introduction & review of literature -...

Amrita Centre for Nanosciences and Molecular Medicine

Chapter 1

Introduction & Review of Literature


1.1 Introduction

Coronary heart disease (CHD) is one of the leading causes of death worldwide

1,2,3. Atherosclerosis is the commonest form of CHD characterized by arterial

narrowing (stenosis) due to the accumulation of fatty deposits beneath the

endothelium (i.e. the layer of cells forming the inner lining of a blood vessel), mainly

affecting the coronary arteries owing to its smaller diameter. This buildup can cause

restriction of blood flow, thus depriving heart muscles and other tissues of oxygen 4-6

.

Under such cases of severe atherosclerosis, endoprosthetic devices such as

cardiovascular stents are used for mechanically keeping the vessel open, thereby

restoring the blood flow 7-11

. Stents are elongated tubular metallic structures, with

either solid or lattice-like walls, deployed inside the artery by an angioplastic

procedure. Because of the need for the appropriate mechanical property compatible

with stent deployment to keep the vessels open, metals (including Ti, stainless steel,

Nitinol, and CoCr alloys) have been widely used as coronary stents 12

. A successful

vascular stent should support natural monolayer coverage of vascular endothelial cells

over the stent surface, so as to integrate it to the vascular wall 13-15

. The reappearance

of plaques at the stented site or in-stent restenosis (ISR) is a known problem

associated with using bare metal stents, eventually leading to thrombosis 16-19

. This

occurs mainly due to excessive vascular smooth muscle cell (VSMC) proliferation

and dysfunction of endothelial cells (ECs) as a result of the injury and inflammation

at the time of stent implantation. Numerous strategies have been adopted to overcome

this limitation pertaining to BMS, of which stent coatings using polymer, inorganic

materials, and drugs 12,20-23

are widely explored. Amongst them, the most promising

solution till date to address ISR is the use of drug eluting stents (DES) which relies on

the release of anti cell-proliferative, immunosuppressive or anti-thrombotic drugs

from a polymeric coating on the stent surface that inhibits the proliferation of smooth

muscle cells as well as reduces thrombus formation within the lumen 24

. However,

this strategy of drug induced inhibition of hyperplasia can also interfere with the re-

establishment of a healthy endothelium, leaving an unendothelialized bare metallic

surface or a dysfunctional endothelium. The polymeric degradation products from


stent coatings can also further aggravate this situation. These can potentially enhance

the chance of thrombosis after the cessation of the drugs, forcing the patients to thrive

on treatments such as anti-platelet therapies for their lifetime 25-27

. Thus the problem

of thrombosis still remains unresolved. In this context, an approach, which is

essentially non-destructive to cells (unlike the case of DES), by specifically

promoting endothelialization process is desirable 28

. Establishment of a healthy

endothelial layer on stents could act as a natural anti-thrombotic barrier, preventing

thrombus formation and subsequently inhibiting smooth muscle proliferation and

platelet aggregation 29-31

. Immobilization of biological agents to promote

endothelialization and hemocompatibility such as fibrin, heparin sulfates, anti platelet

agents, antibodies, etc exhibited promising in-vivo results. However, the technical

complexities, processing cost, and their stability (activity) inside animal body are the

major associated challenges 12, 20-22

.

Moreover, permanent body implants such as stents which are in direct contact

with blood, requires a surface that is hemocompatible throughout its lifespan 21

. The

degree of thrombus formation subsequent to stent deployment is influenced primarily

by the choice of the stent material, site of implantation and blood flow. Despite

extensive anticoagulation approaches in stent design or adjuvant therapy during

angioplastic procedure, instances of stent thrombosis occurred as late angiographic

outcomes 32,33

. Hence, it is important that these device surfaces have minimal

adsorption of thrombogenic blood proteins and reduced interaction with coagulation

factors. Literature reports various strategies such as the use of heparin, antiplatelet

agents, inorganic coatings, polymers, etc for modifying biomedical prosthesis to

improve its hemocompatibility 34

. However, their extensive clinical uses on stents are

limited by the short plasma half life of antiplatelet agents/ anticoagulants and high

cost 34,35

. Herein, inorganic coatings are suggested as stent surface coatings to

improve their electro-mechanical properties and serve as a passive layer preventing its

exposure to blood 22,24

. However, ceramic coatings such as SiC, TiO2, diamond like

carbon (DLC) etc., on stent struts are brittle in nature, making it necessary to be

applied as a thin coating to make it crack-free and stable 36-54

. Conventional


techniques such as dip coating, electrodeposition, etc make it extremely difficult to

achieve a reduced coating thickness, which remains their major challenge 55,56

.In this

context, nanosurface modification has been highlighted as a promising approach to

improve the cellular behavior and tissue integration of biomedical implants 57-63

.Such

surface modifications can be achieved by various methods which include

electrochemical processing, hydrothermal treatment, sandblasting/polishing etc.,

resulting in the formation of diverse ordered/disordered nanostructures on metallic

substrates 64-66

. The creation of nanoscale features on vascular metallic stent surfaces

can mimic the natural structure of the healthy vessel wall, making it more

cytocompatible 67

. Nanostructures could actively interact with the surrounding

biological environment and deliver specific signals to guide and control cell activities,

without the addition of exogenous chemical agents (i.e. growth factors, drugs) 64,65

.

Several studies have demonstrated the utility of nanostructured metallic titanium

surfaces in improving osteoblast response, osseointegration, and also in accelerating

endothelialization 65,67-69

. Titanium despite being well accepted for biomedical

implant applications is afflicted with inadequate mechanical properties, and hence is

not preferred for making stents 12,21

. Nevertheless, titanium in its oxide form, viz.,

titania (TiO2), owing to the excellent tissue response it renders, is proposed as a

suitable stent coating on bare metal stents, providing a hemocompatible surface; with

alterations in its surface chemistry and topography contributing to improved

biocompatibility 34,35,70-73

. The fact that nanoscale surface topography stimulates and

controls several molecular and cellular events at the tissue/implant interface has

prompted investigations of such topographies in the design of implantable metals 74,75

.

1.1.1 Scope of the Thesis

Stent surfaces that promote re-endothelialization would be a good option for

preventing stent thrombosis after long term implantation. Nanostructuring on vascular

metallic stent surfaces can mimic the natural structure of the healthy vessel wall,

making it more favorable for endothelial cells, promoting ECM secretion and

hemocompatibility, while subsequently minimizing VSMC growth, which is the

strategy adopted for this thesis work. This thesis work describes the generation of an


array of different TiO2 nanostructures of distinct morphologies viz., Nanoleaves,

Nanopores, and Nanorods on metallic surfaces via a simple, scalable and cost-

effective technique and comparing their vascular cell response in-vitro. This strategy

of creating specific nanoscale topographies which are capable of naturally modulating

vascular cell behavior can be a cost-effective option for the regulation of the

fundamental factors responsible for in-stent restenosis (such as hyper smooth muscle

proliferation, endothelial dysfunction, platelet aggregation etc). Such biocompatible

metals would present a superior treatment modality without the use of costly drug

eluting stents that demand continual use of anti-platelet therapy for prolonged

durations and have long term toxicity effects. Hence such nanostructures are of great

translational value as a vital ‗polymer-less and drug-free‘ approach for surface

modifying clinically used bare metal stents based on stainless steel, and can be

proclaimed as a viable alternative to the drug eluting strategies.

1.2 Review of literature

1.2.1Blood vessel structure and function

Oxygenated blood is distributed to the heart via the coronary arteries which

are on the surface of the heart. The two primary coronary arteries are the left and right

main coronary arteries. These originate near the cusps of the aortic valve. Additional

arteries branch off the two main coronary arteries to supply the heart muscle with

blood. Since coronary arteries deliver blood to the heart muscle, any coronary artery

disorder or disease can have serious implications by reducing the flow of oxygen and

nutrients to the heart, which may lead to myocardial infarction and possibly death.

The vascular system is composed of arteries, capillaries and veins. The wall of an

artery consists of three layers which are represented in Fig. 1.3. The innermost layer,

which is in direct contact with the flow of blood, is the tunica intima. This layer is

lined with endothelial cells (ECs) which form flat pavement-like patterns on the

inside of the vessels and are surrounded by a connective tissue basement membrane.

The basement membrane separates endothelium from the underlying layers. ECs act

as a protective barrier and control the exchange of nutrients and fluid between blood


and tissue, which is the basis for the maintenance of stability of the physiological

environment, essential to cell survival. The middle layer is the tunica media and it is

usually the thickest layer. It is composed of smooth muscle cells (SMC) and

extracellular matrix (ECM). It not only provides support for the vessel but also

changes the vessel diameter to regulate blood flow and blood pressure. The outermost

layer, which attaches the vessel to the surrounding tissue, is the tunica adventitia. It is

composed of connective tissue and contains primarily fibroblasts, elastic and

collagenous fibers. These fibers allow the vessel to stretch whilst preventing

overexpansion due to the pressure that is exerted on the walls by blood flow. Blood

passes through capillaries into venules, it then flows into progressively larger veins

until it reaches the heart. The walls of the veins have the same three layers as the

arteries, but develop thinner walls. The tunica media (smooth muscle) and tunica

externa (connective tissue) are much reduced and have the media composed primarily

of SMC, with relatively low amounts of elastic tissue 76

. Fig. 1.3 shows a clear picture

on comparison of the inner structure of arteries and veins.

Fig. 1.3 Schematic representation of the three layers of vascular vessel wall comparing that of

arteries and veins. Courtesy: http://apocketmerlin.tumblr.com


1.2.2 Coronary heart disease

Cardiovascular disease broadly covers a range of conditions affecting both the

heart and the blood vessels. It is the leading cause of death in the aged western

population 77

. In America alone, almost 2,400 deaths occur each day as a result of the

disease, an average rate of one death every 37 seconds 78

. When atherosclerosis,

which is described as the hardening and loss of elasticity of arteries, occurs in the

coronary arteries it is referred to as coronary artery disease (CAD). This causes about

2100 deaths annually per million of the population in England and Wales (about

110,000 deaths in total). The disease is typically caused by the deposition of

atherosclerotic plaques (inflamed fatty deposits) on the inner wall of arteries 7, which

narrows the vessel lumen and obstructs the coronary arteries. This narrowing is

commonly referred to as a stenosis and leads to a consequential restriction of the

blood supply to the muscle cells surrounding the heart.

Coronary artery stenosis may be asymptomatic or may lead to shortness of

breath and angina, a chest pain. A critical reduction of the blood supply to the heart

may result in myocardial infarction or death. The atherosclerotic lesions that form

may also become fragile and rupture, resulting in thrombus formation and a further

restriction of blood flow. Fig. 1.1 depicts how atherosclerotic lesions gradually lead

to thrombosis.

Fig. 1.1 Diagram representing the atherosclerotic condition leading to thrombosis.

Image courtesy: http://www.thenewstribe.com


The symptoms and treatment depend on which set of arteries are affected.

Peripheral artery disease mainly relates to vessels of the lower limbs 80

. However, the

symptomatic prevalence is mainly in vessels smaller than 6 mm internal diameter.

1.2.3 Stents and Coronary Disease

Stents are expandable meshed tubes used either to reinforce body vessels

possessing weak walls or to increase the internal diameter of an arterial vessel to

allow an improved flow of fluids such as blood. The use of arterial stents in particular

has grown significantly over the last 20 years due to an ageing population and the

changes in diet which has led to an increase in cardiovascular illness 81

. Estimates

vary, but it is predicted that coronary stents had a market value of $7.2 billion in 2012

and will continue to grow at a rate of 6% per annum thereafter. In 2009, over a

million US citizens received angioplasty/stent interventions. In the treatment of

coronary artery disease, stents offer a less invasive alternative to the Coronary Artery

Bypass Graft. It is estimated that for every bypass operation, there are four instances

of stents being employed as an alternative approach 82

.

Stents are fabricated by laser cutting shaped sections from a metal tube 83

.

For stenting, an opening is made in the patient (groin, arm or neck) and a catheter is

used to guide a deflated balloon inside a stent to the correct position in the artery 84-86

.

X-rays and dye flow are used to identify the area of the artery suffering from plaque

build-up and for associated stent positioning. Once positioned, the balloon is inflated,

causing the stent to expand and therefore the plaque to be pushed back against the

inner walls87,88

. Upon deflation and withdrawal of the balloon, the stent remains in

place. Fig. 1.2 depicts the steps involved in the intervention procedure of stent

deployment. In some instances, balloon inflation inside the artery is performed

without a stent (to assist initial widening) and then a stent is subsequently positioned

89,90.


Fig. 1.2: Coronary stenting procedure (A) a catheter is fed into the femoral artery of the upper

leg. (B) The catheter is fed up to coronary arteries to an area of blockage. (C) A dye is released,

allowing visualization of the blockage. (D) A stent is placed on the balloon-tipped catheter. The

balloon is inflated, opening the artery. (E) The stent holds the artery open after the catheter is

removed. Courtesy: http://www.surgeryencyclopedia.com

1.2.4 History of coronary stenting

At the beginning of the 20th century, glass tubes that became a prototype of

stents were implanted into blood vessels of animals. In the next stage, with the origin

of percutaneous transluminal angioplasty (PTA), the expansion of blood vessels was

achieved by increasing the diameter of a catheter tube91

. This trial failed because of

migration and the development of thrombus. In 1985, Palmaz and his colleagues

developed the first balloon expanded stent 92

and just 1 year later, Gianturco

developed a balloon-expanded coil stent 93

made of stainless steel. The first self-

expandable stent, Wallstent, made of a cobalt (Co)–chromium (Cr) alloy (Elgiloy),

was clinically applied in 1986 94

. Another popular self-expandable stent, SMART,

consists of a superelastic nickel (Ni)– titanium (Ti) alloy 95

. Since the 1990s, stents

have been used in coronary arteries. Typical popular stents for this purpose are the


Palmaz–Schatz stent and the Gianturco–Roubin stent. New designs and functions of

stents have also been developed. Today, for example, the majority of patients

undergoing percutaneous transluminal coronary angioplasty (PTCA) receive a stent.

However, restenosis follows PTCA in 30–40% of coronary lesions within 6 months

96,97. Although providing intra-arterial support with bare metal stents (BMS)

dramatically improves the angiographic and clinical outcome of patients to a

restenosis rate of 20–30% 96,97

, in-stent restenosis still remains a major limitation for

this approach with exaggerated intimal hyperplasia 98

. The advent of DES, which

release drugs such as sirolimus and paclitaxel for localized delivery, is a major

advancement in the evolution of stents. However, there is a risk of late stent

thrombosis (LST) associated with DES 99,100

.

1.2.5 Bare metallic stents (BMS)

Balloon expandable stents should have the ability to undergo plastic

deformation and then maintain the required size once deployed 101

. Self-expanding

stents, on the other hand, should have sufficient elasticity to be compressed for

delivery and then expanded in the target area101

.

The characteristics of an ideal stent have been described in numerous reviews 102-104

.

In general, it should have

(1) low profile—ability to be crimped on the balloon catheter supported by a guide

wire;

(2) good expandability ratio—once the stent is inserted at the target area and the

balloon is inflated, the stent should undergo sufficient expansion and conform to the

vessel wall;

(3) sufficient radial hoop strength and negligible recoil—once implanted, the stent

should be able to overcome the forces imposed by the atherosclerotic arterial wall and

should not collapse;

(4) sufficient flexibility—it should be flexible enough to travel through even the

smaller diameter atherosclerotic arteries;

(5) adequate radiopacity/magnetic resonance imaging (MRI) compatibility—to assist

clinicians in assessing the in-vivo location of the stent;


(6) thromboresistivity—the material should be blood compatible and not encourages

platelet adhesion and deposition

(7) drug delivery capacity—this has become one of the indispensable requirements

for stents of the modern era to prevent restenosis.

1.2.6 Materials for BMS

Generally, the metals commonly used for manufacturing stents are 316L

stainless steel (316L SS), platinum–iridium (Pt–Ir) alloy, tantalum (Ta), nitinol (Ni–

Ti), cobalt– chromium (Co–Cr) alloy, titanium (Ti), pure iron (Fe), and magnesium

(Mg) alloys owing to their adequate mechanical properties which are compared in

Table1.1. A list of stent materials that are recommended for coronary stenting

applications based on their characteristics, with a rationale for its choice has been

depicted in Table 1.2.

i) Stainless steel (SS)

Stents fabricated from stainless steel, which is an iron (Fe)-based alloy,

contain over 18 % Cr, 8 % Ni73

. Stainless steel does not corrode in an oxygen-

containing atmosphere, but they corrode locally and sometimes form pits in chloride

solutions as found in some body fluids. Elements such as Ni, molybdenum (Mo),

copper (Cu), Ti, niobium (Nb), and nitrogen (N) are added to stainless steel to

improve its corrosion resistance, heat resistance, strength, and workability. The

metallurgical structure, strength, and corrosion resistance of stainless steel depend on

the concentrations of Ni and Cr, and stainless steels are categorized as ferritic (Fe–Cr

system), martensitic (Fe–Cr system), or austenitic (Fe–Cr–Ni system) types according

to the crystal phase. Amongst the wide variety of stainless steel available, 316L SS is

the commonly used one for biomedical applications, possessing low carbon content

(below 0.03%). Type 316L stainless steel has Fe, <0.03% C, 16-18.5% Cr, 10-14%

Ni, 2-3% Mo, <2% Mn, <1% Si, <0.045% P, <0.03% S. 2- 3% Molybdenum added

enhance the corrosion resistance steel as the stability of the passive film increases.

The main reason for the use of stainless steel for coronary applications is the good

balance of strength and elongation, which facilitates the manufacture of the stent, the


plasticity for balloon expansion, and the maintenance of morphology to resist the

elastic recoil of blood vessels.

ii) Tantalum

Ta has excellent corrosion resistance because of its highly stable surface oxide

layer, which prevents electron exchange between the metal and the adsorbed

biological species 105,106

. It has been coated on a 316L SS surfaces to improve

corrosion properties, thereby enhancing the biocompatibility of 316L SS107

. It has

excellent fluoroscopic visibility because of its high density. It is an MRI compatible

material as it produces no significant artifacts because of its non-ferromagnetic

properties108,109

. Ta is also known for its good biocompatibility 110,111

. Enhanced

hemocompatibility was achieved by adding Ta to Ti oxide and the films showed

improved endothelialization rate as the percentage concentration of Ta increased 112-

113. Though the biocompatibility and visibility properties of Ta are superior to 316L

SS, the commercial availability of Ta stents is lower than 316L SS stents. This is

mainly because of its poor mechanical properties. Since the yield strength of Ta is

closer to its tensile strength, these stents have a higher possibility of breaking during

deployment.

iii) Titanium and alloys

Ti and its alloys have been extensively used in orthopedic and dental

applications because of their excellent biocompatibility114,115

. Its highly stable surface

oxide layer provides excellent corrosion resistance115, 116

.However, Ti is not

commonly used for making stents. Although Ti and Co–Cr both have high yield

strength in approximately the same range, Ti has a significantly lower tensile

strength. Thus, there is a higher probability of tensile failure of the Ti stents when

expanded to stresses beyond their yield strengths, which is the norm in balloon

expandable stent deployment. Alloying Ti with materials that reduce its yield strength

while retaining tensile properties might prove to be optimum. Because of its low

ductility, Ti stents are more prone to fracture. Because of these inadequate

mechanical properties, commercially pure Ti failed to make an impact as the sole


stent material. However, the applications of Ti are not limited to coronary stent

applications. Ti-nitride oxide coating on 316L SS was found to be biologically inert

with reduced platelet and fibrinogen deposition, thereby reducing neointimal

hyperplasia117

. The Titan stent (Hexacath, France), which has implemented this

coating technique, has shown promising results in human clinical trials 118,119

. Also,

Ti-based Ta and niobium alloys, which have potential applications for stents, showed

excellent hemocompatibility 120

. One of the Ti alloys which is extensively used for

making stents is Ni–Ti.

iv) Co–Cr alloy

Co–Cr alloys, which conform to ASTM standards F562 and F90, have been

used in dental and orthopedic applications for decades 121

and recently have been used

for making stents. These alloys have excellent radial strength because of their high

elastic modulus (Table 1.1).

Table 1.1 Mechanical properties of the metals that are used for making stents; G. Mani et al.

Biomaterials 28 (2007)

The thickness of the struts is a critical issue in designing a stent 122,123

, hence,

the ability to make ultra-thin struts with increased strength using these alloys is one of

their main attractions124

. In addition to this, they are radio-opaque 125

and MRI-

compatible 126

. The cobalt alloy platform DRIVER stents (Medtronic Inc, USA) are

commercially available in Europe. Recently, FDA approved the L-605Co–Cr alloy

Guidant Multi-Link Vision stent for clinical use127

. Co–Cr alloys show high strength,

toughness, castability, corrosion resistance, and wear resistance. The corrosion


resistance of Co–Cr alloys is better than that of stainless steel. The surface oxide film

of a Co–Cr–Mo alloy contains oxides of cobalt and chromium which impart corrosion

resistance to the alloy. Table 1.2 provides a list of materials that are recommended for

coronary stenting applications based on their characteristics, with a rationale for its

choice.

Table 1.2 Materials with ideal characteristics for coronary stent applications; G. Mani et al.

Biomaterials 28 (2007)

1.2.7 In-stent restenosis - mechanism, pathogenesis

Restenosis is the re-narrowing of a blood vessel causing a reduction of the

luminal size, consequently restricting blood flow after intravascular procedure.

Restenosis is mainly characterized by intimal hyperplasia and vessel remodeling128

.

The hyperplasia is an abnormal or unusual increase in the cells composing the intima.

Restenosis is a combined result of a biological response and mechanical reaction to

percutaneous coronary intervention (PCI). At the early phase of restenosis, elastic

recoil takes place due to the mechanical response of the elastic fibers of vascular wall

to overstretching by balloon catheter. The recoil occurs within minutes following

balloon deflation; the recoil may cause up to a 40% lumen loss 129

. However, this

phase may be totally eliminated by introducing a stent. The biological response to the

procedure is more complicated to eliminate, and it may consist of the following four

phases, 129-133

as described in Fig. 1.4.

i) Platelet Aggregation: Immediately after stent placement, endothelial denudation

and medial dissection results due to the mechanical injury of PCI [Fig. 1.4a]. The


injury causes platelet aggregation and activation, producing a countless number of

various cell-signaling factors, initiating an inflammatory cascade and releasing

adhesion molecules that cause a thrombus formation129,130

.

ii) Inflammatory Phase: Over the next few days to weeks, a variety of white cells

will gather at the injury site, secrete their own factors, and exert their own influence

on the healing tissue. The inflammatory response can persist for months [Fig. 1.4b].

iii) Proliferation Phase: The inflammatory phase stimulates smooth muscle cells

(SMCs) migration and proliferation, in an attempt to repair the wound. This process is

enabled by leukocytes cells releasing and activating tissue-digesting enzymes,

forming a path for the SMCs to move. SMCs migrate to the thrombus that acts as a

scaffold, providing the substrate for neointimal formation. The migrating SMCs form

an overgrown, obstructing scar [Fig. 1.4c].

iv) Late Remodeling Phase: The final mechanism of restenosis response is the late

remodeling of the vessel. This produces a neointimal layer, which is mainly formed

by proliferating SMC and extracellular matrix (ECM). Inflammatory mediators and

cellular elements contribute to trigger a complex array of events that modulates

matrix production and cellular proliferation. As the amount of scar develops [Fig. 1.4

d], blood flow is gradually reduced. Additionally, there is evidential re-

endothelialization of partial segments of the injured vessel surface.

Fig. 1.4. A schematic representation of the restenosis process. (a) Platelet Aggregation:

Immediate result of stent placement with endothelial denudation and platelet/fibrinogen (not

shown) deposition. (b) Inflammatory Phase: A variety of white cells will gather at the injury site.


(c) Proliferation Phase: Smooth muscle cells migrate and proliferate, creating the neointima in

an attempt to repair the wound. (d) Late Remodeling Phase: The neointima is changed from

predominantly cellular to a less cellular and more extracellular matrix-rich plaque. Kraitzer A, J

Biomed Mater Res B Appl Biomater. 2008

1.2.8 Strategies for combating in-stent restenosis

A considerable amount of research was invested in the prevention of in-stent

restenosis due to the substantial rate of the phenomenon. This interest prompted

various strategies that have been investigated and employed for the treatment of

restenosis. The stent strategy reduced restenosis caused by balloon angioplasty from

~50% to ~15% 129

. Early attempts at the prevention of restenosis focused on the

administration of antithrombotic agents, but there was limited success in trials134

.

With technologic advances and greater understanding of vascular pathobiology, novel

therapeutic strategies, such as local delivery of ionizing radiation, pharmacological

agents, and gene therapy, have been utilized to prevent coronary restenosis134,135

. It

became clear that although coronary stenting combined with antithrombotic therapies

had essentially eliminated the problem of elastic recoil, thrombus formation, and

vessel remodeling, in-stent restenosis remained a major problem129

. Today, it is

obvious that vascular SMC (VSMC) growth and migration trigger intimal

hyperplasia, which is the main cause of in-stent restenosis135

. Researchers have

adopted several strategies as below to address these issues.

i) Polymer coatings

Polymeric coatings on stents serve as a surface layer preventing its exposure to blood

and also to coat drugs. The most widely used materials for developing resorbable

stents or stent coatings on bare metal stents are the aliphatic polyesters including

poly-L-lactic acid, polyglycolic acid, poly-DL-lactic acid and poly-ε -caprolactone.

These polymers have been used as a coating material on stents to improve the

antithrombogenic properties of Ta137

, corrosion resistance of Ni–Ti138

and

biocompatibility of SS139

. A porcine model using poly-L-lactic acid demonstrated

minimal inflammatory and thrombotic response with good initial radial strength 140

.


Phosphorylcholine is a naturally occurring neutrally charged phospholipid, a

component of the plasma membrane. It causes less platelet activation than stainless

steel. Stents coated with phosphorylcholine have been shown to be non-inflammatory

and non-thrombogenic in-vitro and in-vivo for six months or longer136

.

Phosphorylcholine also provides a basis for a drug eluting coating. Clinical trials with

this and other polymer-based stents to date have yielded highly variable results.

ii) Inorganic/Ceramic coatings

Inorganic strategies may also have potential. Silicon carbide has been

investigated for its ability to alter the electrochemical properties of the stent surface.

It has been suggested that the initiation of thrombosis is at least partly due to the

degeneration of blood proteins by electron transfer to the metal. The ideal surface,

from this point of view, is a semi-conductor such as silicon carbide. But, being brittle,

silicon carbide can only be applied as a thin layer. Systematic testing of the effect of

the silicon-carbide coated Tensum (Biotronik) stent upon cytotoxicity, hemolysis,

mutagenicity and hemocompatibility produced favourable results when compared

with Palmaz–Schatz (Cordis) and HepaMed (heparin) coated Wiktor (Medtronic)

stents 141

. Tantalum stents coated with silicon carbide deployed in rabbit iliac arteries

demonstrated complete endothelialization with minimal intimal proliferation142

.

Placement of eight silicon carbide-coated Palmaz–Schatz stents into patients suffering

from abrupt closure post-PTCA showed, at coronary angiography the next day,

patency of all the stents with no visible thrombus 143

. A series of 165 patients with

215 stents carried out using the Tensum (Biotronik) tantalum, balloon expandable,

silicon carbide-coated stent deployed in a group at high risk of restenosis and

thrombosis demonstrated 2% stent thrombosis. At six months, 32% of patients (24%

of stents) had a cardiac event144

. Likewise, other inorganic coatings have been tested

to demonstrate useful properties. A ‗diamond-like‘ carbon-coated stent exposed to

flowing platelet-rich plasma, produced less platelet activation and deposition and ion

release than uncoated stents 145,146

. Gold was suggested to be the ultimate inert stent

coating. A 5 µm thick gold coating applied to a stainless steel stent indeed showed

more than a halving of adherent thrombus mass compared with an uncoated stent 147

.


However, disappointingly, a randomized study of 730 patients receiving a gold-

coated or bare stent revealed an excess of clinical events in the gold-coated group at

one year (24% vs 13%) 148

.

Biocompatible titanium-based coating was yet another material tested to coat

coronary stent surfaces. Previous examinations of the biocompatibility of various

metals have indicated that in addition to their excellent mechanical performance,

stainless steel and gold have an increased electrochemical surface potential, allowing

the transfer of electrons to proteins, such as fibrinogen. This promotes thrombus

formation and neointimal hyperplasia. In contrast, titanium, with its low

electrochemical surface potential, is biologically inert and has excellent

biocompatibility and inflammatory response149

. Titanium-nitride-oxide (TiNOX) is

proposed as a novel titanium alloy that renders metallic stents biologically inert and

prevents metallic foreign body reaction. Several preclinical and clinical studies were

taken up on this biocompatible coating offered by TiNOX on metallic stents of SS

and Cobalt Chromium. Porcine studies compared TiNOX-1 (ceramic) or TiNOX-2

(metallic) with uncoated stainless steel stents. Initial findings demonstrated reduced

neointimal formation and platelet adhesion 150

. A reduced neointimal hyperplasia up

to 50% at 6 week follow-up was observed. Clinical studies also demonstrated the

non-inferiority of TiNOX -coated stent in comparison with the first generation drug

eluting stents. A 3-year follow up report from the TITAX AMI trial demonstrated

better clinical outcome for TITANOX-coated stent compared with paclitaxel eluting

stents in terms of lower rate of Myocardial infarction (MI), cardiac death and stent

thrombosis in patients presented with acute MI151

.A 6-month follow-up study (TIBET

registry)showed excellent immediate clinical and angiographic outcome, with a low

incidence of major adverse cardiac events (MACE) at mid-term follow-up of Titan

stent implantation in 156 consecutive diabetic patients admitted to undergo

percutaneous intervention for at least one significant (50%) coronary lesion152

. A

similar follow up study conducted by the French Ministry of Health prospective

multicentre sought to explore the immediate outcome of the titanium-nitride-oxide-

coated bioactive stent, Titan2, in real-world practice and the incidence of major


cardiac events(EVIDENCE registry).In real-world practice, Titan2® stent

implantation revealed an excellent immediate outcome, with a low incidence of major

adverse cardiac events at 12-month follow-up153.

iii) Heparin-coated stents

Several anti-thrombogenic stent coatings 154,155

have been investigated,

with heparin being the most well known and extensively tested. Heparin has been

studied while covalently bound to the stent, as a ―passive‖ coating, and also as an

eluted drug. Heparin coated stents were associated with reduced platelet and

endothelial activation when compared to bare metal controls by plasma P-selectin and

E-selectin assessment in a small human study 156

.White cell and platelet activation in-

vitro have been shown to be reduced when heparin coated stents are compared to

gold, 157

SiC, 158,159

or bare metal stents, with prolongation of thrombosis time 160,161

.As

with some non-randomized series evaluating heparin coated stents, the clinical

incidence of sub acute stent thrombosis (SAT) was low in BENESTENT II

(randomized comparison of implantation of heparin coated stents with balloon

angioplasty in selected patients with coronary artery disease) although there was no

control bare stent group compared to the heparin coated Palmaz-Schatz group 162

.

Other randomized (COAST, heparin-coated stent placement for the treatment of

stenoses in small coronary arteries of symptomatic patients) 163

and non-randomized

comparisons to bare metal stents suggest similar SAT and ISR rates 164,165

.

iv) Drug eluting stents

Early difficulties with coronary stents included a risk of

early thrombosis (clotting) (Fig. 1.6A) resulting in occlusion of the stent166

. Coating

stainless steel stents with other substances such as platinum or gold did not eliminate

this problem167

. Drug eluting stents (DES) were developed and approved first in

2003, with significant initial success 168

. DES is capable of releasing single or

multiple bioactive agents from a polymeric coating on its surface into the bloodstream

and surrounding tissues. By adding a drug eluting coating, the rate of restenosis has

been reduced to about 5% or less169

. The drugs that may be useful in preventing ISR


fall into four major categories: anti-neoplastics, immunosupressives, migration

inhibitors, and enhanced healing factors24

. Heparin has been effective in reducing

both thrombosis and neointimal proliferation while sirolimus and paclitaxel were

mainly used for their anti-proliferative effects in blocking neointimal hyperplasia170

.

The anti-proliferative compounds commonly used in developing drug eluting stents

include paclitaxel, QP-2, actinomycin, statins, and many others. Paclitaxel was

originally used to inhibit tumor growth by assembling microtubules that prevent cells

from dividing. It has recently been observed to attenuate neointimal growth as well24

.

Paclitaxel stents, when placed in a culture of smooth muscle cells obtained from

human coronary atherosclerotic plaque, produced severe destruction of cytoskeletal

components of the cells, suggesting a possible strategy for in-vivo use, assuming the

problems of inflammation and radial strength can be overcome171

.

Immunosuppressive such as limus drugs are generally used to prevent the immune

rejection of allogenic organ transplants. The general mechanism of action of most of

these drugs is to stop cell cycle progression by inhibiting DNA synthesis. Everolimus,

sirolimus, tacrolimus (FK-506), ABT-578, interferon, dexamethasone, and

cyclosporine belong to this category. The sirolimus derived compounds appear to be

promising in their ability to reduce intimal thickening 24

.The first successful trials

were of sirolimus-eluting stents. A clinical trial in 2002 led to the approval of

sirolimus-eluting Cypher stent in Europe in 2002 172

.Soon thereafter, a series of trials

of paclitaxel-eluting stents (PES) led to FDA approval of the Taxus stent in 2004 173

.

The Xience V everolimus eluting stent was approved by the FDA in July 2008

and has been available in Europe and other international markets since late 2006.

However, late stent thrombosis has been a major concern with DES platforms due to

delayed endothelialization as seen clearly from Fig. 1.6B. The newer generation of

DES stents which utilize strategies such as biodegradable polymeric coatings,

biodegradable stents and polymer-free approaches can be a possible solution to this

problem. To date, the most successfully tested drug eluting stents have been coated

with synthetic polymers: poly-n-butyl methacrylate and polyethylene–vinyl acetate


with sirolimus, and a poly(lactide-co-caprolactone) copolymer with paclitaxel eluting

platforms 24

.

Fig. 1.6 (A) Schematic representation of late stent thrombosis associated with DES. (B) Partially

endothelialized DES in comparison to BMS ; EES- everolimus-eluting stent; PES- paclitaxel-

eluting stent; SES- sirolimus-eluting stent; ZES- zotarolimus-eluting stent (C) Different

generations of DES. Image courtesy: medimoon.com/2013; M. Joner et al., J Am Coll Cardiol.

2008; M.J. Patel et al, Acta Pharm, 2012

In addition to coated metallic stents, a new category of completely

biodegradable stents appeared in the early 2000. Bioabsorbable DES helps to achieve

excellent acute and long-term results, but disappear completely within months,

thereby avoiding the need for prolonged dual antiplatelet therapy. In the late 1990s, a

bioabsorbable (Igaki–Tamai, Japan) stent, made of a high-molecular-mass poly-L-

lactic acid (PLLA), was implanted in 15 patients (25 stents) to evaluate the feasibility,

safety, and efficacy of the PLLA stent174

. No major cardiac events, except for repeat

A B

C


angioplasty, developed within 6 months. A still larger study of 50 elective patients

(63 lesions, 84 stents) also showed promising results175

.

The IDEAL biodegradable stent (Bioabsorbable Therapeutics, Inc, Menlo

Park, CA) consisting of a backbone of poly-anhydride ester based on salicylic acid

and adipic acid anhydride and a coating of sirolimus, potentially rendered the stent

with both anti-inflammatory and antiproliferative properties176

. Most notably, the

polymer was associated with reduced inflammation compared with a standard BMS

and Cypher stent, which was likely due to the anti-inflammatory properties of

salicylic acid following absorption by the vessel wall after its release. Drug elution

was found to be complete after 30 days, and complete stent degradation occurred over

a 9- to 12-month period. In July 2009, the results from 11 patients enrolled in the

multicenter first-in-humans Whisper trial showed stent safety and confirmed

structural integrity of the stent, with no evidence of acute or chronic recoil.

Unfortunately, insufficient neointimal suppression was demonstrated, which could be

possibly due to the consequence of inadequate drug dosing, particularly considering

that the surface area dose of sirolimus is only a quarter of that found on the Cypher

stent177

. ABSORB (Abbott Vascular, Santa Clara, CA, USA)is a Bioresorbable

Vascular Scaffold (BVS) system that elutes everolimus in a similar way to XIENCE

V and then resorbs naturally into the body leaving no permanent scaffold. The safety

and performance of the Absorb BVS system was previously established in 131

patients from the first-in-man ABSORB trial. Further, clinical outcomes after twelve-

month follow up trials in 512 patients (ABSORB EXTEND) showed low rates of

MACE and scaffold thrombosis 178

.

The polymer-free approaches of developing DES include carbon based

films, micro structured reservoirs, ceramic coatings, bioabsorbable coatings, drug

coatings etc 179

. Limited evidence exists regarding the long-term performance of

polymer-free (PF) DES in comparison to permanent polymer DES. Recent studies

report the results of the long term performance and clinical outcomes of different

polymer-free and polymeric DES after randomized trials. The LIPSIA Yukon trial

randomized 240 patients with diabetes mellitus to a polymer-free sirolimus eluting


stent (Yukon Choice, Translumina) versus a polymer-based paclitaxel-eluting stent

(Taxus Liberté, Boston Scientific)180

.In another randomized trial, the 5-year efficacy

and safety of a PF sirolimus-eluting stent versus a permanent polymer paclitaxel-

eluting stent (PES) in the setting of the Intracoronary Stenting and Angiographic

Restenosis-Test Equivalence Between Two Drug-Eluting Stents (ISAR-TEST) was

investigated in a total of 450 patients undergoing percutaneous coronary intervention

181. Overall, both trials demonstrated no significant differences in the rates of

myocardial infarction, definite stent thrombosis, target lesion revascularization, non-

target vessel revascularization, and stroke between patients assigned to the polymer-

free and polymeric DES platforms. The polymer-free platforms were unsuccessful in

establishing any significant advantage over the polymer-based DES in terms of the

primary end point late lumen loss, while the 5 years follow up showed similar clinical

for both181

.

An updated meta-analysis of bioabsorbable (BP-DES) versus durable

polymer drug-eluting stents (DP-DES) was performed in 20,005 patients with

coronary artery disease by Lupi et al. in 2014. The study concluded significantly

reduced late lumen loss and late stent thrombosis rates for BP-DES in comparison to

DP-DES, without clear benefits on harder endpoints. The efficacy of BP-DES in

preserving lumen patency seemed larger for sirolimus and novolimus DES182

.

Likewise, bioabsorbable polymer-based biolimus-eluting stents established superior

clinical outcomes as against BMS and first-generation DES, and similar rates of

cardiac death/MI, and Target Vessel Revascularization (TVR) compared to second-

generation DP-DES, but higher rates of definite ST than cobalt-chromium

everolimus-eluting stents183

.

1.2.9 Limitations of polymer-based drug eluting stents

Drug eluting stents with durable polymer coatings have by now replaced the

bare metallic stents by significantly reducing late lumen loss and have revolutionized

the practice of percutaneous coronary intervention. However, there are still

unresolved issues pertaining to the use of polymer based drug eluting stents.


i) Polymeric degradation

The first generation DES such as, Cypher and Taxus employed thick durable

polymeric layer to incorporate the antiproliferative drugs and to render their

controlled release. The presence of these durable polymers has been linked to

inflammatory responses and local toxicity in preclinical analysis184, 185

. Furthermore,

durable polymers used in first-generation DES were associated with mechanical

complications such as polymer delamination, ―webbed‖ polymer surface etc.,

leading to stent expansion issues(Fig. 1.5A) as well as non-uniform coating, resulting

in erratic drug distribution(Fig. 1.5B) 140

.

Fig. 1.5 Limitations of stent polymeric coatings (A) Degradation/ delamination (B) non uniform

coating Image courtesy: Waksman, Interventional Cardiology, 2010

In recent years, considerable trials have been made in clinical research on

the development of novel drug carrier systems including absorbable (or

biodegradable) polymers and non-polymeric stent surfaces. These new platforms

provide better deliverability, radio-opacity, flexibility, and radial strength and also

facilitate reduced dosage of currently used antiproliferative drugs. However, such

polymeric coatings are also afflicted by certain issues as discussed below.

ii) Delayed endothelialization

The long-term safety of polymer-based DES has been a major concern due to

impaired arterial healing, which in turn leads to late stent thrombosis. The anti-

proliferative drugs from DES would hamper the growth of endothelial cells apart

from inhibiting smooth muscles and platelets [Ref].This affects the re-

A B


endothelialization of stent struts, leaving the bare stent surface to contact with blood

provoking inflammatory cell infiltration and/or cause long-term drug sequestration

within the arterial wall186

. However new research on using more endothelial cell

friendly drugs to coat stents such as everolimus and zotarolimus have resulted in

improved endothelial coverage after implantation. In addition, literature on the use of

polymer-besides for interventions involving multiple stents (such as bifurcations or

overlapping stents) indicate that, there are chances for local arterial toxicity

aggravation that occurs when drug and polymer concentrations are substantially

increased187

.

iii) Stent thrombosis

Polymer-based drug eluting stents have been associated with an increased

risk of very late stent thrombosis compared with bare metal stents. This could be

endorsed fairly to hypersensitivity reactions towards the polymeric coatings, resulting

in inflammation, late stent malapposition, and stent thrombosis 24, 186

. Moreover,

delayed stent endothelialization results in platelet aggregation and stent thrombosis at

a later stage, mainly after the exhaustion of the drugs (late stent thrombosis) 188

. This

makes the patients depend on prolonged anti-platelet therapy after coronary stent

implantation.

1.2.10 Role of regenerative medicine and nanotechnology for

cardiovascular implants

The endothelial monolayer that lines the normal blood vessel serves as a

bioregulator of cardiovascular physiology. The vascular endothelium is versatile and

multifunctional with many synthetic and metabolic properties. These include the

regulation of thrombus and platelet activation, adhesion and aggregation 189

as well as

modulation of vascular tone and blood flow 190

. It also controls SMC migration and

proliferation 191

. ECs secrete and express numerous growth factors, extracellular

matrix products, anti-thrombotic and pro-coagulant factors. ECs are intimately

involved in maintaining a non-thrombogenic blood-tissue interface. Re-

endothelialized segments of artery are often associated with less neointimal


thickening and the absence of an intact EC lining therefore predisposes

cardiovascular implants; including bypass graft and stents; to platelet deposition,

thrombus and implant failure. As a result, investigators have developed methods to

promote the endothelialization of vascular grafts prior to implantation by

transplantation of ECs in-vitro, a process called EC seeding. The cellular engineering

approach called ‗seeding‘ involves lining the lumen of the graft in-vitro with EC 192,

193. Recent advances in science and technology, especially progressive development

in nanoscience and nanotechnology, offer novel nano-structured materials with

enhanced characteristics which make them the material of choice for various

applications 194, 195

.

Nanotechnology is concerned with manipulation at the molecular or atomic

level to provide useful applications 196,197

. The main unifying theme is the control of

matter on a scale smaller than 1 micrometer, normally between 0.1-100 nanometers,

as well as the fabrication of devices on this same length scale. Materials reduced to

the nanoscale can show very different properties compared to what they exhibit on a

macroscale, enabling unique applications197

. To cite a typical example, a material

such as gold, which is chemically inert at normal scales, can serve as a potent

chemical catalyst at nanoscales. Much of the fascination with nanotechnology stems

from these unique quantum and surface phenomena that matter exhibits at the

nanoscale. The difficulties involved in fulfilling the numerous ideal characteristics

using traditional synthetic materials have lead biomaterials research towards the fields

of nanotechnology and tissue engineering.

In cardiovascular arena, researchers have utilized nanotechnology to create

stent surfaces that express novel advantageous properties such as anti-

thrombogenicity and biostability, either through surface texturing or via the

incorporation of functionalities on material surfaces198

. Surface characteristics of the

biomaterial used for stent fabrication is well known to influence stent-biology

interactions. The primary surface characteristics of a stent material, which influence

thrombosis and neointimal hyperplasia, include surface energy, surface texture,

surface potential, and the stability of the surface oxide layer 199,200

. An in-vitro study


showed that the adherence, growth and proliferation of endothelial cells on Ta films

were much better than on 316L SS and Ti films 201

. In another study, the sputter

coating of a Ti–Ta target produced a surface that showed better endothelialization

because of the changes in the microstructure of the natural Ti-oxide film produced 202

.

The above cases imply that changes in surface energies or surface texturing of a

biomaterial can induce changes in their biological behavior. However, the nature of

the coating although biocompatible, should not lose its integrity during the stent

placement and expansion, causing adverse effects. Also, similar to surface charge, an

optimal range of surface energy and surface roughness can promote better

endothelialization on the metallic biomaterials.

It is well established in literature that nanoscale surface topography has

significant effects on cellular behavior. This has been corroborated by the recent

studies by our research group on nanostructuring of metallic titanium showing that

osteoblast/endothelial cells do respond to nanoscale surface differently203,204,70

.

Literature reports also provide ample evidences for the effects of nanostructuring in

promoting enhanced cellular adhesion in-vitro and in-vivo. A nanoscale rough surface

of poly(lactic-co-glycolic acid) (PLGA) produced by NaOH treatment when

compared with smooth PLGA substrate, enhanced rat aortic SMC adhesion and

proliferation, and decreased rat aortic EC adhesion and proliferation 205,206

. Similar

effect was observed on nanostructured Ti versus conventional Ti 207

. Human coronary

artery SMCs were found to adhere and proliferate significantly more on aligned

electrospun poly (l-lactid-co-ε-caprolactone) or P(LLA-CL) nanofibers than on

randomly oriented nanofibers and the P(LLA-CL) film as well as tissue culture PS

(TCPS) control 208

. These altered cell responses to nanoscale topography stems from

several factors such as (i) changes in protein adsorption at the nanoscale surface, (ii)

the dimension of the nanoscale structures which regulates cell adhesion and

spreading, subsequently gene expression, proliferation and differentiation and (iii) the

nanostructure-induced alteration of the elasticity of substrate surface 209,210

. These

effects are beautifully illustrated in Fig. 1.7 wherein surface roughness and stiffness

are found to influence the attachment and proliferation of cardiomyocytes.


Fig. 1.7 Schematic representation of the effect of nanorough surface on cell behavior compared

to a flatter surface. (A) Shows the adsorption of ECM proteins immediately when substrates

implanted or soaked in media. (B) Indiactes the cardiomyocytes adhered to the substrates and

begin to grow. (C) Due to mimicking native myocardium ECM in surface features, more

cardiomyocytes on nanorough stiff substrates were adhered and grown than on conventional and

plain nanorough substrates.Stout et. al Int J Nanomedicine. 2012

1.2.11 Titania nanosurface modification

Current efforts are directed towards maximizing biocompatibility of

the implant material, which means to optimize cell adhesiveness and to support

physiological cell reactions such as spreading, proliferation, migration and

differentiation. Major strategies to improve cell adhesiveness of implant biomaterial


include surface roughening, etching or modifying by physical and chemical methods,

and/or coating with adhesive proteins of the extracellular matrix such as collagens,

fibronectin, and laminins. In addition, evidence is accumulating that not only the

surface chemistry of the biomaterial, but also the surface topography at nanoscale is a

critical parameter for cellular recognition of biomaterials. It is well known that the

native oxide layer on titanium (Ti) implants is responsible for its superior

biocompatibility and tissue integration. Recent efforts have targeted titanium dioxide

(TiO2 or titania) as a good candidate for surface modification at the nanoscale,

leading to improved nanotextures for enhancing host integration properties. TiO2

nanotubes on a Ti implant surface with different diameters (30, 50, 70, and 100 nm)

created by a simple electrochemical anodization process showed minimal

inflammatory response of macrophages in-vitro, clearly demonstrating the nanosize

effect on immune cells 211

. Studies also demonstrated that bone marrow stem cells

(MSC) respond to titanium chips covered with TiO2 nanotubes in a size-dependent

manner, with a maximum of cell adhesion, proliferation, and migration rates on 15

nm diameter nanotubes indicating the influence of nanosructural dimensions on cell

behavior 212,213

. It was demonstrated that fabrication of such surface nanofeatures by

compressing Ti nanoparticles 207

or by creating rationally designed nanopatterns 214

on Ti resulted in a substantial increase in endothelial adherence compared to bare Ti.

Surface modification of Ti by the creation of ordered TiO2 nanotubes has resulted in

improved osteoblast attachment, function, and proliferation215,216

.

In a previous study conducted by our own group, osteoblast cell response

was found to be greatly improved on hydrothermally modified TiO2

nanostructures203,204

. Nanomodification of Ti generating TiO2 surfaces has proved

beneficial in promoting endothelial proliferation as well as migration onto stent

surfaces214, 216, 217

Recent studies also suggest that TiO2 nanotubes are beneficial for

endothelial cells, encouraging more ECM secretion and functions, while inhibiting

VSMC growth 101,102

.Nanostructures developed by coating the Ti surface with rosette

nanotubes of DNA base analogues have been reported to influence endothelial

attachment and spreading, 216

implying in general that nanostructuring of metallic


surfaces might be a promising approach to faster endothelialization. Brammer et al

also reported that, TiO2 nanotubular structures have influence on endothelial

functionality with increased Nitric oxide/endothelin-1 ratio pointing to their

antithrombogenicity 217

. A recent report on stainless steel 316L coated with nano-

structured TiO2 layer exhibited improved blood compatibility, in terms of both blood

platelet activity and coagulation cascade, which can decrease the thrombogenicity of

blood contacting devices which were made from stainless steel 218

.

1.2.12 Micro and nanomodifications on coronary stents

Surface topography as mentioned earlier, is known to modulate cell response

significantly219

. The technology of creating nano and micro features on coronary

stents to modulate vascular cell behavior has been employed by various stent

companies. Recent examples of stents that contain a microporous coating include the

Corel-C™ stent that is fabricated from CoCr with a carbon nanoparticle coating

(Relysis Medical, India) 220

. Microporous surfaces have also been incorporated into

polymer based bioabsorbable stents 221

. Sandblasting has been used to create pores

between 1 to 100 μm on the surface of stainless steel stents, an example of which is

the Yukon™ stent platform (Translumina, Hechingen, Germany) 222

. The

BioFreedom (Biosensors Inc.) is a 316L stainless steel stent platform without polymer

and coated with biolimus A9 (Fig. 1.8A).

Fig. 1.8 Micro and nanomodifications on coronary stents. (A) Without polymer. The image of

scanning electron microscope showing the surface of the BioFreedom stent with biolimus A9

A

B

C Ci

Cii

Ciii

Civ


impregnated micropores in the abluminal side of the strut. (B): SEM view of a Setagon stent (x

20,000) showing alumina nanoporous stent surface (Ci) Scanning electron micrographs of

microporous Hap Coating, (Cii) cross section of the Hap coating (Ciii) final coating including the

Hap filled with sirolimus formulation, (Civ) cross section of the final coating. (Image courtesy:

Interven Cardiol, 2011, Future Medicine Ltd; Informa Healthcare 2007; A. Abizaid, Circulation:

Cardiovascular Interventions, 2010)

Preclinical studies have reported lower lesion scores, fewer struts with fibrin,

granulomas and giant cells, a significantly lower percentage diameter stenosis and

greater endothelialization than the sirolimus eluting Cypher stent223

. However, the

rapamycin eluting Yukon™ stent was found to be clinically inferior to the Cypher

stent, as well as a version of the Yukon™ stent coated in a drug containing

bioabsorbable polymer222

. At least two examples exist of nanoporous stent surfaces

are in the market. The Jomed™ coating consists of a thin aluminum base layer, which

is then subjected to an acidic solution that converts the aluminum into a thin ceramic

nanoporous aluminum oxide 224

(Fig. 1.8B, Setagon stents). A recent report suggests

that particle debris may be released from the stent surface 225

. MIV Therapeutics, Inc.

has developed a stent with a hydroxyapatite (Hap) coating that is 0.30 to 1 μm in

thickness with a porosity of 40-60% in volume 226

(Fig. 1.8C). This stent has shown

promising responses in both animal studies and in an initial clinical study after

adsorption of sirolimus 227

.


Summarizing, various strategies have been adopted by researchers to address the

limitations of bare metallic stents as tabulated in Fig 1.9.

Fig. 1.9 Strategies to combat in-stent restenosis

Drug coated stents

• Anti-neoplastic– Paclitaxel, ABT 578 • Immunosuppressant– Sirolimus– Tacrolimus

Inorganic coatings HAp, Iridium oxide, TiN & TiO

2 coatings, carbon-based films such as

diamond-like carbon (DLC) and carbon nitride (CN), Al2O

3, silicone

carbide

Endothelialization promoting modifications

• Antibody immobilization (CD31) for EPC capture • Bioactive peptide, Estradiol– VEGF

Bio-chemical modifications

• Heparin coating • Biomimetic polymers – Phosphorylcholine,

Hyaluronic acid (HA)


1.3 Aim and overall hypothesis of the thesis:

This thesis work focuses on the development of inorganic (ceramic) coatings

of bio/hemocompatible, nanotextured titania on Stainless Steel substrates and

thereafter on stents as a cell-friendly approach to realize three-fold benefits:

provide an anti-thrombotic layer

inhibit platelet aggregation

concurrently improve endothelialization and inhibit SMC over-proliferation.

AIM:

The aim of the study is to improve the current limitations of existing bare

metallic stents through a simple surface nanostructuring approach, without the use of

any anti-proliferative drugs or polymers, to combat in-stent restenosis.

HYPOTHESIS:

Surface modification of bare metal stents by creating hemocompatible

titania nanostructural features can present a more cell-friendly

environment for vascular cells, aiding in rapid functional

endothelialization, without inducing hyperproliferation of smooth muscle

cells or platelet adhesion, thereby regulating the causative factors for in-

stent re-stenosis and thrombosis.

1.4 Specific objectives of the thesis

The following are the specific objectives of this thesis work:

Fabrication of an array of uniform nanofeatures of TiO2 on metallic titanium

substrates/stent prototypes via hydrothermal processing by varying different

processing parameters such as time, temperature and concentration of the

reaction medium.


Comparative in-vitro evaluation of endothelial and smooth muscle cell response

(proliferation, apoptosis and cytoskeleton organization, etc.) on various

nanosurfaces, functionality of the formed endothelium and time for complete

endothelialization.

Checking for in-vitro hemocompatibility of nanomodified Ti plates/stents under

static and dynamic flow conditions, and assessment of the same on

endothelialized surfaces.

Optimization of hydrothermal process for nanotexturing of TiO2 onto Ti coated

stainless steel substrates for better biological response, corrosion resistance and

mechanical integrity.

Assessment of ion leaching profiles, corrosion resistance and mechanical

properties of the titania coating on stainless steel substrates as per ISO standards.

Translation of TiO2 nanostructured coating on to stainless steel coronary stents.

Crimping and expansion tests to prove the adhesion and stability of TiO2 coating

on SS stent and evaluation of the coating durability in-vitro under high shear

stress dynamic flow.


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