spending time to make the braiding of tubular

structure possible.

I sincerely thank to Mauricio Malheiro for helping me with my tests and arranging me a time in his

busy schedule, whenever I needed.

I thank the members of the laboratory of environmental chemistry and laboratory of research in

dyeing,Vergina Pinto, Frederico Maia, Sandra Sampaio for letting me use their equipment for coating

the samples and for providing me a friendly atmosphere.

I express my sincere thanks to Prof. Sam Hudson for sharing his deepest knowledge and encouraging

me via emails.

I would like to thank Erkan Turker Baran for his support and supplying the solutions needed to

complete my work.

I appreciate the help given by Katrien Hooreman during my two years studies.

My sincere appreciation goes to Steve and Lynne Robson for being there for me and to all my friends

whom made my time more enjoyable during my stay in Guimaraes.

I am deeply grateful to my parents and my sister for their unconditional love and encouragement.

Without their support, it would not be easy for me to complete my study abroad.

Aylin Tekin

2009

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Copyright : The author gives admission to make this Master’s thesis available for consultation and to copy parts of the

Master’s thesis for personal use.

Any other use falls under the limitations of the copyright, especially with regard to the obligation of mentioning the source

explicitly on quoting the results of this Master’s thesis.

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TABLE OF CONTENTS

ACKNOWLEDGEMENTS.........................................................................................ii

LIST OF TABLES....................................................................................................vii

LIST OF FIGURES.................................................................................................viii

LIST OF SYMBOLS or ABBREVIATIONS..................................................................x

ABSTRACT............................................................................................................xi

CHAPTER 1 INTRODUCTION..................................................................................1

1.1 Background..............................................................................................1

1.2 Objective..................................................................................................2

1.3 Methodology............................................................................................2

1.4 Thesis Disposition.....................................................................................2

CHAPTER 2 STATE OF ART

2.1 Introduction............................................................................................3

2.2 The Morphology of Blood Vessels..........................................................4

2.3 Atherosclerosis.......................................................................................6

2.4 Treatments.............................................................................................6

2.5 Drawbacks of Current Grafts..................................................................6

2.6 Vascular Prostheses................................................................................7

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2.6.1 Types of Vascular Prostheses........................................................7

2.6.1.1 Autografts..........................................................................7

2.6.1.2 Synthetic Prosthetic Grafts................................................8

2.6.1.3 Biological Prosthetic Grafts...............................................9

2.6.1.4 Tissue Engineered Vascular Graft....................................10

2.7 Desirable Properties of Vascular Prostheses........................................11

2.8 Polymeric Biomaterials.........................................................................12

2.8.1 Teflon..........................................................................................12

2.8.2 Polyester......................................................................................13

2.8.3 Nylon...........................................................................................13

2.8.4 Polyurethane Elastomer..............................................................14

2.9 Fabric Construction..............................................................................16

2.9.1 Woven.........................................................................................19

2.9.2 Knitted.........................................................................................20

2.9.3 Braided........................................................................................21

2.9.4 Nonwoven...................................................................................21

2.10 Finishing..............................................................................................22

2.11 Testing and Evaluation.......................................................................22

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CHAPTER 3 MATERIALS AND PROCESSES...........................................................24

3.1 Material Characterization.....................................................................24

3.1.1 Mechanical Properties of PET and PLA........................................24

3.1.2 Winding PET Yarns.......................................................................26

3.2 Braiding Technique...............................................................................26

3.3 Coating Process....................................................................................30

3.3.1 Preparation of CS Solution..........................................................31

3.1.2 Preparation of NaOH Solution.....................................................31

CHAPTER 4 RESULTS AND DISCUSSIONS.............................................................33

4.1 Fabric Characterization.........................................................................33

4.1.1 Uncoated Tubular Structures Fabric Characterization................33

4.1.2 Coated Tubular Structures Fabric Characterization....................34

4.2 Physical Properties of Tubular Structures............................................36

4.2.1 Uncoated Tubular Structures Physical Properties.......................36

4.2.1.1 Change in Braiding Angle................................................37

4.2.1.2 Change in Crimp Distance...............................................38

4.2.1.3 Change in Crimp Height...................................................39

4.2.1.4 Change in Wall Thickness................................................39

4.2.1.5 Change in Mass...............................................................40

4.2.2 Coated Tubular Structures Physical Properties...........................41

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4.3 Comparison of Coated and Uncoated Tubular Structures....................43

4.4 Evaluating the Influence of Coating and Washing Time.......................44

4.5 Evaluating Chitosan Adhesion to Braided Fabric Surface ..................45

4.5.1 SEM Analyses..............................................................................45

4.5.2 A Direct-Staining Method............................................................45

CHAPTER 5 CONCLUSIONS AND FUTURE RECOMMENDATIONS........................49

5.1 Summary...............................................................................................49

5.2 Limitation of Approach.........................................................................50

5.3 Future Recommendations....................................................................50

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LIST OF TABLES

Table 2.1: A comparison of fabric formation technique 21 Table 2.2 Sample test methods for large diameter textile graft 23 Table 3.1 The average of force-extension values of PET and PLA 26

Table 3.2 The effect of different coating and washing times on 100% PET 31

braided

Table 3.3 The effect of coating time on 100%PET and 75%PET-25%PLA 32

samples

Table 4.1 The structural characteristic of uncoated 100% PET samples 36

Table 4.2 The structural characteristic of uncoated 75%PET-25%PLA 37

samples

Table 4.3 Structural characteristic of coated 100% PET samples 41

Table 4.4 Structural characteristic of coated 75%PET-25%PLA samples 41

Table 4.5 Change in mass and length after coating 44

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LIST OF FIGURES

Figure 2.1 Structure of arteries 4

Figure 2.2 Structure of a muscular artery 5

Figure 2.3 Tissue engineered vascular grafts 10

Figure 2.4 Constituent elements of medical textile products 16

Figure 2.5 Examples of fibrous materials developed for use in 17 medicine

Figure 2.6 Fabric techniques 18

Figure 2.7 Examples of woven, knitted, and braided structures 18

Figure 2.8 Fabric and vascular grafts made using a Leno weave 19 Figure 2.9 Woven and knitted fabrics used for vascular grafts, 20 showing differences in porosity Figure 2.10 Vascular graft finishing operation 22 Figure 3.1 Force-extension curve for 68.2/2 tex PET multifilament 25 Figure 3.2 Force-extension curve for PLA monofilament 25 Figure 3.3 Braiding Machine 27 Figure 3.4 Braiding Process 28 Figure 3.5 Oven 28 Figure 3.6 Braided tubular structures 29 Figure 3.7 3% w/v CS solution 30

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Figure 3.8 3% w/v CS Solution; 0.5 % NaOH; Methanol 31 Figure 4.1 Uncoated 100% PET samples (200x) 33 Figure 4.2 Uncoated samples 75% PET-25%PLA (200x) 34

Figure 4.3 Coated samples 100% PET (200x) 35

Figure 4.4 Coated samples 75% PET-25%PLA (200x) 35

Figure 4.5 Variation of the braiding angles according to the diameter 38

Figure 4.6 A comparison of crimp distance 38

Figure 4.7 A comparison of crimp height 39

Figure 4.8 A comparison of wall thickness 40

Figure 4.9 A comparison of mass according to diameter variation 40

Figure 4.10 Effect of Coating on Wall Thickness 42

Figure 4.11 Coated and Uncoated 100% PET d: 6 (200x) 43

Figure 4.12 Coated and Uncoated75% PET-25%PLA d: 5 (200x) 43

Figure 4.13 The effect of coating and washing time with 0,5 % NaOH 44

on the quantity of chitosan coated. ( 100% PET d:6 )

Figure 4.14 Micrographs of Coated 100% PET Braided Structure 46

Figure 4.15 Micrographs of Coated 75/25 PET/PLA Braided Structure 47

Figure 4.16 Stereoscope images of coated and uncoated structure 48

were taken after treatment with Eosin. (200x)

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LIST OF SYMBOLS or ABBREVIATIONS

PET: Polyester

PLA: Poly-lactic acid

ePTFE: Poly-tetrafluoroethylene

FDA: Food and Drug Administration

ASTM: American Standart Test Method

AAMI: Association for the Advancement of Medical Instrumentation

ANSI: American National Standards Institute

ISO: International Organization for Standardization

CS: Chitosan

Group 1: Samples were produced using 100% PET yarn

Group 2: Samples were produced using 75%PET yarn, 25%PLA yarn

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ABSTRACT

This study has been focus on designing and producing braided graft having a biostable and

biodegradable material. Braided tubular structure with diameters ranging from 5 to 8 mm are

developed using a horizontal braiding machine according to a patented braiding process. The

material used was Polyester biostable yarn with 68.8/2 tex, Polylactide biodegradable yarn with 6.3

tex. The braided structure was coated with chitosan (medium molecular weight ~ 0.3 x 105,

desacetylation of 90%).

The experimental work involved the braiding and coating process and the fabric characterization of

tubular structures. Physical properties of the fabrics, the change in wall thickness and mass after

coating and the effect of coating time were investigated. Finally, the morphology of the coated

surface was analyzed using scanning electron microscopy.

KEY WORDS: coronary artery disease, vascular prostheses, braiding, tubular fabrics, chitosan.

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CHAPTER 1

INTRODUCTION

1.1 Background

Atherosclerosis is a disease of arteries that causes more death and disability than any other disease

in the industrialized world, more than all types of cancer combined.

Vascular grafts are used to replace, bypass or maintain function of damaged or diseased blood

vessels in small, medium and large diameters. [1]

The ideal cardiovascular bypass graft requires a broad range of characteristics including strength,

viscoelasticity, biocompatibility, blood compatibility, biostability and also needs to have the

capability to adapt to the prevailing hemodynamic conditions both immediately and in the long term.

[2] The degree of host tissue infiltration into the biomaterial depends on the pore size, surface

texture, anatomical location and the material’s biocompatibility. [3] However, the two main

problems that were found with the biomaterials were; the difficult task of making biocompatible

materials that heal the body and making materials that can be punctured and have quick sealing

rates.[4] Therefore, the objective of this thesis is to choose a biomaterial, design a method for

manufacturing a tubular vascular graft which eliminates the suturing process and treat with chitosan

in order to match the tissue’s mechanical properties, its topography and compatible with the

host tissue.

The tubular structures with a range of diameters are manufactured with 100% polyester yarn and

%75 PET (polyester) – 25% PLA (polylactide) yarns by using available horizontal braiding machine

according to a patented braiding process. The fabricated grafts are analyzed and the physical

properties are measured and compared. Since the surface modification of vascular grafts becomes

important when they come into contact with physiological components such as blood and living

tissues, the samples were coated with chitosan.

The final properties of the resultant fabricated grafts are measured.

Several tubular structures were manufactured and characterized to obtain an average result.

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1.2 Objective

The main objectives of this study were to produce tubular braided structures, investigate their

properties and coat the structures with chitosan in order to use them in vascular prostheses

application.

1.3 Methodology

The study was conducted within several tasks:

The first task was literature survey, in order to collect and determine all scientific information related

to vascular prostheses.

The second task was the experimental work which included selection of materials, braiding and

coating process of tubular fabrics.

The third task was fabric characterization, surface analysis of uncoated braided structure, evaluate

the physical properties and examine the coated structure in order to control the homogeneity of

chitosan film between filaments.

The conclusion from the data analysis fulfilled the last task.

1.4 Thesis Disposition

The thesis comprises five chapters:

Chapter 1 is the INTRODUCTION, where the general information about the topic is given. The

problem in this field defined in background, the main objective , the way of work realization and the

organization of the thesis were mentioned.

Chapter 2 presents the STATE-OF-ART where the latest studies given in the field of vascular

prostheses. The morphology of the vessels, the type vascular prostheses, the materials and the

techniques that has been used to manufacture the vascular grafts was described. The drawbacks of

current vascular prostheses were mentioned.

Chapter 3 describes the EXPERIMENTAL WORK of the thesis where the materials and the methods

were stated. Selection of the materials, preparation of the yarns and manufacturing of the braided

structure and the coating process were explained.

Chapter 4 is RESULTS and DISCUSSIONS where the results of the test evaluated. The characterization

of the braided fabrics was explained according to vision analyses.The physical properties of the

braided fabrics, the quality of the coating process and the comparison of the coated and uncoated

structure were discussed.

Chapter 5 give the CONCLUSION and the FUTURE RECOMMENDATIONS of the thesis with the main

stress on the experimental work.

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CHAPTER 2

STATE OF ART

2.1 Introduction

The purpose of the vascular graft prosthesis is to replace, repair, or otherwise correct a damaged or

diseased blood vessel. To maximize the effectiveness, the prosthesis should be able to fully replicate

the vascular tissues’ regenerative properties. The major differences between natural vessels and the

prosthetic grafts may be caused by the absence of an endothelial monolayer on the graft surface.

Especially, the small diameters (< 6 mm diameter) synthetic vascular grafts have a failure rate of 53%

after 4 year. [5] In order to increase the long-term patency rates, the biomaterial surface should be

more compatible with the human body and present similar chemistry, morphology and mechanical

properties to cell surfaces and have long-term mechanical stability.

Newer synthetic prostheses have enabled successful reconstructions in large diameter, high flow

vessels, such as the aorta and its primary intra-thoracic and abdominal branches. However, the

success of medium and small size arteries has been low. The ideal desirable features of vascular

prostheses are handling, performance, durability and cost considerations. The graft must be pliable

yet resist to excessive dilation and ideally stimulate the viscoleastic properties of the vessel. It must

be sterilizable and resistant to infection. The flow surface should be maximally thrombosis-resistant

and allow generation of a thin pseudo-intima or complete spontaneous re-endothelialization . A

controlled inflammatory and healing response should occur to the bio-chemically stable and

relatively inert prosthetic material to allow tissue incorporation.

However, there are limitations to present understanding of the blood and tissue material

interactions affecting thromboresistence, flow surface healing and tissue incorporation. The hyper-

plastic response of the blood vessel constitutes the greatest long-term problem leading to failure of

heart valves and vascular grafts. The new tissue is often particularly liable to the pathological

changes that characterize arteriosclerosis and thereby result in the production of a highly

thrombogenic surface. In this manner, graft failure may be consequence of a tissue overgrowth

reaction not predicted by the previous tests used to evaluate biomaterials, or by the flow

characteristics in the region of the implant, nor by the impact of biomaterial upon haemostatic

reactions.

The two main problems that were found by practical application with the biomaterials were:

the difficult task of making biocompatible materials that heal the body;

making materials that can be punctured and have quick sealing rates. [4]

Improving the prostheses’ chemical and mechanical properties and the construction of such grafts

are the requirements for ideal vascular prostheses. In order to find greater compatibility between

implants and their hosts, and solution for durability, an investigation of the materials’ properties and

their host responses and the structure of the arteries are essential.

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2.2 The Morphology of Blood Vessels

Arteries are essentially high-pressure pipes that carry blood pumped by the heart out to all of the

tissues of the body, such as muscle, skin, bone, liver, kidneys, etc. The arterial system, shown in

Figure 2.1, resembles a tree with a large main trunk - the aorta - and progressively smaller arteries

branching out from the aorta.

Figure 2.1- Structure of arteries [source: http://biology.about.com/library/organs/heart/blarteries.htm]

The thick-walled aorta, somewhat less than an inch in diameter, emerges from the top of the heart,

gives off branches that supply the head and arms and then curves downward, running just in front of

the spine, to supply the trunk and internal organs, and eventually splits into two large arteries that

supply the legs. After many stages of branching, the tiniest arteries become capillaries - the smallest

blood vessels. In the capillaries, oxygen and nutrients in the blood are delivered to tissues, and waste

products are picked up. Then capillaries join together to form tiny veins, which join together many

times again to form larger veins. Veins are thin-walled, low-pressure blood vessels, which bring blood

from the tissues back to the heart.

At the very beginning of the aorta, just beyond the heart valve that separates the pumping chamber

of the heart from the aorta, the left and right coronary arteries branch off to supply the heart muscle

itself. The left main coronary artery splits almost immediately into two branches, called the left

anterior descending coronary artery and the left circumflex coronary artery. Coronary artery disease

due to atherosclerosis is often described as 1-, 2-, or 3-vessel disease, depending on whether major

blockages are found in various combinations of the right, left anterior descending, and left circumflex

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coronary arteries and their branches. The most important coronary artery is the left anterior

descending artery, which supplies the front side of the heart as well as the muscular septum that

separates the right and left pumping chambers. [6]

All arteries are comprised of three distinct layers: intima, media and adventitia, the proportion and structure of each varies with the size and function of the particular artery. Most prominent is the middle layer, called the media, which consists of tightly packed smooth muscle cells, fibrous tissue proteins such as collagen and elastin, and gel-forming proteoglycans. The inner tissue layer is the intima, which is a much looser structure with fewer cells, less elastin, and considerably more open spaces between tissue components. The outer tissue layer, called the adventitia, is also a relatively loose tissue consisting mostly of bundles of collagen along with a few connective tissue cells. A large artery, like aorta, shown in Figure 2.2, is comprised of the following layers, going from the lumen to the most external ones:

1. The intima, or intermost layer, consists of a layer of endothelial cells separated from the inner layer by a narrow layer of connective tissues which anchors the cells to the arterial wall.

2. A large layer of elastic fibres forming the elastica interna layer. 3. Below this layer are concentric waves of smooth muscle cells intermixed with elastic fibres.

Elastic lamellae and smooth muscle cells are imbedded in a ground substance rich in proteoglycans. Proteoglycans are formed of disaccharides bound to protein and serve as binding or "cement" material in the interstitial spaces. The outer layer of the media is penetrated by branches of the vasa vasorum.

4. Between the smooth muscle layer and the adventitia, there is again another layer of elastic fibers, the elastica externa. Layers 2, 3 and 4 form the media.

5. The outer layer or adventitia is formed of irregularly arranged collagen bundles, scattered fibroblasts, a few elastic fibers and blood vessels which, because of their location, are called vasa vasorum or vessels of the vessels. [4,7]

Figure 2.2 - Structure of a muscular artery [source: Christiaan Hendrikus Gerardus Arnoldus van Oijen, Mechanics and design of fiber-

reinforced vascular prostheses]

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2.3 Atherosclerosis

Atherosclerosis is a disease of the arterial intima, which is the innermost layer of the arterial wall.

The intima, like the other arterial tissue layers, is a type of connective tissue. Connective tissues give

form and structure to the body, and keep the organs in place. Examples of connective tissues are the

deeper layers of skin, the fascia that separate muscle layers (think of "gristle" in meat), bones, and

tendons. Much of the strength in connective tissues comes from fibrous tissue proteins - elastin and

various kinds of collagen - which are located between the cells in the tissue and are laid down by the

cells. In the arterial intima, scattered smooth muscle cells manufacture the collagen, elastin, and

proteoglycans that form the bulk of intimal tissue. When injured, connective tissues form scars, and

some features of atherosclerosis are very similar to scar formation. [6]

Atherosclerosis is known by several other names - arteriosclerosis (though technically

arteriosclerosis also includes some other rare and minor arterial conditions), hardening of the

arteries, cholesterol deposits in the arteries, and arterial blockages. Coronary heart disease is the

result of atherosclerosis in the coronary arteries, which supply the heart muscle. Cholesterol carried

by lipoproteins in the blood enters the artery wall and builds up in enormous amounts, leading to

tissue damage, inflammation, and fibro proliferative scarring. Breakdown of tissue in the inner part

of the artery wall sets the stage for blood clots causing heart attacks and strokes. [8]

2.4 Treatments

Medications (nicotinic acid, cholestyramine etc) are usually the first step as treatment. Advance

studies in vascular surgery have found opportunities for treatment and prevention. Coronary bypass

operations, heart catheterization balloon and vascular graft procedures can relieve the pain of

blocked coronary arteries, but these operations and procedures do not prevent future heart attacks

very well at the moment. [6] This can be achieved through harvesting of blood vessels from the

patient’s own body (autogenous grafts) or through the use of manufactured prosthesis (artificial

grafts). The autografts can be harvested from the patient’s body, such harvesting can be a tedious

surgical task, does not always result in the best quality graft and is often not therapeutically

recommended. A tissue-engineered graft could fulfil the ideal characteristics present in an artery.

However, the great disadvantage of such a conduit is the time required for maturation leading to

unacceptable delays once the decision to intervene surgically has been made. Therefore synthetic

grafts are being developed for coronary artery bypass as well as for other blood vessel replacement.

[1]

2.5 Drawbacks of Current Grafts

The development of textile based vascular grafts of either synthetic or natural origin has been one of

the most important biomedical applications. This development has been facing significant progress

and allowed the reconstruction of obstructed or injured blood vessels with remarkable success.

However, some serious problems still remain unsolved. While homografts are still considered to be

preferred arterial replacement for small diameter vessels (below knee), they are used in limited

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quantity and are insufficient to meet the increasing needs for vascular replacement, due to mainly to

an inadequate supply, non-uniform properties and difficulty in preparation.[9]

Although the synthetic polymers make up by far the broadest and most diverse class of biomaterials,

many of them are chemically non-reactive in human and other animal bodies. Their lack of chemical

reaction, inertness and therefore biocompatibility may be attributed. [10] The synthetic materials

have been experimented and those that include rigid nonporous tubes made of gold, silver

aluminium, glass and polyethylene completely failed due to the lack of porosity and compliance.

These two important requirements of an ideal vascular graft have been partially fulfilled by the use

of fibrous based polymers and textile structure. [9]

The synthetic vascular grafts such as Dacron and expanded poly-tetrafluoroethylene (ePTFE) have

been successfully used in treating the pathology of large arteries( > 6 mm inner diameter ), however

the success is not proved for small diameter blood vessels. In medium size (6 mm diameter) and

small-diameter arteries (4 mm) bypass or replacement with no segments of textile grafts have

resulted in unacceptable lower long-term patency rates. [10] Patency rate is the performance of

vascular prostheses that is measured, amongst other parameters, by the length of time they remain

open. Mechanical matching of vascular grafts and host vessels are also important in determining the

graft patency rates. Improvement of the mechanical compatibility between natural arteries and

synthetic grafts may lead to better patency rate for synthetic grafts. Biomaterials that successfully

integrate into surrounding tissue should not only match the tissue’s mechanical properties, but also

its topography. [4]

2.6 Vascular Prostheses

One of the main uses for vascular grafts is in the treatment of arteriosclerosis in order to prevent the

reduction in vein size and blood flow. Vascular grafts provide the necessary space and strength

needed by the patients for the transfer blood from larger arteries to smaller veins.

According to the Food and Drug Administration (FDA) the approved uses for vascular grafts include

the repair, replacement and bypass of sections of native or artificial vessels. [4]

2.6.1 Types of Vascular Prostheses

Vascular grafts can be classified as: autogenous grafts, prosthetic grafts, biological prosthetic grafts

and tissue engineered vascular grafts.[5]

2.6.1.1 Autografts

In biology, autologous refers to cells, tissues or even proteins that are reimplanted in the same

individual as they come from. Bone marrow, skin biopsy, cartilage, and bone can be used as

autografts. (source: wikipedia ) Therefore, autogeneos grafts are harvested from same individual as

the graft is placed into. They may be arterial or venous.

Arterial autografts are widely accepted as the best bypass conduit currently available. Arterial grafts

are able to maintain their viability, demonstrate proportional growth when used in children, exhibit

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normal flexibility at joints and do not degenerate with time. Most arterial autografts are used in the

abdomen for visceral vessel reconstruction or in the groin region. Due to its size and availability,

arterial conduits are rarely used in the aortic location or for lower extremity arterial bypass. [11]

2.6.1.2 Synthetic Prosthetic Grafts

Autogenous grafts are not always available, or the graft available may be of inadequate length or

calibre. In order to provide a substitute conduit, the prosthetic vascular grafts were developed.

Grafts are constructed in a variety of ways that affect their porosity, thrombogenicity and

compliance. Prosthetic grafts are classified according to their construction method as textile or non-

textile grafts. Most currently available textiles are knitted or woven polymer grafts. Non-textile

grafts, such as poly(tetrafluoroethylene) (PTFE) and polyurethane are manufactured using techniques

of precipitation or extrusion of the polymer from solutions or sheets of the material. In order to

become an adequate conduit, the material must be biocompatible and toxic free, allergic and

carcinogenic side effects. The graft must be durable and resist to degradation and deterioration over

time. The graft should be readily available in a variety of diameters and lengths and should have an

adequate flexibility yet to be resistant to kinking and should have some degree of structural integrity

so that it can be sutured in place. [11]

2.6.1.3 Biological Prosthetic Grafts

Arterial Allografts: Arterial allografts have been used intermittently since the early days of arterial

reconstruction. Fresh or preserved allografts inserted from that time into the late 1940s, generally

were met with rapid rejection and degeneration. Allograft preservation with formalin, glycerine,

ethylene dioxide, gamma irradiation, freeze drying produced generally unsatisfactory, although

highly variable results. Failure was evidenced by mural degeneration, aneurysm formation,

haemorrhage and death. Since the biological characteristics of artery being so highly desirable, this

area remains possible that well-handled, well-preserved arterial grafts may have a role in the modern

area. [12]

Venous Allografts: Venous allografts have been studied more thoroughly than arterial allografts but

have had similar results. Satisfactory retention of viability secondary to modern techniques of

cryopreservation has recently renewed interest in cadaveric venous allografts.

Some very recent evidence suggests that an immunological privilege may be conferred by

cryopreservation. Perhaps the cryopreservation process itself leads to a general proliferative

response in the vessel providing results similar to that with synthetic graft materials in the small-

calibre application.

However, currently there is no method of graft procurement and preservation that preserve normal

endothelial and smooth muscle cell function or eliminate antigenicity available. [12]

Arterial and Venous Xenografts: Xenografts are arteries or veins of animal origin. These also have

received a great deal of interest over the years, mainly due to their availability. Xenografts have been

modified in a number of ways. They must be fixed to prevent an aggressive xenogenic immune

response that will cause early implant failure and degeneration. The most popular approach of

fixation was with some form of chemical digestion to achieve decellularization and removal of

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foreign animal proteins followed by chemical cross-linking. However, most evidence suggests that

these methods do not effectively blunt the xenogeneic immune response sufficiently to make this

approach a viable alternative. On the other hand this does not mean that there is no merit in this

strategy as an alternative for future investigations. [12]

Human Umbilical Cord Vein Allografts: Umbilical vein was introduced in the 1970s and became

popular in 1980s. The umbilical cord vein is a long, unbranched conduit, acquired from delivery

suites, mechanically or manually stripped of its surrounding tissue, fixed in glutaraldehyde and

encased in a loose mesh polyester( Dacron). The glutaraldehyde tanning procedure increases tensile

strength, masks antigenicity and sterilizes the tissue. This graft processes physical, chemical and

mechanical properties that initially seemed superior to that of existing synthetic graft materials. [12]

Despite excellent properties, aneurismal dilatation occurs due to the use of a Dacron mesh wrap. The

manufacturing is improved in 1989 but dilatation is still a long term problem. [13]

2.6.1.4 Tissue Engineered Vascular Graft

Although synthetic vascular grafts such as Dacron and ePTFE have been successfully used in treating

the pathology of large arteries( >6 mm inner diameter ), they are not suitable for replacing the

smaller diameter vessels. Tissue engineering offers the potential of providing vessel that can be used

to replace diseased and damaged native blood vessels.

The success of a tissue-based graft depends on its ability to meet several requirements. A graft must

possess a confluent, adherent and quiescent endothelium to resist thrombosis in vivo. The

mechanical behaviour of the graft must mimic the mechanical properties of a native vessel and it

must contain an elastin network to provide compliance and recoil. As it shown in Figure 2.3,

collagen and elastin are secreted by smooth muscle cells. Cross-linking stabilizes collagen and elastin,

making them less susceptible to proteolysis. Well-organized layers of insoluble collagen and elastin

result in a strong, compliant vessel.

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Figure 2.3 – Tissue engineered vascular grafts.

[ source: Shannon L., Mitchell and Laura E. Niklason, Requirements for growing tissue engineered

vascular grafts, Cardiovascular Pathology, Volume 12, Issue 2, March-April 2003 ]

Tissue engineering using either polymer or biological based scaffolds, represents the newest

approach to overcoming limitations of small diameter prosthetic vascular grafts. Their disadvantages

include thromboembolism and thrombosis, anticoagulant related haemorrhage, compliance

mismatch, neointimal hyperplasia, as well as aneurysm formation.

One of the first attempts to tissue engineering vascular grafts is to develop a biocompatible and

mechanically stable vascular graft combining human cells and a xenogenic acellular matrix as a result

of this attempt stable mechanical properties were achieved at physiological perfusion pressures in

vitro. [5]

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2.7 Desirable Properties of Vascular Prostheses

Any material that is used for vascular prostheses should include the following properties:

1. Biocompatibility and non-immunogenicity

2. Long term chemical and mechanical stability.

3. Processability.

4. Viscoelastic properties similar to blood vessels.

5. Adequate, manageable pore size and distribution.

6. Prevent graft leakage which can lead to seroma formation and blood loss, be abrasion resistant.

7. Promotes cell attachment and angiogenesis.

8. Negligible toxicity, locally, systemically and from degradation products.

9. Absorbable/Non-absorbable

10. Ability to release bioactive compounds.

11. Smooth blood flow surface

The most important criteria in designing vascular fabrics are porosity and nonthrombogenic surface.

These two criteria are closely related to each other because vascular fabrics require porous space for

tissue in-growth, which would ultimately lead to the formation of nonthrombogenic surfaces. The

generation of a nonthrombogenic surface on synthetic vascular graft materials is one of the most

important goals in the repair of damaged or diseased vascular systems.

It is desirable to have prosthesis that can be readily manufactured and that can have a diameter

which can be closely controlled to increase the efficiency of preclotting and yet allow for tissue

ingrowth and complete healing. [1]Since a certain period of time passes before blood begins to form

clots after exposure to a foreign surface, the cardiac prosthesis inside structure should be designed

to continue blood flow.

All of the existing commercial vascular grafts are constructed from a single type of non-absorbable

fibre (e.g. Dacron), their porosity is relatively constant and does not change with time. It is therefore

impossible to vary the porosity of these single-component fabrics so that they would be very tight

during implantation (i.e. low bleeding porosity) to prevent the occurrence of blood leakage and be

very porous during healing (i.e. high healing porosity) to promote fibrous tissue ingrowth for a full-

wall healing.

There are several approaches to try to design ideal vascular fabrics that would meet these two most

important criteria. The use of absorbable fibres as the sole component or as one of two components

of a fabric and the theoretical prediction of the porosity of a fabric based on some mathematical

formula appear to be promising for the design of the next generation of vascular fabrics. [9]

12

2.8 Polymeric Biomaterials

Biomedical textile materials have a large range of applications in vascular, ligament, heart valve,

heart wall and other replacement devices. The main applications for medical implant textiles are

patches for the heart/vessel system prostheses for blood vessels and for ligaments. Textile structures

for implantation are identified by structure, material composition and behaviour of the fibre surface

and degradation. [9] A major concern is the bio-compatibility of such textiles with the human body

on which they are used. [14] A biotextile in implantation must meet mechanical requirements and it

must be biocompatible.

The materials that have been used can be classified into three groups: metallic, ceramic/glass and

organic polymer.

The use of metals in prolonged direct contact with blood is of limited interest due to their stiffness;

except under very special conditions such as a frame for valve and only a few metals ( platinum,

titanium, tantalum ) are un-reactive toward 0.30 N NaCl solution at pH 7.04(blood plasma).

The ceramics, however, have such high modulus in bending that periodic deformation of any sort is

inconceivable. Like metals, they are impermeable to any component of blood (unless intentionally

created with porosity, such as sintered powder metal compacts). [4]

The third category, organic polymers, which can be matched to whatever degree that is wanted with

the gross mechanical properties of the living vessel wall such as in density, viscoelastic response by

giving appropriate design.[4] The polymeric based grafts are far more successful than metallic ones,

but they still exhibit various problems. Almost all non-absorbable polymeric based grafts share one

common disadvantage, they must be removed, in most cases, if infection develops. Therefore

absorbable vascular grafts have shown better chronic tissue tolerance because they are biodegraded

completely so that no traces of a foreign body remain. [9]

Vascular grafts are made of biomaterials such as: Teflon, Dacron-Polyester (PET), Nylon,

Polyurethane, Polytetrafluoroethylene (PTFE) and Gore-tex. [4]

The two most important polymeric biomaterials used for making vascular grafts are poly(ethylene

terepthalate ) and expanded poly(tetrafluoroethylene). The former includes Dacron and has the

largest market share. The latter is the same as Gore-tex. In addition to these two common polymeric

biomaterials as the major source for fabricating vascular grafts, the use of polyurethane based

elastomeric fibres (i.e.Spandex ) [9]

2.8.1 Teflon

Teflon is quite effective when it is woven, because it needs no previous clotting. When it is knitted

into a porous mesh, before the operation, the mesh created is highly malleable and can be

manipulated easily. It must have high resistance to in vivo degradation, low thrombogenicity and

exceptional physical and mechanical properties also it must have high elastic capacities because

13

veins are arteries are usually stretched, compressed and decompressed by muscular tissue and

ligaments in motion.

FDA has approved the use of Teflon grafts that have sizes smaller than 6 mm for arterial use and

have sizes of 6 mm or larger for veins and other vacuoles. FDA also stipulates that the design must

include an outer coating made of a biological substance such as albumin or collagen or a synthetic

coating such as silicone.

Teflon grafts are high competitors in the market and were designed to meet the need of

cardiovascular diseased patients. [4]

2.8.2 Polyester (Dacron)

The polyester prostheses have emerged as the best bio-stable synthetic material for the vascular

grafts due to its strength, endurance, long-term patency and its biostability.

Similar to Teflon, Dacron is designed to be fabricated into small strands that are woven into a mesh

and to be excellent conductors. Dacron grafts are able to be stretched and elongated by the body’s

muscular tissues and tendons. Dacron graft has a lower thrombogenicity and a higher compatibility

with the human hosts than Teflon. Dacron also resists the process of hemodialysis (seal quickly after

being punctured.)

PET or Dacron has been successfully used in large diameter grafts; however, small calibre grafts still

show an unsatisfactory high percentage of failure. Due to surface forces, blood plasma proteins

adsorb to the graft, resulting in inflammation, infection, thrombus formation and ultimately, vessel

reclosure. To prevent thrombus formation, grafts are sometimes bonded with heparin, an

anticoagulant.

The synthetic polymer polyesther terephthalate (PET), known as Dacron, is mostly recognized due its

biocompatible, resilient, flexible, durable and resistant to biodegradation and sterilization. [4]

Another benefit of Dacron polyester prostheses is that under certain porosity conditions, they can be

penetrated from the outside of the graft through pores in the graft walls by perigraft tissues thereby

fastening the prosthesis to surrounding perigraft tissues and making the prosthesis blood tight. [1]

2.8.3 Nylon

Nylon, which is different to Teflon and Dacron, is a tough material combined with an excellent

coefficient of friction and good abrasion resistance. Nylon was designed in the 1930s for fabrics and

industrial uses. Its mechanical properties make it an ideal material because it has qualities similar to

Teflon and Dacron. Nylon also has the quality of elasticity necessary for proper function and mobility

as part of the vascular system. Additional characteristics of Nylon are: processability, heat

resistance, dry/wet service capability, non-abrasive to other materials, fatigue resistance, non-

lubricated operation, noise dampening, electrical properties and chemical resistance.

Its limitations are: it does not have a level of thrombogenicity as low as Teflon and Dacron. To avoid

the clots inside of the nylon grafts have to be lined with other materials whether synthetic or

biological. [4]

14

2.8.4 Polyurethane Elastomer

Elastomers have been used as biomaterials in many cardiovascular and soft-tissue applications due

to their high elasticity, impact resistance and gas permeability. Applications of elastomer are: flexible

tubing for pacemaker leads, vascular grafts and catheters, biocompatible coatings and pumping

diaphragms for artificial hearts, grafts for reconstructive surgery and maxillofacial operations, wound

dressings, breast prostheses etc.

Elastomers are typically amorphous with low crosslink density. This gives them low to moderate

modulus and tensile properties as well as high elasticity. The majority of biomaterials used in humans

are synthetic polymers such as the polyurethane, resilient elastomers found in short and long term

cardiovascular devices, made out of natural rubbers or rubber like materials

Most common thermoplastic polyurethane, such as polyethylene and polyester, are used as

biomaterials. Thermoplastics usually exhibit moderate to high tensile strength (50 to 10.000

atmospheres of pressure) with moderate elongation (2 to 100%), and they undergo plastic

deformation at high strains.

Thermoplastics consist of linear or branched polymer chains. Depending on the structure and

molecular organization of the polymer chains, thermoplastics may be amorphous (polystyrene),

semicrystalline (low-density polyethylene), or high crystalline (high-density polyethylene) or they

may be processed into high crystalline textile fibres (polyester Dacron). Some thermoplastic

biomaterials such as poly-lactic acid and poly-glycolic acid are polymers based on a repeating amino

acid subunit.

These polypeptides are biodegradable. The rate of biodegradation in the body can be adjusted by

using copolymers. These are polymers that link two different monomer subunits into a single

polymer chain. The resultant biomaterial exhibits properties, including biodegradation, that are

intermediate between the two homo-polymers. Many polyurethane elastomers are thermoplastic in

nature. The molecular structure of a typical biomedical TPU (thermoplastic polyurethane) consists of

alternating high-melting “hard” urethane segments and liquid like “soft” segments. Hard segments

are usually the reaction product of an aromatic or aliphatic diisocyanate and a low molecular-weight,

chain-extending dialcohol or diol. In the TPUs used as biomaterials, soft segments are usually built

from (polyether or polycarbonate) polyols with terminal hydroxyl (-OH) groups.

Conventional polyether and polycarbonate TPUs generally have excellent physical properties,

combining high elongation and high tensile strength to form tough, although fairly high-modulus,

elastomers. Natural rubber latex may have an initial modulus of a few hundred pounds per square

inch (psi), an 80A aromatic polyetherurethane might have a modulus of >2000 psi, making it

considerably less compliant. On the other hand, aromatic polyether TPUs can have excellent flex life,

a tensile strength of >5000 psi (34 MPa) and ultimate elongations of >700 %. They are often used for

continuously flexing, chronic implants such as ventricular-assist devices, intra-aortic balloons and

artificial heart components.

15

The two most important diisocyanates used in biomedical TPUs are aromatic diphenylmethane

diisocynate (MDI) and its hydrogenated analog (HMDI). TPUs with hard segments made from MDI

typically have superior physical properties and chemical resisitance relative to analogous TPUs made

from HMDI, especially when compared at body temperature in an aqueous environment typical of

blood or tissue. In device applications, the use of TPUs of different hardness values within a single

family provides a considerable versatility in design and manufacturing.

Even as the first polyurethane was being used in medical devices, some investigators recognized the

possible advantages of combining silicone and polyurethane in a single biomaterial. Since 1970, the

approaches have included coatings, blends, interpenetrating networks, surface modifying additives in

urethane base polymers and most recently high-strength “structural” copolymers of silicone and

polyurethane.

In 1970, the solvent-cast silicone-polyurathane Avcothane-51 was introduced by Avco Everett

Research Laboratory (Everett, MA) as the material of construction for the first clinical intra-aortic

balloon. This combination of silicone and polyurethane was first proposed to improve the

thromboresistance of early cardiac-assist devices, which at the time were plagued by gross

thrombogenicity . Cardiothane-51 has since been reported to have excellent thromboresistance, flex

life, abrasion resistance and biostability. [4]

16

2.9 Fabric Construction

Medical textile products are based on fabrics, of which there are four types: woven, knitted, braided,

and nonwoven, shown in Figure 2.4. The first three of these are made from yarns, whereas the

fourth can be made directly from fibers, or even from polymers. Gore-Tex—based products or

electrostatically spun materials from polyurethane are examples of products made directly from

polymers. The performance of the final textile products is affected by the properties of polymers and

the structure of the fabric construction. [14]

Figure 2.4 - Constituent elements of medical textile products.

[source: Bhupender S.Gupta, Medical Textile Structures ]

Although there are many different types of polymers, only a few can be made into useful fibers. This

is because a polymer must meet certain requirements before it can be successfully and efficiently

converted into a fibrous product. Some of the most important of these requirements are:

Polymer chains should be linear, long, and flexible. Side groups should be simple, small, or polar. Polymers should be dissolvable or meltable for extrusion. Chains should be capable of being oriented and crystallized.

Common fiber-forming polymers include cellulosics (linen, cotton, rayon, acetate), proteins (wool, silk), polyamides, polyester (PET), olefins, vinyls, acrylics, polytetrafluoroethylene (PTFE), polyphenylene sulfide (PPS), aramids (Kevlar, Nomex), and polyurethanes (Lycra, Pellethane, Biomer). Each of these materials is unique in chemical structure and potential properties. For example, among the polyurethanes is an elastomeric material with high elongation and elastic recovery, whose properties nearly match those of elastin tissue fibers. This material-when extruded into fiber, fibrillar, or fabric form-derives its high elongation and elasticity from alternating patterns of crystalline hard units and noncrystalline soft units.

17

Although several of the materials mentioned above are used in traditional textile as well as medical applications, various polymeric materials—both absorbable and nonabsorbable—have been developed specifically for use in medical products. Chemical structures of some of these materials are illustrated in Figure 2.5. [14]

Figure 2.5 - Examples of fibrous materials developed for use in medicine.

[ source: Bhupender S.Gupta, Medical Textile Structures ]

The reactivity of tissues in contact with fibrous structures varies among materials and is governed by both chemical and physical characteristics. Absorbable materials typically excite greater tissue reaction, a result of the nature of the absorption process itself. Among the available materials, some are absorbed faster (e.g., polyglycolic acid, polyglactin acid) and others more slowly (e.g., polyglyconate). Semiabsorbable materials such as cotton and silk generally cause less reaction, although the tissue response may continue for an extended time. Nonabsorbable materials (e.g., nylon, polyester, polypropylene) tend to be inert and to provoke the least reaction. To minimize tissue reaction, the use of catalysts and additives is carefully controlled in medical-grade products. [14]

Through careful control of morphology, fibers can be manufactured with a range of mechanical properties. Tensile strength can vary from textile values (values needed for use in typical textile products such as apparel) of 2—6 g/den (gram/denier) up to industrial values (values typical of industrial products such as tire cords or belts) of 6—10 g/d. For high-performance applications, such as body armor or structural composites, novel spinning techniques can produce fibers with strengths approaching 30 g/den. Likewise, breaking extension can be varied over a broad range, from 10—40% for textile to 1—15% for industrial and 100—500% for elastomeric fibers. Fibers or filaments are converted into yarns by twisting or entangling processes that improve strength, abrasion resistance, and handling. Yarns are interlaced into fabrics by various mechanical processes such as weaving, knitting, braiding and non-woven. [14]

18

Figure 2.6 – Fabric techniques (a) Braided, (b) Woven, (c) Knitted [source: Frank K. Ko, Drexel University, ASM Handbook, Volume 21: Composites]

Woven, in which two sets of yarns are interlaced at right angles; knitted, in which loops of yarn are intermeshed; braided, in which three or more yarns cross one another in a diagonal pattern, shown in Figure 2.6. Knitted fabrics can be either weft or warp knit, and braided products can include tubular structures, with or without a core, as well as ribbon, are illustrated in Figure 2.7.

Figure 2.7 - Examples of woven (top left), knitted (top right, bottom left) and braided (bottom right) structures.

[source: Bhupender S.Gupta, Medical Textile Structures]

19

The properties of fabrics depend on the characteristics of the constituent yarns or fibers and on the

geometry of the formed structure. Whether a fabric is woven, knitted, braided, or nonwoven will

affect its behavior. [14] Each type of construction has positive and negative characteristics.

2.9.1 Woven

Fabrics that are woven provide strength with high dimensional stability and low permeability to

blood and it is less prone to kinking. It is used mainly in large-diameter vessels like the aorta and

major arteries from which uncontrolled bleeding could lead to fatality. The main disadvantages of

woven vascular fabrics are their very low healing porosity leading to poor healing, difficulty in

suturing, fraying of cut edges and poor compliance. [9] Therefore velour woven vascular fabrics

became an alternative to woven vascular fabrics due its high porosity and better compliance.These

new fabrics were made by floating portions of the weft or/and warp yarns so that the numbers of

intersections along both warp and weft directions were reduced considerably from the conventional

1x1 woven structure. Since the number of intersections of both weft and warp yarns are closely

related to the mechanical and physical properties of the resulting fabrics, a significant reduction in

these numbers would make the fabrics more flexible and porous. The floated portions of the yarns

have another advantage: they produce three-dimensional loose surface structures as the famework

for tissue attachment and in-growth that conventional woven fabrics do not have. [9] This stitching

technique known as a Leno weave, as can be seen in Figure 2.8, in which two warp threads twist

around a weft and can considerably diminish the fraying.[14]

Figure 2.8- Fabric and vascular grafts made using a Leno weave.

[source: Bhupender S.Gupta, Medical Textile Structure]

20

2.9.2 Knitted

The weft-knitted structures are more extensible than woven fabrics, but they are also dimensionally

unstable unless additional yarns are used to interlock the loops and reduce the extension while

increasing elastic recovery.

Warp-knitted structures are extremely versatile, and can be engineered with a variety of mechanical

properties matching those of woven fabrics. The major advantage of knitted materials is their

flexibility and inherent ability to resist unraveling when cut. A potential limitation of knitted fabrics is

their high porosity, which—unlike that of woven fabrics—cannot be reduced below a certain value

determined by the construction, as it seen in Figure2. 9 [14]

Figure 2.9 - Woven (left) and knitted (center and right) fabrics used for vascular grafts, showing

differences in porosity. [ source: Bhupender S.Gupta, Medical Textile Structures]

Due to their high porosity, they are difficult to preclot during the time of implantation and therefore

impose the problem of blood leakage. As a result, knitted vascular fabrics are not normally used in

large arteries where bleeding is a major problem. There are far more varieties of knitted than of

woven structure. Straight or branched grafts are possible by using the either weft or warp knitted

technology. However, the current commercially available knitted vascular fabrics are mainly warp

knitted with a single or double velour surface structure. [9]

All woven and knitted vascular grafts have one common appearance, a crimped structure. A

crimping process introduces “hills and valleys” along the longitudinal direction of the grafts. The

purpose of crimping these vascular grafts is to prevent kinking of the graft at its bending site. A

kinked graft would block the flow of blood at the point of bending. Since the crimped process would

destroy the smooth and even flow surface of the fabric, the wells in the valley portions could be

prone to thrombus deposits due to the stagnation of the blood flow in these regions.

The most nonconventional textile structure in vascular fabrics is Gore-tex. It consists of nodes that

are connected by fine fibrils. Because of this unique structure, it is highly porous.

21

2.9.3 Braided

Braided structures are unique in their high level of conformability, torsional stability and damage resistance. [15] Typically employed in cords and sutures, braided structures can be designed using several different patterns, either with or without a core. Because the yarns criss-cross each other, braided materials are usually porous and may imbibe fluids within the interstitial spaces between yarns or filaments. To reduce their capillarity, braided materials are often treated with a biodegradable (polylactic acid) or nonbiodegradable (Teflon) coating. Such coatings also serve to reduce chatter or noise during body movement, improve hand or feel, and help position suture knots that must be transported by pressure from a surgeon's finger from outside the body to the wound itself.

Braided material differs from woven and knitted fabrics in the method of yarn introduction into the fabric and in the manner by which the yarns are interlaced. The comparison of braided, knitted and woven fabrics is shown in Table 2.1.

Parameter Braiding Weaving Knitting

Basic direction of yarn introduction

One (machine direction)

Two (0°/90°) (warp and fill) One (0° or 90°) (warp or fill)

Basic formation technique

Intertwining (position displacement)

Interlacing (by selective insertion of 90° yarns into 0° yarn system)

Interlooping (by drawing loops of yarns over previous loops)

Table 2.1- A comparison of fabric formation technique

[source: Frank K. Ko, Drexel University, ASM Handbook, Volume 21: Composites]

2.9.4 Nonwovens

The properties of nonwoven fabrics are determined by those of the constituent polymer or fiber and by the bonding process. For instance, expanded PTFE products can be formed to meet varying porosity requirements. Because of the expanded nature of their microstructure, these materials compress easily and then expand—a suture, for example, can expand to fill the needle hole made in a tissue—allowing for tissue ingrowth in applications such as arterial and patch grafts. Polyurethane-based nonwovens produce a product that resembles collagenous material in both structure and mechanical properties, particularly compliance (extension per unit pressure or stress). The porosity of both PTFE- and polyurethane-derived nonwovens can be effectively manipulated through control of the manufacturing processes.

22

2.10 Finishing

The next process after fabric production is called finishing. The yarns may contain additives that can

result in cytotoxicity and adverse reactions when in contact with tissue. Each polymer and fabrication

process differ than others, therefore the finishing operation must be material and device specific.

Finishing includes such steps as cleaning, heat setting, bleaching, shrinking, inspection, packaging and

sterilization. Figure 2.10 shows a schematic of a typical finishing operation used in vascular graft

manufacturing. [15]

Figure 2.10- Vascular graft finishing operation

[source: Frank K. Ko, Drexel University, ASM Handbook, Volume 21: Composites]

Testing of the finished product for cytotoxocity and residual extractables is generally used in order to

make sure that all the surface additives are removed from the product’s surface prior to packing and

sterilization.

2.11 Testing and Evaluation

Once the biotextile is given to its final shape, it must be tested and evaluated to confirm that it meets

published standards and its intended end use. When developing and implementing a testing

program, a various pieces of reference information may apply, including ASTM standards, AAMI/ISO

standards, FDA documents, prior regulatory submissions, and the results of failure analyses. Table

2.2 includes the list of the suggested test methods used in the development of a textile-based

vascular graft for large vessel replacement. (ANSI /AAMI/ISO, 2001). [15]

23

Table 2.2- Sample test methods for large diameter textile graft.

[source: Frank K. Ko, Drexel University, ASM Handbook, Volume 21: Composites]

24

CHAPTER 3

MATERIALS AND PROCESSES

3.1 Material Characterization

PET and PLA filaments were chosen to produce the hybrid vascular prosthesis.

Polyester is relatively flexible and resilient. It was also preferred due to its strength, biostabilty and

long term patiency. It can be sterilized by all methods.

PLA is more elastic and bioabsorbable. It is toxicological safety. It presents good biocompatibility,

good biodegradability and excellent mechanical properties. By using PLA with PET, it is expected to

obtain a structure with better quality and an improved compliance match between the vessel and

graft with better surgical handling characteristics.

Since the surface modification of vascular grafts becomes important when they come into contact

with physiological components such as blood and living tissues, the samples were coated with

chitosan.

3.1.1 Mechanical Properties of PET and PLA

PET yarn, 68.2/2 tex, no twist multifilament and the PLA yarn, 6.3 tex monofilament were used.

In order to understand the tensile behaviour of the yarns, a universal tensile machine Hounsfield

HK100 was used for testing according to standard NP EN ISO 2062. 10 specimens for each yarn with

length of 15 cm were prepared and the test carried out. Figures 3.1 and 3.2 present the typical load-

extension curves for both yarns tested.

Analysing the curves, one can see that the materials tested present very different tensile behaviours.

The behaviour of 68.2/2 tex PET yarn may be divided into 4 different steps:

Step 1 - this first step ranges from 0 to 3,5 % extension and its characterised by a deformation under

a small load due to the elasticity of the yarn given by the texturized process; at this stage the load

applied is used to orient the filaments in its direction.

Step 2 – this second step ranges from 3,5 to 83% extension and its characterized by a linear

relationship between force and extension showing an higher slope; at this stage the filaments are

quite aligned and the load is transferred for them which oppose the load applied effectively showing

the yarn rigidity.

Step 3 – this third step ranges from 83 to 100% and starts at the elastic limit of the yarn.

Step 4 – the fourth step ranges from 100 to 120 % and corresponds to the failure of the material.

25

Figure 3.1-Force-extension curve for 68.2/2 tex PET multifilament

Similar analysis may be performed for PLA yarn:

Step 1 – from 0 to 0.2 % extension, the load applied is used to orient the filaments in its direction.

Step 2 – from 0.2 to 15% extension, its characterized by a linear relationship between force and

extension until the yield point is achieved. At this stage the material can go back to its original size.

Step 3 – from 15 to 21 % extension, the curve is almost straight at lower loading. The material is

permanently deformed.

Step 4 – from 21 to 127 % as the specimen yields, it begins to “neck”. This is where the cross

sectional area decreases because of the amount of extension. As the force increases, the yarn

eventually fails and it reaches the ultimate tensile strength.

Figure 3.2-Force-extension curve for PLA yarn

26

Table 3.1 shows the results obtained for the PLA and PET yarns tested.

Yarn Linear

Density Max

Force Elongation

at Max Force at

Break Tenacity

(Tex) (N) (%) (N) (N/Tex)

PET 34.4 12.92 19.01 7.96 0.3757 Mean

0.605 0.1371 0.3228 0.0176 S. D.

13.5 20.4 13.5 0.3924 Maximum

11.71 16.68 13.80 0.3405 Minimum

4.678 7.21 4.053 4.677 C.V.

PLA 6.3 1.578 27.66 1.498 0.2504 Mean

0.1377 0.3843 0.1036 0.0219 S.D

1.810 33.8 1.684 0.2873 Maximum

1.394 23 1.356 0.2213 Minimum

8.73 13.89 6.92 8.73 C.V.

Table 3.1-The Average of Force-Extension Values of PET and PLA

3.1.2 Winding the PET Yarns

The original PET yarn obtained from the supplier was a multifilament 34.4 tex single yarn which led

to breaking of the individual filaments during manufacturing process. The rotation of the carrier track

plate of the braiding machine was weakened the filaments and the filaments broke over time

resulting in failure of braided structure. Therefore, in order to prevent the filaments fail, the PET yarn

was wound in a doubled untwisted multifilament 68.2/2 tex. Besides the the use of this yarn was

intented to get a braided structure with more covered surface.

27

3.2 Braiding Technique

Braiding process can take place vertically or horizontally. In this study horizontally braiding machine

with 40 bobbins was used, which can be seen in Figure 3.3. The carriers were arranged in track as

required position. The mandrel, which is surrounded by a spring, was placed to the mandrel holder.

Then, the yarns were twined to a mandrel by interlacing each other and in order to form the tubular

braided structure with a high yarn tension, shown in Figure 3.4. It was experienced that when low

yarn tension was used, structure distortion was occurred.

Figure 3.3- Braiding Machine ( Trenz-Export, S.A.)

28

Figure 3.4- Braiding process

The braided fabric was heat treated at around 135°C for 10 minutes to stabilize the structure, shown

in Figure 3.5.

Figure 3.5- Oven (Memmert)

29

Two groups of samples with diameters ranging from 5 to 8mm and length around 15 mm, were

manufactured by braiding process. The first group of samples (Group 1) manufactured with 100%

PET and the second group of samples (Group 2) manufactured with 75%PET-25%PLA. The tubular

structures are shown in Figure 3.6.

Figure 3.6- Braided tubular structures

The tubular braided fabrics were analysed for their physical properties under stereoscope (Olympus)

with 200x magnification. Then they were analyzed by Image Tool. The values, such as braiding

angle, crimp density, crimp distance and crimp height were then measured in pixels and converting

to millimetre by using scale.

The wall thickness of the tubular structures was measured by using Digital Thickness Gauge (SDL

International). Thickness measurement was carried out for three times for each sample at a pressure

of 1000 Pascal.

The number of crimps in was counted and the average number of the crimps in 1 cm was calculated.

30

3.3 Coating Process

Since the surface modification of vascular grafts becomes important when they come into contact

with physiological components such as blood and living tissues, the samples were coated with

chitosan. The reasons for choosing chitosan (CS) are that CS is natural biocompatible, non-toxic and

cationic polysaccharides. When sticking to the bacterial cell wall, CS can suppress the metabolism of

the bacteria promoting the healing process and also improves the biocompatibility of the grafts. CS is

also inexpensive, readily available and safe to handle. However it is insoluable in general solvents

apart from acid aqueous solution.

3.3.1 Preperation of Chitosan Solution

3% w/v CS (medium molecular weight ~ 0.3 x 105, desacetylation of 90%) was added into a 5%

aqueous acetic acid solution in order to be dissolved, it was then filtered. The samples were placed in

a glass container of 50 ml solution and vacuumed for 10 minutes, shown in Figure 3.7. Samples 100%

PET and 75%PET-25%PLA were waited in CS solution for 24 hours.

Figure3.7- 3% w/v CS solution

31

3.3.2 Preparation of NaOH

After the samples were dried at room temperature, they were neutralized by immersion in 0.5 %

NaOH solution on methanol about 20 minutes and washed in pure methanol, seen in Figure 3.8.

Figure 3.8- 3% w/v CS Solution; 0.5 % NaOH; Methanol.

For evaluating the influence of the time, 100% PET sample with diameter 6 mm was selected. The

sample was cut into 4 pieces with 3 cm length. The first sample was waited for 5 minutes in CS

solution, second for 10 minutes, the third for 30 minutes and the fourth sample for 4 hours, as it

seen in Table 3.2.

Sample

Method

Coating Time (min)

Washing Time with 5% NaOH

(min)

Washing Time with Methanol

(min)

100%PET d:6 A 5 20 5

" B 10 60 5

" C 30 60 5

" D 240 180 5

Table 3.2- The effect of different coating and washing times on 100% PET braided structure.

After coating the samples, were dried at room temperature and washed in 0,5% NaOH on methanol,

the first sample for 20 minutes, the second and the third for 1 hour and the fourth for 3 hours and

then all samples were washed in pure methanol for 5 minutes.

Another test also carried for observation the changes after coating for 20 minutes and 24 hours and

washed for 20 minutes with NaOH. One sample with diameter 6 mm from Group 1 and one sample

32

with diameter 6 mm from Group 2 were chosen. Each sample cut into two pieces and were grouped

into two, seen in Table 3.3.

Table 3.3- The effect of coating time on 100%PET and 75%PET-25%PLA samples.

The samples were first analyzed with Scanning Electron Microscopy (SEM)(Edax) and then, in order to

observe the homogeneity of the coated surface, dyed with Eosin(Sigma) and examined under

Stereoscope(Olympus).

Sample

Method

Coating Time

Washing Time with 5% NaOH

(min)

Washing Time with Methanol

(min)

100%PET d:6 A 20 min 20 5

" B 24 hours 20 5

75%PET-25%PLA d:6 A 20 min 20 5

" B 24 hours 20 5

33

CHAPTER 4

RESULTS and DISCUSSIONS

4.1 Fabric Characterization

Uncoated and coated tubular structures’ surface, Group 1 (100% PET) and Group 2 (75% PET-25%

PLA) with varying diameters (from 5 to 8 mm), were examined under stereoscope.

4.1.1 Uncoated Tubular Structure Surface Characteristic

As it can be seen in Figure 4.1-4.2, the amount of voids between yarns increases with the increase in

diameter, this is probably due to the fibre density which decreases by increasing the diameter of the

mandrel. It also affects the flexibility of the structure. The sample with highest diameter is looser

than the other. The yarns of the uncoated 75%PET-25%PLA are inhomogeneous when compared to

100%PET surface due to the use of different materials in the same structure. The voids are dense

where the PLA and PET yarns interlaced each other. This is probably also due to the differences in the

yarn counts.

Figure 4.1 – Uncoated 100% PET samples (200x)

34

Figure 4.2- Uncoated samples 75% PET-25%PLA (x200)

4.1.2 Coated Tubular Structure Surface Characteristic

Figures 4.3 shows the images obtained for coated braided fabrics. When compare to coated 100%

PET fabric with uncoated ones, it can be seen that the amount of voids decreases and the cover

factor increases However, the surface of 75%PET-25%PLA were not covered as much as it was in

100%PET and the yarns look more disorganised, as shown in Figure 4.4. This may be occurred due to

the linear mass difference between PET and PLA resulting in irregular criss-cross surface.

35

Figure 4.3- Coated samples 100% PET (200x)

Figure 4.4- Coated samples 75% PET-25%PLA (200x)

36

4.2 Physical Properties of Tubular Structures

4.2.1 Uncoated Tubular Structures Physical Properties

One of the objective of this thesis was to examine the structural properties of the braided tubular

graft and to observe the changes according to the variation of the samples diameter. Therefore,

after manufactured and applied the heat treatment, the physical properties of the braided structure

were examined including the change in braiding angle, crimp distance, crimp height, crimp density,

wall thickness and the mass . The values were determined based on the analysis of the images

presented in Tables 4.1 and 4.2.

Diameter (mm)

5 6 7 8

Braiding Angle (°) 95.51 92.67 97.34 92.07 Average

7.77 4.14 5.36 7.21 Sd

0.08 0.044 0.05 0.078 cv(%)

Crimp Distance(mm) 2.76 3.15 3.471 2.93 Average

0.105 0.132 0.159 0.177 Sd

0.037 0.041 0.045 0.06 cv(%)

Crimp Height (mm) 0.589 0.71 0.775 0.695 Average

0.01 0.06 0.08 0.083 Sd

0.178 0.088 0.106 0.119 cv(%)

Crimp Density(crimp/cm) 3.9 3.45 3.26 3.54 Average

0.07 0.058 0.114 0.055 Sd

0.018 0.016 0.035 0.015 cv(%)

Wall Thickness (mm) 0.92 0.98 0.865 0.956 Average

0.033 0.04 0.109 0.269 Sd

0.035 0.041 0.126 0.282 cv(%)

Mandrel Radius (mm) 2 3 4 5

Spring Radius (mm) 1.5 1.5 1.5 1.5

Mass (g/10 cm) 0.225 0.234 0.236 0.225 Average

0.01 0.003 0.015 0.008 Sd

0.045 0.016 0.062 0.036 cv(%)

Table 4.1- The structural characteristic of uncoated 100% PET samples.

37

Table 4.2- Structural characteristic of uncoated 75% PET- 25% PLA samples.

4.2.1.1 Change in Braiding Angle

Analysing the data presented in Table 4.1, the diameter of the sample does not have significant

effect on braiding angle. To confirm this conclusion for each group of the samples, a plot between

braiding angle and the diameter was obtained and compared (Figure 4.5). The differences in braiding

angle between samples which have same diameter indicate that is it more likely, the braiding angle

influenced by the yarn properties and construction parameters.

Diameter (mm)

5 6

Braiding Angle (°) 83.62 91.52 average

9.16 10.73 sd

0.11 0.12 cv(%)

Crimp Distance(mm) 2.27 2.47 average

0.09 0.18 sd

0.38 0.07 cv(%)

Crimp Height (mm) 0.63 0.55 average

0.11 0.08 sd

0.18 0.14 cv(%)

Crimp Density(crimp/cm) 4.6 4.2 average

0 1.63 sd

0 0.039 cv(%)

Wall Thickness (mm) 0.75 0.74 average

0.042 0.066 sd

0.056 0.089 cv(%)

Mandrel Radius (mm) 2 3

Spring Radius (mm) 1,5 1,5

Mass (g/10 cm) 0.23 0.24 average

0.012 0.007 sd

0.054 0.028 cv(%)

38

Figure 4.5- Variation of the braiding angles according to the diameter

4.2.1.2 Change in Crimp Distance

The Figure 4.6 shows that with an increase in diameter from 5 mm to 7 mm, the crimp distance of

the fabric increased. The loose structure of the sample with diameter 8 mm, which can be discussed

before, affects the crimp distance therefore unlike other samples the crimp distance decreases.

Figure 4.6 – A comparison of crimp distance.

39

4.2.1.3 Change in Crimp Height

As it can be seen in Figure 4.7, the similar curve was also obtained for 100% PET samples as it was in

crimp distance-diameter relationship graph. The loose structure of sample with diameter 8 mm also

shows significant effect on crimp height and reduces the value. The curve for the 75%PET-25%PLA

samples drops by decreasing the diameter. As it can be is seen in the graph, samples with diameter 6

mm show almost the same value.

Figure 4.7 - A comparison of crimp height.

4.2.1.4 Change in Wall Thickness

Thickness of the tubular structures is an important factor that effects their mechanical properties.

Thickness measurement was made at pressure 1000 Pascal. Figure 4.8 represents the relationship

between wall thickness and diameter. Comparing the thickness values of the samples, it can be seen

that there is no significant difference between the thickness values at different diameters.

40

Figure 4.8- A comparison of wall thickness.

4.2.1.5 Change in Mass

Values of the mass of the fabrics were determined and are presented in Figure 4.9. As it can be seen

in the graph, with an increase in diameter from 5 mm to 7 mm, the mass of the fabric increases. This

result also confirms the observation of the loose structure of the sample with greatest diameter.

Figure 4.9- A comparison of mass according to diameter variation

41

4.2.2 Coated Tubular Structures Physical Properties

After coating the samples, the physical properties such as change in braiding angle, change in crimp

distance, change in crimp height, change in crimp density, change in wall thickness and the mass

were examined. The values were determined and are presented in Tables 4.3 and 4.4 for 100% PET

and 75/25 PET/PLA respectively.

Diameter (mm)

5 6 7 8

Braiding Angle (°) 86.758 94.26 90.232 80.18 Average

6.32 5.805 3.666 7.389 Sd

0.073 0.062 0.0406 0.092 cv(%)

Crimp Distance(mm) 2.75 2.933 3.338 3.053 Average

0.044 0.042 0.0932 0.19 Sd

0.016 0.014 0.0279 0.0622 cv(%)

Crimp Height (mm) 0.536 0.55 0.613 0.67 Average

0.0409 0.101 0.035 0.111 Sd

0.076 0.184 0.057 0.1658 cv(%)

Crimp Density (crimp/cm) 4.5 4 4 4 Average

Wall Thickness (mm) 1 1.04 1.10 1.18 Mandrel Radius (mm) 2 3 4 5

Spring Radius (mm) 1.5 1.5 1.5 1.5

Mass (g/10 cm) 0.244 0.284 0.271 0.27

Table 4.3- Structural characteristic of coated 100% PET samples.

Table 4.4- Structural characteristic of coated 75%PET-25%PLA samples.

Diameter (mm)

5 6

Braiding Angle (°) 91.88 93.73 Average

7.618 11.038 Sd

0.083 0.118 cv(%)

Crimp Distance(mm) 2.677 2.284 Average

0.095 0.137 Sd

0.0353 0.0598 cv(%)

Crimp Height (mm) 0.664 0.639 Average

0.119 0.159 Sd

0.18 0.249 cv(%)

Crimp Density(crimp/cm) 6 5 Average

Wall Thickness (mm) 1.60 1.40 Mandrel Radius (mm) 2 3

Spring Radius (mm) 1.5 1.5

Mass (g/10 cm) 0.293 0.297

42

It was observed that the coating process had no significant effects on crimp height, crimp distance, crimp density. As expected, the noticeable difference obtained in the values of wall thickness and mass. As can be seen in Figure 4.10, the 75/25 PET/PLA fabrics coated with CS present higher value when compared to 100% PET ones. This may be due to PLA interacts with CS more than PET does.

Figure 4.10- Effect of Coating on Wall Thickness.

43

4.3 Comparison of Coated and Uncoated Tubular Structures

Analysing Figures 4.11 and 4.12, one can be seen that the CS treatment leads to a braided structure

with more irregular surface. This fact is more visible when PLA is used, this probably due to the PLA

interacts with CS more than PET.

Figure 4.11- Coated and Uncoated 100% PET d: 6. (200x)

Figure 4.12- Coated and Uncoated 75% PET-25%PLA d: 5. (200x)

44

4.4 Evaluating the Influence of Coating and Washing Time

Four samples were coated and washed at different waiting period in order to see the influence of the

time on the quantity of chitosan that was adsorbed.These methods were described before in

Chapter 2 under the title “Coating Process.” From the Figure 4.13 it is seen that the maximum

amount of chitosan was adsorbed after coating the samples for 10 minutes and washed with NaoH

for 1 hour.

Figure 4.13- The effect of coating and washing time with 0,5 % NaOH on the quantity of chitosan

coated. ( 100% PET d:6 )

In order to determine the changes in fabrics after treated with different coating times but the same

washing time, two braided structure with diameter 6 mm were chosen from each group of samples

and cut into two pieces. First set of samples waited in CS solution for 20 minutes and the other set

for 24 hours. Then they were weighted by using sensitive scale (Mettler PM 300). In order to

determine the change in length the number of crimps in 2 cm was counted before and after the

coating process, as it can be seen in Table 4.5.

Uncoated Coated

Coating Time

Sample

Weight (g)

Crimps (crimp/2cm)

Weight (g)

Crimps (crimp/2cm)

20 min 100% PET 0.26 7.5 0.3 7

24 hour " 0.24 7.5 0.26 7

20 min 75/25

PET/PLA 0.14 8.5 0.15 8

24 hour " 0.15 8.5 0.18 8.5

Table 4.5- Change in mass and length after coating.

45

Analysing the results shown Figure 4.13, for 100% PET samples the maximum amount of CS was adsorbed when the fabric was impregnated with CS solution for 20 minutes, however the result for 75/25 PET/PLA was not the same, the maximum CS amount was obtained after waiting in CS solution for 24 hour.

However, at this stage it still was not known what was the quality of coated surface and how

homogene it was. Therefore two analayses were done in order to evalute the coating quality which

will be discussed in following section.

4.5 Evaluating Chitosan Adhesion to Braided Fabric Surface

Both PLA and PET are known for their strong hydrophobic character which directly affects the quality

of coating process, therefore it was necessary to analyse the surface into more detailed.

4.5.1 SEM Analyses

The SEM has a large depth of field , which allows a large amount of the sample to be in focus at one

time and examined at a high magnification. SEM analyses requires the sample to be conductive

therefore the samples examined were covered with Au-Pd powder and then mounted on a specimen

and placed on the stage of the chamber in order to be analysed. The micrographs obtained from SEM

are presented in Figures 4.14 and 4.15.

46

Figure 4.14 Micrographs of Coated 100% PET Braided Structure

47

Figure 4.15- Micrographs of Coated 75/25 PET/PLA Braided Structure

SEM images of braided surface shows that the CS used has good film forming ability which can be

seen clearly in Images 2, 5, 11, 14. However it is also clearly seen that there was a compatibility

problem between the fibre surface and the CS. The CS looks to be bonded better in Image 12 and 15

which represent the surface of PLA monofilament, therefore it can be said that PLA reacted with CS

more than PET.

The effect of CS which flexed the fibres can also be seen in Image 1, 4, 10, 13 when compare to

Image 7 and 16.

48

4.5.2 A Direct-Staining Method

In order to evaluate CS adhesion to braided structure, the coated structures were visualised after

staining with 0.1% w/v of Eosin(Sigma) solution.

CS treated and untreated fabrics were stained with 5 ml of 0.10 w/v solution for 10 min at 32 °C and

washed on three occasions in 0.25 M sucrose solution in order to remove uncomplexed dye. The

stained fabrics then examined by stereoscope which can be seen in Figure 4.16.

Figure 4.16- Stereoscope images of coated and uncoated structure were taken after treatment with Eosin. (200x)

From the Images 2 and 5, it is seen that the 24 hour coating time affected the intensity of CS since

the color intensity is higher than for Images 1 and 4, however the homogenity of the color was not

the same in these images. 100%PET fabric shows better homogenity after 24 hour coating period

than 75/25 PET/PLA fabric.

As expected, no evidence of CS is seen in Images 3 and 6, however it can be seen in Image 6 that PLA

was stained with Eosin which react with CS, more than PET. This also supports that the reaction

between PLA and CS is more effective than PET and CS which is also showed in previous sections with

other analyses.

49

CHAPTER 5

CONCLUSION AND FUTURE RECOMMENDATIONS

5.1 Summary

PET which is more commonly known as Dacron, is a widely used polymer in large vessels

reconstruction even though the surface of polyester is prone to protein adhesion so, in the long term

uses, it leads to compatibility problems. The surface structures in vascular grafts play a key role in

patency rate, therefore new biodegradable materials based on hydrophobic polyester and

hydrophilic polysaccharide have taken much attention in vivo medical applications because they

complete each other to improve their surface quality.

The most commonly used Dacron grafts are produced either by weaving or knitting technology.

Although there have been researches about other manufacturing techniques such as braiding or non-

woven; they have not reached the success.

The goal of this study is to design and produce braided vascular grafts with diameter ranging from 5

mm to 8 mm using PET and PLA yarns and to study their physical properties. Moreover, the influence

of coating with CS in order to promote the healing process and improve the biocompatibility of the

grafts, is also studied. First, the materials selected were characterised. Second, the process

parameters were determined; the braided structures were manufactured using a horizontal braiding

machine according to a patented braiding process. Two groups of samples were produced: in the first

group 100% PET was used and in the second group 75%PET-25%PLA were used and the samples were

coated with CS. Third, the physical properties of such tubular fabrics were determined before and

after coating process and the changes in their properties were noted. Fourth, the effect of coating

process and the adhesion of CS on the surface were determined.

The physical analyses showed that the physical properties changed when the diameter of the

structure is increased. The yarns interlaced each other more tightly with diameter 5 mm and 6mm

compare to 7 mm. Relatively loose structure was obtained with diameter 8 mm. The surface of the

samples in Group 1 showed uniform surface compare to Group 2, yarns interlaced each other

uniformly and covered the surface more effectively than in Group 2. The noticeable change was seen

in the mass of the fabrics with diameter 8 mm which was the result of its loose structure.

The results obtained from the evaluation of coating time showed that the coating time affected the

quantity of CS that was adsorbed and the homogeneity of the adhesion. The quantity of CS was the

highest after 20 minutes treatment with CS, however the homogeneity was much better after 24

hours treatment. The results also showed that there was a compatibility problem between the fibre

surface and CS, however comparing Group 1 with Group 2, it was seen that CS bonded better with

PLA. CS is a very hydrophilic material and spreads well on a hydrophilic surface. However, PLA and

PET are hydrophobic materials, so the CS did not spread well on them and it appears to be

delaminated. The use of non-ionic wetting agent may help the CS to spread on PET. In order to

prevent the delamination a binding agent may be necessary to bond the CS to PET since there are no

functional groups for CS to interact with.

50

5.2 Limitation of Approach

The first approach is limited by the needs for materials (PET and PLA) with the same linear mass in

order to obtain a regular surface that the yarn interlaced each other more uniformly. The second

approach requires stable process parameters to be sure that all the structures are produced with the

same yarn tension. The third approach requires a surface treatment before coating CS in order to

improve the quality of coating. The fourth approach requires a mechanical test in order to

determine properties of tubes such as burst strength, water permeability, porosity. The fifth

approach requires more experimentation in the wet finishing of a hydrophobic fibre.

5.3 Future Recommendations

Although this study has given some idea on the design of CS coated braided tubular structure using

PET and PLA yarn, further investigation is necessary in the following areas:

Investigate the materials behaviour during the manufacturing process, understand their

coherence to each other.

Evaluate the effect of manufacturing process on the structure’s physical and mechanical

properties.

Apply the burst strength, tensile strength, the water permeability and porostiy test in order

to evaluate the mechanical properties.

The materials that were used in this work are PET due to its strength, biostabilty and long

term patiency and PLA due to its good biocompatibility, good biodegradability and excellent

mechanical properties. CS which is a hydrophilic polysaccharide was selected to coat the

tubular structure due its non-toxicity, biocompatibility properties. However, PET and PLA are

strong hydrophobic material which limited CS to be spread on them resulting in delaminated.

Therefore it is proposed to use non ionic wetting agent to help the CS to spread on PET.

Also, in order to prevent the delamination a binding agent may be necessary to bond the CS

to PET surface since there are no functional groups for CS to interact with.

51

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