+ All Categories
Home > Documents > Swamy Science

Swamy Science

Date post: 14-Apr-2018
Category:
Upload: deva-raj
View: 223 times
Download: 0 times
Share this document with a friend

of 22

Transcript
  • 7/30/2019 Swamy Science

    1/22

    Biomaterials 28 (2007) 16891710

    Review

    Coronary stents: A materials perspective

    Gopinath Mania, Marc D. Feldmanb,c, Devang Patelb, C. Mauli Agrawala,

    aDepartment of Biomedical Engineering, College of Engineering, The University of Texas at San Antonio, One UTSA Circle,

    San Antonio, TX 78249 0619, USAbDivision of Cardiology, Department of Medicine, The University of Texas Health Science Center at San Antonio, 7703 Floyd Curl Drive,

    San Antonio, TX 78229 3900, USAcThe Department of Veterans Affairs South Texas Health Care System, San Antonio, TX 78229 4404, USA

    Received 5 September 2006; accepted 29 November 2006

    Available online 22 December 2006

    Abstract

    The objective of this review is to describe the suitability of different biomaterials as coronary stents. This review focuses on the

    following topics: (1) different materials used for stents, (2) surface characteristics that influence stentbiology interactions, (3) the use of

    polymers in stents, and (4) drug-eluting stents, especially those that are commercially available.

    r 2006 Elsevier Ltd. All rights reserved.

    Keywords: Stent; Surface treatment; Surface modification; Drug delivery

    Contents

    1. Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1690

    2. Metallic stents. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1690

    2.1. 316L SS . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1691

    2.2. PtIr alloys . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1691

    2.3. Ta . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1691

    2.4. Ti . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1692

    2.5. NiTi. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1692

    2.6. CoCr alloy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1693

    2.7. Biodegradable metallic stents . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1693

    2.7.1. Pure Fe . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1693

    2.7.2. Mg alloys. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1693

    3. Surface characteristics . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1694

    3.1. Surface energy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1694

    3.2. Surface texture . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 16943.3. Surface potential. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1694

    3.4. Stability of surface oxide layer . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1695

    4. Rationale for coatings . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1695

    4.1. Types of coatings . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1695

    4.1.1. Inorganic coatings . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1695

    4.1.2. Endothelial cells . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1696

    4.1.3. Porous materials. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1696

    5. Polymers. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1697

    ARTICLE IN PRESS

    www.elsevier.com/locate/biomaterials

    0142-9612/$ - see front matterr 2006 Elsevier Ltd. All rights reserved.

    doi:10.1016/j.biomaterials.2006.11.042

    Corresponding author. Tel.: +1 210458 5526; fax: +1 210458 5556.

    E-mail address: [email protected] (C.M. Agrawal).

    http://www.elsevier.com/locate/biomaterialshttp://localhost/var/www/apps/conversion/tmp/scratch_8/dx.doi.org/10.1016/j.biomaterials.2006.11.042mailto:[email protected]:[email protected]://localhost/var/www/apps/conversion/tmp/scratch_8/dx.doi.org/10.1016/j.biomaterials.2006.11.042http://www.elsevier.com/locate/biomaterials
  • 7/30/2019 Swamy Science

    2/22

    5.1. Biostable polymers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1697

    5.2. Biodegradable polymers. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1697

    5.3. Copolymers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1698

    5.4. Biological polymers. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1698

    5.4.1. PC . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1698

    5.4.2. HA . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1698

    5.4.3. Fibrin . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1699

    6. Rationale for DES. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 16996.1. Techniques for drug-loading and release kinetics . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1699

    6.2. DES . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1699

    6.2.1. Heparin-coated stents . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1699

    6.2.2. Sirolimus-eluting stents (SES) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1700

    6.2.3. Paclitaxel-eluting stents (PES) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1701

    7. Conclusion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1702

    References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1703

    1. Introduction

    Percutaneous transluminal coronary angioplasty

    (PTCA) is an invasive procedure performed to reduce

    blockages in coronary arteries [1,2]. However, restenosis

    follows PTCA in 3040% of coronary lesions within 6

    months [3,4]. Although providing intra-arterial support

    with bare metal stents (BMS) dramatically improves the

    angiographic and clinical outcome of patients to a

    restenosis rate of 2030% [3,4], in-stent restenosis still

    remains a major limitation for this approach with

    exaggerated intimal hyperplasia [5]. The biology of rest-

    enosis in stents includes plaque redistribution, thrombosis,

    and neointimal hyperplasia [6]. The basic mechanisms [79]underlying thrombus formation and neointimal muscle cell

    proliferation, followed by extracellular expansion are

    understood to some extent, but the basic biology of

    restenosis still remains an active area of research. As a

    result of the inadequacies of BMS, different kinds of

    materials, designs, and techniques have been explored to

    further optimize stent design. Coronary stents developed to

    date can be grouped in four categories: bare metallic stents,

    coated metallic stents, biodegradable stents and drug-

    eluting stents (DES). 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 [10,11].

    This review evaluates the pros and cons of choosing

    different materials for the manufacture of coronary stents.

    The physical properties of each material that are relevant

    for this application are discussed. The influence of a

    materials surface characteristics on the biology of rest-

    enosis will be discussed as well. A variety of coating

    materials are commonly used in an attempt to improve the

    performance of stents; including inorganic materials,

    polymers, endothelial cells, and porous ceramics. The role

    of these different types of coatings is described in detail.

    The materials and the coating techniques used in commer-

    cially available DES are described. A list of ideal

    characteristics for coronary stents and the materials and

    processes that best meet these requirements are tabulated

    in the concluding section. The physical design of a stent,

    another important parameter, is not covered here as the

    discussion is confined to biomaterials.

    2. Metallic stents

    Balloon expandable stents should have the ability to

    undergo plastic deformation and then maintain the

    required size once deployed [12]. Self-expanding stents,

    on the other hand, should have sufficient elasticity to becompressed for delivery and then expanding in the target

    area [12]. The characteristics of an ideal stent have been

    described in numerous reviews [1315]. In general, it should

    have (1) low profileability to be crimped on the balloon

    catheter supported by a guide wire; (2) good expandability

    ratioonce 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 recoilonce implanted,

    the stent should be able to overcome the forces imposed by

    the atherosclerotic arterial wall and should not collapse; (4)

    sufficient flexibilityit should be flexible enough to travel

    through even the smaller diameter atherosclerotic arteries;

    (5) adequate radiopacity/magnetic resonance imaging

    (MRI) compatibilityto assist clinicians in assessing the

    in-vivo location of the stent; (6) thromboresistivitythe

    material should be blood compatible and not encourage

    platelet adhesion and deposition; and (7) drug delivery

    capacitythis has become one of the indispensable

    requirements for stents of the modern era to prevent

    restenosis.

    Generally, the metals commonly used for manufacturing

    stents are 316L stainless steel (316L SS), platinumiridium

    (PtIr) alloy, tantalum (Ta), nitinol (NiTi), cobalt

    chromium (CoCr) alloy, titanium (Ti), pure iron (Fe),

    ARTICLE IN PRESS

    G. Mani et al. / Biomaterials 28 (2007) 168917101690

  • 7/30/2019 Swamy Science

    3/22

    and magnesium (Mg) alloys. These are briefly discussed

    below:

    2.1. 316L SS

    Whether it is a bare stent or with a coating material, 316L

    SS is the most commonly used metal for stents. It has well-

    suited mechanical properties (Table 1) and excellent

    corrosion resistance (carbon contento0.030 wt%), making

    it the preferred material for this application [12]. However,

    the clinical limitations of using 316L SS are its ferromag-

    netic nature (6065 wt% pure Fe) and low density. These

    properties make SS a non-MRI compatible and poorly

    visible fluoroscopic material. Also, biocompatibility is anissue with bare SS stents. The weight percentage of nickel,

    chromium, and molybdenum in 316L SS are 12, 17, and 2.5,

    respectively [16]. Allergic reactions to the release of nickel

    can occur among SS implants [17]. In particular, the release

    of nickel, chromate, and molybdenum ions from SS stents

    may trigger local immune response and inflammatory

    reactions, which in turn may induce intimal hyperplasia

    and in-stent restenosis [18]. SS stents made of grades with

    lower nickel content can reduce the concern over allergic

    reactions to nickel. A number of SS grades with low nickel

    concentration (4.59%) are available [16]. However, it has

    been shown that higher nickel content (1014%) can be

    advantageous in decreasing the ferromagnetic properties of

    SS by stabilizing Fe in a non-magnetic state [19]. Hence, the

    SS grades (321 and 321H) with optimal nickel (10.5%) and

    carbon (0.08%) concentrations are promising. The addition

    of Ti (0.4%) in these grades makes them more attractive. A

    variety of materials have been used as a coating for 316L SS

    stents, mainly to circumvent its visibility limitations and to

    improve its biocompatibility by preventing the release of

    ions from the metal surface. However, the supremacy of

    316L SS platforms for making stents is evident from

    Table 2. Out of the eight coronary stents approved by the

    US Food and Drug Administration (FDA), seven are made

    from 316L SS.

    2.2. PtIr alloys

    An alloy of 90% platinum and 10% iridium was used for

    making bare stents and successfully implanted in animal

    models [20,21]. These stents showed excellent radiopacity

    [20] and it is even possible to take the three-dimensional

    image of the lumen of these stents using MRI [22]. The

    artifacts produced by the PtIr alloy in MRI are much

    lower when compared with 316L SS stents [19,22]. The

    presence of iridium in the alloy could pave the way for

    potential applications in radioactive stents [21]. In general,

    these alloys show excellent corrosion resistance but poor

    mechanical properties [23,24]. Although a reduction in

    both thrombosis and neointimal proliferation with lessinflammatory reactions was observed for these stents, their

    recoiling percentage was higher (16%) than the 316L SS

    stents (5%) [20,21]. Though a human clinical trial [25]

    encouraged the use of these stents as safe and effective, the

    literature on the biocompatibility and haemocompatibility

    of PtIr (90/10) alloys remains limited.

    2.3. Ta

    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 [26,27]. It has been coated on a 316L SS surfaces to

    improve corrosion properties, thereby enhancing the

    biocompatibility of 316L SS [28]. It has excellent fluoro-

    scopic visibility because of its high density. It is an MRI

    compatible material as it produces no significant artifacts

    because of its non-ferromagnetic properties [29,30]. Ta is

    also known for its good biocompatibility [31,32]. Enhanced

    hemocompatibility was achieved by adding Ta to Ti oxide

    and the films showed improved endothelialization rate as

    the percentage concentration of Ta increased [33,34].

    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

    ARTICLE IN PRESS

    Table 1

    Mechanical properties of the metals that are used for making stents

    Metal Elastic modulus (GPa) Yield strength (MPa) Tensile strength

    (MPa)

    Density

    (g/cm3)

    References

    316L stainless steel (ASTM F138 and

    F139; annealed)

    190 331 586 7.9 [23,59]

    Tantalum (annealed) 185 138 207 16.6 [23]

    CpTitanium (F67; 30% cold worked) 110 485 760 4.5 [23,59]

    Nitinol 83 (Austenite phase) 195690 (Austenite

    phase)

    895 6.7 [23,37]

    2841 (Martensite phase) 70140 (Martensite

    phase)

    Cobaltchromium (ASTM F90) 210 448648 9511220 9.2 [23,37,59]

    Pure iron 211.4 120150 180210 7.87 [269]

    Mg alloy (WE43) 44 162 250 1.84 [71,270]

    G. Mani et al. / Biomaterials 28 (2007) 16891710 1691

  • 7/30/2019 Swamy Science

    4/22

    because of its poor mechanical properties. Since the yield

    strength of Ta is closer to its tensile strength (Table 1),

    these stents have a higher possibility of breaking during

    deployment. Hence, the pressure applied for the deploy-

    ment of these stents is usually low and this might result in

    recoiling. The recoiling percentage was significantly higher

    for Ta stents when compared with 316L SS stents andresulted in enhanced neointimal formation [35]. Although

    no Ta-based stents have been approved by the FDA for

    general use to date, Cordis (Johnson & Johnson, USA) has

    used a bare Ta stent in clinical trials and released this stent

    commercially in Japan, Canada and Europe [36].

    2.4. Ti

    Ti and its alloys have been extensively used in orthopedic

    and dental biomedical applications because of their

    excellent biocompatibility [24,37]. The highly stable surface

    oxide layer provides excellent corrosion resistance [24,37].

    However, Ti is not commonly used for making stents.

    Although Ti and CoCr both have high yield strength in

    approximately the same range, Ti has a significantly lower

    tensile strength (Table 1). 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. This is one of the

    reasons why CoCr is used for making stents and not Ti.

    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 impactas 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 and there-

    by reducing neointimal hyperplasia [38]. The Titan stent

    (Hexacath, France), which has implemented this coating

    technique has shown promising results in human clinical

    trials [39,40]. Also, Ti-based Ta and niobium alloys, which

    have potential applications for stents, showed excellent

    haemocompatibility [41]. One of the Ti alloys which is

    extensively used for making stents is NiTi.

    2.5. NiTi

    NiTi constitutes 49.557.5 at% nickel and the remain-

    ing is Ti [42]. It is used for fabricating self-expanding stents

    mainly because of its shape memory effect. Self-expanding

    stents have a smaller diameter at room temperature and

    expand to their preset diameter at body temperature [43].

    NiTi is plastically deformed at room temperature

    (martensitic phase) and crimped on to the delivery system

    [37,42,43]. After implantation it regains its original shape

    (already memorized austenite phase according to the

    diameter of the target vessel) and conforms to the vessel

    wall because of the increase in temperature inside the body

    ARTICLE IN PRESS

    Table2

    FDAapprovedcoronarystents

    St

    entname

    Manuf

    acturer

    Barestentmaterial

    Coating

    FDAapprovaldate

    References

    BiodivYsio

    TM

    AS

    Biocom

    patiblesCardiovascular

    Inc.CA

    316Lstainlesssteel

    Cross-linkedphosphorylcholine

    September2000

    [271]

    BeStentTM

    2

    Medtronic,Inc.,Minnesota

    316Lstainlesssteel

    Nil

    October2000

    [272]

    CYPHERTM

    Cordis

    Corporation,FL

    316Lstainlesssteel

    Firstcoat:ParyleneC;second

    coat:mixtureof

    polyethylene-co-

    vinylacetate,po

    lyn-butyl

    methacrylate,an

    dSirolimus;third

    coat:mixtureof

    polyethylene-co-

    vinylacetate,po

    lyn-butyl

    methacrylate

    April2003

    [230]

    M

    ULTI-LINKVISIONTM

    GuidantCorporation,CA

    L-605cobaltchromiumalloy

    Nil

    July2003

    [65]

    NIRflex

    TM

    MedinolLtd.,Israel

    316Lstainlesssteel

    Nil

    October2003

    [273]

    TAXUSTM

    Express2TM

    Boston

    ScientificCorporation

    316Lstainlesssteel

    Mixtureofpoly(styrene-b-

    isobutylene-b-styrene)triblock

    copolymerandpaclitaxel

    March2004

    [274]

    LiberteTM

    MonorailTM

    Boston

    ScientificCorporation,

    MN

    316Lstainlesssteel

    Nil

    April2005

    [275]

    Rithron-XR

    BiotronikGmbH,Germany

    316Lstainlesssteel

    Amorphoussilic

    on-carbide

    April2005

    [276]

    G. Mani et al. / Biomaterials 28 (2007) 168917101692

  • 7/30/2019 Swamy Science

    5/22

    [37,43]. The maximum strain recovery is 8.5% after plastic

    deformation [37]. NiTi also has suitable mechanical

    properties [37] (Table 1). However, the corrosion resistance

    of NiTi is actively debated. Though the literature

    generally portrays NiTi as a corrosion resistant material

    [37,42,44], the release of nickel ions and their toxic effects

    to tissues have been reported in many cases [45,46]. Inorder to overcome this problem, the surface is passivated to

    increase the Ti oxide concentration at the surface and

    thereby reduce the nickel concentration [47,48]. This can be

    achieved by plasma-immersion ion implantation [48], nitric

    acid treatments [49], heat treatments [50], and electro-

    polishing [50]. Also, some of the materials like polyur-

    ethane [51], Ti nitride [52], and polycrystalline oxides [53]

    have been coated on NiTi stents mainly to improve the

    corrosion resistance. NiTi stents are not adequately visible

    by fluoroscopy and this is an issue [54]. Although MRI

    visualizes the stent [55], most stent deployment is

    performed under fluoroscopy. ACT-OneTM

    (Progressive

    Angioplasty Systems, USA) [56], Paragon (Progressive

    Angioplasty Systems, USA) [57], and Radius (Scimed,

    USA) [58] are some of NiTi stents used in clinical trials.

    2.6. CoCr alloy

    CoCr alloys, which conform to ASTM standards F562

    and F90, have been used in dental and orthopedic

    applications for decades [59] and recently have been used

    for making stents. These alloys have excellent radial strength

    because of their high elastic modulus (Table 1). The

    thickness of the struts is a critical issue in designing a stent

    [6062], hence, the ability to make ultra-thin struts withincreased strength using these alloys is one of their main

    attractions [63]. In addition to this, they are radiopaque [63]

    and MRI-compatible [64]. The cobalt alloy platform

    DRIVER stents (Medtronic Inc, USA) are commercially

    available in Europe. Recently, the FDA approved the L-605

    CoCr alloy Guidant Multi-Link Vision stent [65].

    2.7. Biodegradable metallic stents

    Pure Fe [66] and Mg alloys [67] are the two metals that have

    been used for making biodegradable coronary stents recently.

    2.7.1. Pure Fe

    Pure Fe (more than 99.5%) is the major component in

    degradable Fe stents [66,68]. Fe has superior radial

    strength because of its higher elastic modulus (Table 1).

    This can be helpful in making stents with thinner struts.

    Since the yield strength and tensile strength of pure Fe are

    close to each other (Table 1), theoretically, these stents may

    fracture during deployment. However, these stents were

    successfully deployed in rabbit and porcine arteries with

    balloon pressures of 3.5 and 10 atmospheres, respectively

    [66,68]. Fluoroscopy was used to view these stents (strut

    thickness varied from 100 to 120mm) [66,68]. The

    biodegradation involves the oxidation of Fe into ferrous

    and ferric ions and these ions dissolve into biological media

    [69]. Ferrous ions reduce the proliferation of smooth

    muscle cells in in-vitro conditions, and thus may inhibit

    neointimal hyperplasia [69]. Thrombogenicity and neointi-

    mal proliferation were reduced and no local toxicity was

    observed [66,68]. Endothelialization of Fe stents was also

    observed in animal models [66,68]. These studies werelimited by the small study groups [66] and slow degradation

    kinetics of Fe [66,68].

    2.7.2. Mg alloys

    Mg and its alloys have been previously used for

    biodegradable orthopedic implants [70]. However, these

    materials are novel in their application to coronary

    stents [67]. The mechanical and corrosion properties of

    pure Mg [7173] do not suit the requirements for stent

    material. Hence, Mg alloys with improved mechanical

    and corrosion properties [7173] were chosen for the

    purpose. AE21 [67] and WE43 [74] are the two Mg-based

    alloys reported in the literature for making stents. AE21

    contains aluminum (2%) and rare earth metals (1%) [67].

    WE43 contains 4% yttrium, 0.6% zirconium, and 3.4%

    rare earth metals [71]. The remaining component is Mg in

    both of these alloys. The typical mechanical properties of

    commercially available WE43 are tabulated in Table 1. It

    has poor radial strength because of its low elastic modulus.

    In order to provide proper vessel wall support, the struts

    have to be thicker and this increases the area of

    metalartery interaction. Mg alloy stents may fracture

    because of their low ductility. Also, these stents are

    radiolucent and cannot be imaged by X-rays. However,

    intravascular ultrasound and MRI techniques have used tovisualize these stents [75]. In the physiologic environment,

    Mg corrodes into soluble Mg hydroxide, Mg chloride, and

    hydrogen gas [76]. However, it is vital to investigate the

    corrosion products of Mg alloys that are actually used for

    making stents. The Lekton Magic stent (Biotronik, Switzer-

    land) is made from WE43 and implanted in porcine models

    [74]. Reduced smooth muscle cell growth and enhanced

    endothelialization were observed [74]. A Biotroniks Mg

    absorbable metal stent (AMS) was implanted in a baby and

    was well tolerated [77], but not in another baby [78]. These

    mixed results show the need for further research.

    The biodegradable metallic stent looks promising for the

    growing artery in children. However, the types of

    degradation products, size of these products, and their

    biocompatibility still need to be studied. Theoretically, the

    mechanical properties of Mg are poor for a coronary stent.

    Also, the degradation behavior of these stents is not

    controllable. Local toxicity of the degradation products of

    these stents is unlikely because Fe and Mg are present

    naturally in the human body [66,74]. However, the impacts

    of elevated local concentration of these elements are

    unknown. A detailed investigation is needed in this area

    based on large clinical trials.

    The materials for metal stents are often chosen with an

    emphasis on their engineering properties. Hence, there is

    ARTICLE IN PRESS

    G. Mani et al. / Biomaterials 28 (2007) 16891710 1693

  • 7/30/2019 Swamy Science

    6/22

    a need to emphasize the materialbiology interactions early

    on in the technology development process.

    3. Surface characteristics

    Surface characteristics of a stent material, which

    influence thrombosis and neointimal hyperplasia, includesurface energy, surface texture, surface potential, and the

    stability of the surface oxide layer [79,80]. In many

    circumstances a combination of one or more of these listed

    factors predicts the outcome. The surface properties of a

    material may depend on the surface treatment of the

    material. For example, microblasting produced a rough Ta

    surface with particle contaminants [81]. Reactive ion

    etching on the other hand, produced an even rougher

    surface on the same material but removed all the

    contaminants [81], thus demonstrating that different sur-

    face treatments can produce different surface textures and

    surface chemistries on the same material surface. This can

    provide different surface energies to the material surface.

    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 [32]. This result can be

    attributed to the technique that the Ta films were actually

    deposited (pulsed metal vacuum arc source deposition) and

    processed (annealed to 700 1C) rather than the nature of

    the material (Ta oxide) itself. In another study, the sputter

    coating of a TiTa target produced a surface that showed

    better endothelialization because of the changes in the

    microstructure of the natural Ti-oxide film produced [34].

    To improve the corrosion resistance, Ta is coated on a

    316L SS substrate by physical vapor deposition sputtering[28]. However, when the material is plastically deformed,

    cracks appear in the coating [28]. Though it was claimed

    that the crack surfaces were repassivated by treating with

    the physiological saline, this kind of acute change (crack-

    ing) in surface morphology might pose a serious threat

    when the material is exposed to in-vivo conditions. The

    nature of the coating may be biocompatible, but, if the

    coating loses its integrity during the stent placement and

    expansion, it can cause adverse effects.

    3.1. Surface energy

    The surface energy of a material as well as its surface

    chemistry affects its wettability. The thrombogenicity of a

    material surface increases with increasing surface energy

    [82,83]. The thrombogenic potentials of PET and PTFE

    were compared and the thrombogenicity was significantly

    higher for PET (high surface energy) than PTFE (low

    surface energy) [84]. This result has initiated the use of

    coating metal surfaces that usually have higher surface

    energy with polymeric materials having low surface energy.

    A polyurethane coating on Ta and SS reduced the surface

    energy which resulted in the significant reduction of

    thrombosis [85,86]. One limitation of these thromboresis-

    tant coatings (hydrophobic polymers) is that they not only

    prevent the adherence of platelets but also endothelial cells

    [87]. This problem can be prevented by coating the polymer

    surfaces with fibronectin, which enhances the endothelia-

    lization process [87,88]. A combination of human plasma

    proteins and heparin coating on copolymers, which are

    derived from hydrophilic 1-vinylpyrrolidinone and hydro-

    phobic n-butyl-methacrylate, provided better endotheliali-zation and thromboresistance to the material, respectively

    [89]. Seeger et al. [90] coated hydrophilic polymers such as

    N-vinylpyrrolidone and potassium sulfopropyl acrylate on

    SS to reduce the accumulation of platelets. These coatings

    indeed smoothened the irregularities of underlying SS. In

    spite of their similar hydrophilicity, there was a significant

    difference in the platelet accumulation between the

    polymers. This can be attributed to different surface

    charges of the polymers [90]. Hence, when polymers are

    coated on a metal stent, the surface energy is not the only

    parameter that is altered but also the other parameters like

    surface texture and surface charge.

    3.2. Surface texture

    Thrombogenicity is usually higher for rougher surfaces

    [9193], thus polishing is essential for stent materials. Acid

    pickling followed by electrochemical polishing has been

    used to remove the slag (formed on the stent during its

    laser cut) and to polish the stents, respectively [94]. It has

    been shown that polishing of coronary stents resulted in

    decreased thrombogenicity as well as neointimal hyperpla-

    sia in different animal models [95,96]. Also, the coating

    techniques used for surface deposition can directly

    influence the surface texture. Hehrlein et al. [97] investi-gated the effect of two surface deposition methods,

    galvanization and ion implantation, on the biocompat-

    ibility of endovascular stents. Both thrombogenicity and

    neointimal hyperplasia were higher for the stents that are

    coated by galvanization because of the pores and cracks

    created during expansion [97]. Close control over the

    surface texture of stents is relatively difficult due to

    morphological changes during its expansion. Hence, when

    a stent is polished or coated, sufficient care should be taken

    to evaluate the effects of expansion. Recently, Sprague

    et al. [98] observed that the grooved surfaces double the

    migration rate of endothelial cells over polished and smooth

    controls. Larger grooves (on the order of 22mm) resulted in

    greater migration rates and faster endothelialization times.

    Thus, clearly the surface roughness of the stent is an

    important parameter in their clinical success. This needs to

    be taken into consideration as new coatings are being

    developed, especially for drug elution, as coatings may have

    different surface textures compared to the bare metal.

    3.3. Surface potential

    The net electrical charge on a material surface is also

    critical to the success of a stent [26,91]. Zitter et al. [26]

    investigated the current densities of different metals in

    ARTICLE IN PRESS

    G. Mani et al. / Biomaterials 28 (2007) 168917101694

  • 7/30/2019 Swamy Science

    7/22

    in-vitro conditions and ranked the current densities in the

    following order: Au4316L SS4CoCr4Ti6Al4V4

    Ti4Nb4Ta. The reasons for the good and poor

    biocompatibilities of Ta and Au, respectively were dis-

    cussed with respect to the decreasing order of current

    densities [26]. Most metals are electropositive while blood

    elements tend to be electronegative, which accentuates thethrombogenicity problem [91]. Ta has a net negative

    electrical charge and thus has a theoretical advantage over

    other metals [26,91]. However, experimental results have

    been contradictory; Scott et al. [99] compared the thrombo-

    genicity of 316L SS and Ta stents of identical design in

    baboon and porcine models. The platelet and fibrin

    deposition on these stents were determined and it was

    concluded that there was no significant difference in

    thrombogenicity between 316L SS and Ta stents [99]. In

    another study, the least electropositive metals like copper

    induced more neointimal hyperplasia than the most electro-

    positive metals like gold and platinum [97]. This illustrates

    that thrombogenicity and biocompatibility depend on multi-

    ple factors and that surface charge is just one such parameter.

    3.4. Stability of surface oxide layer

    The stability of surface oxide layer directly influences the

    biocompatibility of a material as the surface layer acts as a

    barrier to the release of ions from the bulk materials

    underneath the surface. For example, the percentage of

    nickel in NiTi and SS is 50% and 12%, respectively

    [16,100]. Atomic adsorption spectrophotometry analyses

    revealed a significant release of nickel and chromium metal

    ions from non-coated SS stents over 96 h in human plasma[101]. The release of nickel ions from NiTi has been

    reported in few cases [45,46]. The released metal ions

    induced platelet and leukocyte activation [101,102], which

    resulted in thrombogenicity of the stents. Also, the

    endothelial cell damage caused by the release of very low

    concentration metal ions may be considered as a poten-

    tially toxic effect [102]. The stability of oxide layer is also

    key to several of the surface characteristics discussed

    above. It influences surface energy by providing hydro-

    philicity to a material surface and surface potential by

    preventing the release of electrons. Since the stability of the

    natural surface oxide layer in 316L SS and NiTi is not

    very high, the possibility of metal ions being released is

    greater. Coating of stents is the most common approach to

    prevent this effect.

    4. Rationale for coatings

    Since the basic mechanisms underlying the interaction

    between a metal and tissue/blood are still not completely

    understood, the biocompatibility and the hemocompat-

    ibility of metallic stents still remains an issue. Thrombosis

    and neointimal hyperplasia were commonly reported

    among bare metallic stents [103107]. Coating the metallic

    surface with other materials to alter its surface character-

    istics without interfering with the bulk properties of the

    metal stent has been one rational approach to address this

    issue [108,109]. The coating of a stent can improve the

    surface properties significantly; surface energy can be

    reduced, surface texture can be smoothened, surface

    potential can be neutralized and the stability of the surface

    oxide layer can be enhanced. These modifications coulddirectly influence thrombosis and neointimal proliferation,

    which both can reduce restenosis.

    4.1. Types of coatings

    Galvanization [97], sputtering followed by ion bombard-

    ment [97], pulsed biased arc ion plating [110], dipping

    [111,112], spraying [113,114], and plasma-based deposi-

    tions [115,116] are some of the techniques that have been

    commonly used for coating stents. Initially, the coatings

    were used to increase the biocompatibility of stents, but

    later this technique became a platform for the controlleddelivery of drug to inhibit intimal hyperplasia. The coating

    materials for stents can be broadly classified into four

    types: inorganic materials, polymers, porous metals, and

    endothelial cells. Inorganic materials and endothelial cells

    are exclusively used as coating materials, while the

    polymers are used as both as a coating material as well

    as the sole stent material. Porous metal coatings can be of

    the same metal as the stent or an alternative metal.

    4.1.1. Inorganic coatings

    There are many inorganic coating materials which are

    potentially suitable for the treatment of medical implantsurfaces. Gold, silicon-carbide, iridium oxide, and dia-

    mond-like carbon are some of the commonly used

    inorganic-coating materials on stents.

    4.1.1.1. Gold. At one juncture gold coating was a

    preferred coating on SS stents to enhance fluoroscopic

    visibility for reduced strut thicknesses of 5080 mm [117].

    Since gold has six times the radiopacity of steel, a 5 mm

    coating on each side doubles the radiopacity of an 80 mm-

    thick steel stent [117]. Edelman et al. [118] investigated the

    vascular response in porcine coronary arteries by compar-

    ing the standard gold coating with thermally processed

    gold coating. The reduction in neointimal hyperplasia and

    inflammation for the thermally processed coating over

    standard coating was mainly attributed to smoothened

    gold surface and removal of embedded impurities in the

    gold coating. This study indicated that surface properties

    and material purity may play a significant role in

    tissuematerial interactions. However, the human clinical

    trials on gold-coated stents were not satisfactory. Dahl et

    al. [119] reported that the neointimal proliferation was

    more in patients who received gold-coated stents. Danzi

    et al. [120] reported that the morphology of the restenosis

    was proliferative in 83% of the cases and the remaining

    17% were occluded.

    ARTICLE IN PRESS

    G. Mani et al. / Biomaterials 28 (2007) 16891710 1695

  • 7/30/2019 Swamy Science

    8/22

    4.1.1.2. Iridium oxide. Iridium oxide, generally accepted

    as a highly biocompatible inert ceramic material has been

    used for stent coatings [121,122]. It was found that

    hydrogen peroxide is produced at a metal (cobalt, zinc,

    nickel, copper, silver, chromium, and some of their alloys)

    surface when it is corroded [123]. Hydrogen peroxide, a

    strong oxidizing agent, can be harmful to the artery andcan cause inflammatory reactions. It is claimed that a metal

    coated with iridium oxide promotes the immediate

    conversion of hydrogen peroxide into water and oxygen,

    so it was expected that it might reduce inflammatory

    reactions and promote endothelialization [124]. Initial

    studies in the porcine model showed that this coating can

    reduce the neointimal thickness from 118 mm on a bare SS

    stent to 55mm on an iridium-coated stent [121]. The Lunar

    stent (Inflow Dynamics, Germany) is a 316L SS with a thin

    inner layer of gold for increasing visibility and an outer

    layer of iridium oxide for improving biocompatibility [124].

    A clinical trial evaluated the immediate and long-term

    outcome of these stents and reported that the overall

    angiographic restenosis rate was 13.8% [124]. It was

    reported that the iridium oxide promoted fast endothelia-

    lization because of its capability to prevent the production

    of free oxygen radicals, which can affect the adhesion and

    proliferation of endothelial cells [124]. The surface of SS

    produced little hydrogen peroxide unlike its alloying metals

    (nickel and chromium) [123]. Hence, a detailed investiga-

    tion is needed to evaluate the amount of H2O2 actually

    produced on a 316L SS surface and the role of iridium

    oxide in the conversion of H2O2 into H2O and O2.

    4.1.1.3. Silicon-carbide (SiC). Amorphous hydrogenatedSiC, a semiconductor, has been well known for its

    antithrombogenic properties [125]. The reduction in the

    deposition of platelets, leukocytes, and monocytes over a

    stent made of SiC offers a promising coating material for

    reducing restenosis [126,127]. Although in-vitro studies

    [128] provided encouraging results, the outcomes from the

    various human trials produced contradictory results. For

    instance, the presence of endothelialization was reported in

    a 6-month clinical follow-up of the Tenax coronary stent

    [129]. In contrast, results showing greater neointimal

    hyperplasia were observed in a 6-month clinical follow-up

    for SiC-coated stents in another study [130]. A clinical trial

    compared the SiC-coated stents (Biotronik, Germany) with

    316L NIR stents (Boston Scientific, USA) and concluded

    that both stents had a low rate of major adverse coronary

    events at 81712 weeks of follow-up, with no definite

    superiority [131]. These mixed results indicate the need for

    further research in this area. Bickel et al. [132] compared

    the platelet adhesion on different coating materials and

    found that SiC coating has a significant effect in reducing

    the number of adhesion platelets than the uncoated stents

    but the effect is inferior to heparin and/or carbon coating.

    4.1.1.4. Carbon. Coating of stents with diamond-like

    carbon, a chemically inert hydrocarbon, have shown

    improved biocompatibility [133,134]. The results of a 6-

    month follow-up of Carbostent (Sorin Biomedica, Italy)

    showed that the carbon coating can significantly reduce

    stent thrombosis and restenosis in relatively high-risk

    patients [135]. The observed angiographic restenosis rate

    was 11% with no indication of subacute thrombosis.

    However, the results were not consistent with other studies.More recently, carbon coating has been considered

    inactive because they showed no improvements in

    angiographic restenosis [136]. Another clinical study

    compared uncoated MAC (AMG Raesfeld-Erle, Germany)

    stents with carbon-coated MAC stents [137] and showed

    that the carbon coating did not influence the inflammatory

    response. Other studies showed similar rates of binary

    restenosis, 31.8% for carbon-coated stents and 35.9% for

    bare metallic stents [138]. Thus carbon-coating stents have

    not yielded a significant clinical improvement.

    A variety of these coating materials have been claimed to

    have improved biocompatibility, hemocompatibility, and

    antithrombogenicity properties for stents. However, these

    claims are usually not based on thorough comparative

    experiments and more definitive work remains to be done.

    4.1.2. Endothelial cells

    Endothelial cell damage and exposure of subendothelial

    matrix at the site of arterial injury is a basis for both

    thrombus and neointima formation [7,9]. This clearly

    illustrates the importance of re-endothelialization, an

    approach in which endothelial cells are placed on stents

    before being implanted, and the cells are expected to

    proliferate, differentiate and release growth factors, which

    in turn inhibit thrombosis and neointimal hyperplasia[139]. Van der Giessen et al. [140] were the first ones to seed

    endothelial cells on stents and study their in-vitro behavior.

    Several attempts have been made to seed endothelial cells

    on stents, but most of them have been limited by the quick

    loss of seeded cells, endothelial cell damage upon stent

    expansion, and the inability to maintain cell adherence to

    the artery wall during the blood flow [9,139].

    4.1.3. Porous materials

    One method that has been attempted to promote rapid

    endothelialization is to create micropores on the walls of

    vascular grafts [141,142]. Nakayama et al. [143] implemen-

    ted this technique by covering stents with a segmented

    polyurethane film with micropores of diameter 30mm

    prepared by laser ablation technique. Increased pore

    density resulted in better endothelialization and a thinner

    neointimal layer [143]. However, this technique was limited

    by two factors [144]: (i) the non-flat luminal surface design

    leads to thrombus formation, and (ii) when the edges of the

    polyurethane were overlapped by gluing, the pore density

    at these spots becomes zero which results in high

    neointimal hyperplasia. Later this technique was modified

    by dip coating the stent in polyurethane twice to have a flat

    luminal surface and a microporous outer surface [145].

    Also, heparin and tacrolimus were immobilized on the

    ARTICLE IN PRESS

    G. Mani et al. / Biomaterials 28 (2007) 168917101696

  • 7/30/2019 Swamy Science

    9/22

    luminal and outer surface of the PU matrix, respectively by

    using photoreactive gelatin followed by UV irradiation

    [145]. However, after these techniques were reported for

    immobilizing the drugs, there have been no published

    reports available on drug release profiles from these

    systems. Wieneke et al. [146] created a nanoporous

    aluminum oxide coating on a SS stent for loadingtacrolimus. Though the surface modification by ceramic

    coating did not show significant effect in reducing the

    neointima in a rabbit model, the release of tacrolimus

    reduced neointimal thickness in a dose dependent (52%

    and 56% reduction for 60 and 120 mg, respectively) manner

    [146]. Recently, it was reported that this kind of ceramic

    coating may liberate particle debris which in turn affects

    the antiproliferative effect of tacrolimus and results in a

    significant increase in the neointima as compared to the

    control uncoated stents [147].

    These studies further stress the importance of coating

    integrity. The coating is applied on a stent surface in its

    crimped state; i.e. the surface area of the stent is minimal.

    When the stent is expanded, the surface area increases and

    can result in fissures, cracks and pores on the coating. The

    damaged coating can also release particulate debris. Such

    changes in surface morphology and localized particle

    delivery can increase the chances of restenosis. Also it

    was reported that 80% of loaded drug was released from

    nanoporous coatings within 100 h and the remaining 20%

    was not released [147]. This type of release profile is

    indicative of a burst effect. Since it is preferable to release

    the drug over at least a 30-day period, open porous coating

    may either require smaller pores to trap the drugs longer or

    a second coating, which retards quick drug release.

    5. Polymers

    Polymers used for coating stents can be broadly

    classified into biostable (non-biodegrable) polymers, bio-

    degradable polymers, copolymers, and biological poly-

    mers. Several polymers with previous medical or dental

    applications have been used for coating stents or for

    making the entire stent. Although a wide range of polymers

    have been used to coat the stent, only a few, like

    polyethylene terepthalate (PET), poly-L-lactic acid

    (PLLA), and poly-L-glycolic acid (PLGA) have been tested

    as a lone stent material.

    5.1. Biostable polymers

    The principle of biostable polymeric stents is very similar

    to metallic stents: the stent should have sufficient mechan-

    ical properties for providing stable support in maintaining

    the lumen gain [148]. Besides that, it should be biocompa-

    tible, and should not initiate thrombus formation and

    inflammatory reactions. The elastic modulus of biomedical

    polymers usually lies in the range of 15 GPa [149], which

    raises concerns whether they posses sufficient mechanical

    properties for use as stents. However, Van der Giessen

    et al. [150] have shown that the radial pressure exerted by

    PET braided mesh stents is the same as that of SS stents.

    PET has been investigated for making stents because of its

    good mechanical properties [149] and its reputation as a

    successful material for cardiovascular grafts [149,151].

    Murphy et al. [152] deployed PET stents in the coronary

    arteries of porcine model. Though this study demonstratedthe possibility of percutaneous deployment of polymeric

    stents in coronary arteries, the use of PET was associated

    with a chronic foreign body inflammatory reaction and an

    intense proliferative neointimal response that resulted in

    the complete occlusion of the vessel. In another study in a

    porcine model, the PET-stented vessels were endothelia-

    lized and the neointimal thickening was 44113 mm 4 weeks

    post implantation [153]. Though it was stated that the

    extent of neointimal proliferation was limited when

    compared to the responses induced by bare metallic stents,

    the foreign body reaction was significant and highlighted

    the inflammatory reactions elicited by PET. Lack of

    radiopacity is also a concern for PET stents [154].

    5.2. Biodegradable polymers

    Due to the above-mentioned limitations of bare metallic

    stents and biostable polymer stents, biodegradable stents

    have been considered as an option. Biodegradable stents

    have the theoretical advantage of no longer being present

    as a foreign material in arteries once they have scaffolded

    the vessel for a relevant period of time [148,155]. The other

    significant advantage is that drugs can be released in a

    controlled manner [148,155]. Stack et al. [156] developed

    the first biodegradable stent made of PLLA and deployed itin canine model. The initial results showed the occurrence

    of limited thrombosis and minimal neointimal proliferation

    in the short term and also at 18 months. The mechanical

    behavior of these stents was also investigated [157].

    However, the implantation of biodegradable polymers in

    a porcine model showed extensive inflammation and

    neointimal proliferation [158]. To investigate the issues in

    detail, Van der Giessen et al. [159] implanted 3 different

    biodegradable polymers (polyorthoester, polycaprolac-

    tone, and polyethylene oxide/polybutylene terepthalate

    (PEO/PBTP)) and 3 biostable polymers (polyurethane,

    silicone, and PET) in porcine arteries. After 30 days of

    implantation, histological examination strongly confirmed

    the presence of neointimal thickening at the polymer-

    coated side of each stent [159]. Extensive fibro-muscular

    proliferation, multinucleated giant cell formation, mono-

    nuclear and eosinophilic smooth muscle cell proliferation

    were seen adjacent to PEO/PBTP and PET samples [159].

    Lincoff et al. [160] showed that high molecular weight

    PLLA was well tolerated in a porcine model while the low

    molecular weight PLLA was not. This study showed that

    molecular weight of polymer also has an impact on

    neointimal hyperplasia. In spite of the controversies of

    using biodegradable stents, PLLA stents were implanted in

    a small clinical trial and the results were encouraging [161].

    ARTICLE IN PRESS

    G. Mani et al. / Biomaterials 28 (2007) 16891710 1697

  • 7/30/2019 Swamy Science

    10/22

    Though it has been speculated that biodegradable

    polymers induce inflammatory reactions because of an

    immune response to degradation products and non-reacted

    monomer compounds [158], the basic mechanism is still not

    completely understood.

    5.3. Copolymers

    Several copolymers which include polyhydroxy

    butyrate/valerate [159], PEO/PBTP [159], methyl metha-

    crylate/2-hydroxy ethyl methacrylate [162], ethylene-vinyl

    acetate [163], laurylmethacrylate/methacryloylphosphoryl-

    choline [164], PLGA [159,165] and polyurethanes (PU)

    have been evaluated either as a coating material or as a

    base stent material. However, PLGA and PU are the most

    investigated copolymers for coronary stents. PLGA, the

    copolymer of polylactic acid and polyglycolic acid, has

    been widely used in bioresorbable sutures, drug delivery

    devices and orthopedic implants [149]. The degradation

    behavior of PLGA is crucial for its use in controlled drug

    delivery stent systems. For example, it has been observed

    that the rate of heparin release was slowest for PLLA,

    followed by PLGA (80/20) and finally PLGA (53/47) for

    coronary stent applications [166]. Thus the rate of drug

    release from PLGA stents would likely depend on the

    copolymer ratio. Recently, Venkatraman et al. [167]

    imparted self-expanding capability to biodegradable PLLA

    stents (at 37 1C) by adding PLGA (53/47). This effect was

    induced in the bilayered stents by fabricating them at 37 1C.

    Interestingly, PLGA stents have been investigated more in

    urological applications than coronary applications

    [168,169]. PU have been extensively used for medicaldevice applications because of their excellent biocompat-

    ibility [83]. These polymers have been used as a coating

    material on stents to improve the antithrombogenic

    properties of Ta [85], corrosion resistance of NiTi [51],

    and biocompatibility of SS [170]. It was also reported that

    PU coatings can improve endothelialization [171]. PU has

    also been investigated in drug delivery systems [172174].

    Lambert et al. [172] successfully studied the drug release

    kinetics and distribution of the model drug Forskolin

    delivered from the PU coated metallic stents. However,

    some studies have shown that PU coating can be

    accompanied by extensive inflammatory reactions [175].

    Thus while some studies advocated the use of PU for

    stents, others show contradictory evidence. Hence it is

    difficult to categorize PU as a good/poor stent material.

    Although PU has successfully been used in many

    cardiovascular devices (pacemaker lead wires, vascular

    grafts, artificial heart pumps, and inner surface coatings of

    artificial heart [149,176,177]), it does not necessarily mean

    that the coating may be beneficial for stents.

    5.4. Biological polymers

    Natural polymers are derived from natural resources and

    can be broadly classified into those of plant and animal

    origin. Phosphorylcholine (PC), hyaluronic acid (HA), and

    fibrin are some of the biological polymers that were

    extensively explored for coating stents.

    5.4.1. PC

    Phospholipid PC, an essential part of the red blood cell

    membrane, is structurally composed of both hydrophilic

    and hydrophobic components. This has been coated on

    metallic stents mainly to prevent the adhesion of coagula-

    tion-inducing cells [178]. An initial study of PC as a stent-

    coating material in porcine models showed its excellent

    bio- and hemo-compatibility [179]. Thereafter, extensive

    literature is available on the hemocompatibility and tissue

    compatibility of PC-coated stents [179,180]. The BiodivY-

    sio stent, coated with PC, was evaluated in a human clinical

    study and the results showed that the restenosis decreased

    from 8977% to 5.676% [181]. Many other human

    clinical studies confirmed the antithrombogenic properties

    and decreased restenosis rates of PC-coated stents[182,183]. Also, the PC coating is stable up to 6 months

    of implantation [184186]. These characteristics together

    with its ability to deliver drugs make this material an

    attractive choice of coating for DES [187]. A human

    clinical trial studied the PC coating-based elution of an

    antiproliferative agent, ABT-578, using Endeavor stents in

    humans [188]. The restenosis rate of endeavor stents

    (13.3%) was almost three times less than the bare cobalt

    alloy driver stent (34.2%) [188]. A non-randomized trial

    investigated the PC coating-based elution of dexametha-

    sone and reported the binary restenosis rate of 13.3% for

    60 patients at 6 months [189]. Recently, a human clinical

    trial successfully evaluated the release of angiopeptin from

    a PC-coated stent and showed encouraging results [190].

    5.4.2. HA

    HA, a linear polysaccharide non-sulfated glycosamino-

    glycan present in various tissues of the body, has been

    found to improve the thrombo-resistance of stents.

    Verheye et al. [191] reported a significant reduction of

    platelet deposition on HA-coated SS stents in a baboon

    model. In order to extend the antiproliferative and

    antithrombogenic properties of biodegradable HA, it can

    be made insoluble by self-cross-linking with N-(3-dimethy-

    laminopropyl)-N0-ethyl carbodiimide [192]. Reduced in-

    flammatory responses were found for periods up to a

    month when compared with uncoated SS stents in

    undiseased pig coronary arteries [192]. In another study,

    HA was covalently attached to SS [193]. Epoxy silane was

    covalently attached to SS and then the epoxy group was

    converted to aldehyde group to react it with HA. The

    approaches which involve chemical modification of the

    coating polymer need to be thoroughly characterized

    before implantation. Even traces of the chemical used for

    chemical modification can be non-biocompatible and

    eventually leads to erroneous conclusions about the coat-

    ing material. Though the available literature on HA

    ARTICLE IN PRESS

    G. Mani et al. / Biomaterials 28 (2007) 168917101698

  • 7/30/2019 Swamy Science

    11/22

    coatings is meager, coating the stents with these biopoly-

    mers seems promising.

    5.4.3. Fibrin

    Fibrin is an insoluble protein produced during the

    coagulation of blood. This biopolymer is well known for

    its biocompatible, biodegradable, and viscoelastic proper-ties [194,195]. Holmes et al. [196] demonstrated the

    potential role of exogenous fibrin as a better coating

    material in a porcine model. They compared a circumfer-

    ential fibrin sleeve-coated coil wire stent with a PET stent

    and a PU-coated stent. Markedly less vessel occlusion and

    foreign body reaction was observed with exogenous fibrin

    than with PU and PET stents. Another major advantage of

    the fibrin-film stent is that it provides complete endolum-

    inal paving by covering 100% of the arterial surface,

    compared with the partial coverage achieved with bare

    metallic stents [197]. This strategy particularly facilitates

    site-specific therapies, such as delivering the drugs to the

    entire lesion surface [197]. It may also lead to rapid

    endothelization [198].

    The application of all these biological polymers as stent

    coatings appears very promising. Since such polymers

    facilitate re-endothelialization and show negligible inflam-

    matory reactions, human clinical trials for stents with

    coatings of biological polymers is the logical next step.

    6. Rationale for DES

    Endothelial and smooth muscle cell damage, unavoid-

    able in PTCA and stent placement, is a cause of restenosis

    [9]. The optimization of the architecture and mechanicalcharacteristics of stents has lead to a decrease in restenosis

    but using drug delivery platforms remains a promising way

    to further reduce restenosis. The main reason for the failure

    of systemic pharmacological therapy is the inability to

    deliver an adequate drug dose at the site of injury [199].

    Earlier approaches for local drug delivery by using

    catheter-mounted balloons and needles were not successful

    due to rapid washout of the drugs by the blood stream

    [199,200]. Currently, the treatment available for preventing

    restenosis is the implantation of DES. Heparin was the first

    therapeutic agent attached directly to a stent. The concept

    of delivering medications at the injury site has evolved

    from heparin-coated stents to stents with drugs that inhibit

    neointimal hyperplasia. For preventing neointimal hyper-

    plasia, an appropriate drug concentration has to be

    delivered for at least 30 days during which the biology of

    restenosis is known to occur.

    6.1. Techniques for drug-loading and release kinetics

    The techniques for loading drugs on a stent can be

    categorized into three major types: (i) attaching the drug

    directly onto the metal surface; (ii) loading the drug into

    the pores of porous metal stents; and (iii) incorporating the

    drug in a polymer that is then used as a stent-coating

    material. The drug release depends on the way the drug is

    loaded on the stent. If the drug is physically adsorbed on

    the metal surface or in the porous surface, it can be released

    by simple diffusion. Here, the porous surface offers the

    possibility of incorporating more drugs than the metal

    surface because of the greater surface area. The amount of

    drug release can also be controlled by the size and densityof the pores. If the drugs are trapped inside non-

    biodegradable polymers (techniques used in CYPHERTM

    and TAXUSTM

    Express2TM

    ), they are released by diffu-

    sion. In this case, the amount of drug released depends on

    the thickness of the outer coating, as it modulates the

    amount of drug that can be released per unit time. When

    the drugs are chemically attached to the surface, the drug

    release depends on the rate at which the chemical bonds are

    cleaved. The rate of chemical bond cleavage depends on the

    orientation of drug molecules, which determines the

    triggers access to the bond. Drug delivery through

    biodegradation is the most common phenomenon and it

    has been extensively reviewed in the literature for

    orthopedic [201,202], ocular [203,204], neuro [205,206],

    and cardiovascular applications [155,207]. The same

    concept applies in case of DES as the drug-incorporated

    matrix is coated on the metal surface and the rate of drug

    release depends on the rate at which the matrix is degraded.

    6.2. DES

    The three drugs that have been investigated in depth for

    treating restenosis are heparin, sirolimus, and paclitaxel

    (Fig. 1). Heparin has been effective in reducing both

    thrombosis and neointimal proliferation while sirolimusand paclitaxel were mainly used for their anti-proliferative

    effects in blocking neointimal hyperplasia.

    6.2.1. Heparin-coated stents

    Heparin, a heterogeneous group of unbranched, acidic

    glycosaminoglycans, has been widely used for modifying

    the surfaces of vascular implants because of its antic-

    oagulant properties [208]. Heparin activity depends on the

    interaction between its active sites (carbohydrate se-

    quences) and the circulating antithrombin III. The

    antithrombin which binds to the active sites catalyzes the

    inhibition of thrombin and the resultant inactive anti-

    thrombin/thrombin complex is released into the blood

    stream [209].

    There are various ways in which heparin can bind to a

    stent surface and these include physical adsorption, ionic

    bonding, copolymerization, and polymer encapsulation.

    Physical adsorption has been attained by coating the stent

    with a solution of water-insoluble benzalkonium chloride

    complex [210]. For ionic bonding, the material surface was

    cationically charged through quarternization (treatment

    with tridodecylmethylammonium chloride ammonium salt

    or ethyl bromide) treatment and then the anionic heparin

    molecules are ionically bonded on to the cationic surfaces

    [211,212]. The stability of both physically adsorbed and

    ARTICLE IN PRESS

    G. Mani et al. / Biomaterials 28 (2007) 16891710 1699

  • 7/30/2019 Swamy Science

    12/22

    ionically bound heparin is low as they are easily removed

    from the surface when exposed to plasma [213]. The

    stability issue raised a question about the long-term anti-

    coagulation therapy. To improve the stability, heparin was

    copolymerized with a variety of polymers like poly(methyl-

    methacrylate) [214], poly(vinyl alcohol) [215], and PU

    [216]. Though the copolymerization techniques provided amore stable binding of heparin to the surfaces compared to

    physical adsorption and ionic binding, these techniques

    tend to alter the chemical sequences of heparin, which are

    essential for its therapeutic effect. Larm et al. [213] created

    aldehyde groups in heparin through its reaction with

    nitrous acid. Then, the aldehyde groups were covalently

    bonded to the amine terminated stent surface. This end-

    point attachment technique has the unique advantage of

    having secured the active carbohydrate sequences, which

    are essential for binding antithrombin.

    Heparin delivery has been achieved using biodegradable

    PLGA microspheres [217]. In order to control the heparin

    delivery from the biodegradable polymers (PLLA,

    PLLGA, PLGA), polyethylene glycol, which is a plastici-

    zer, was added to the polymerheparin films [166]. The

    effect of plasticizers on the release profiles of heparin was

    found to be dependent on the copolymer ratios of PLA and

    PGA. Though several techniques have been tried to

    immobilize and/or deliver heparin at the target site, each

    technique has its own advantages and disadvantages and

    none has proved to be optimum.

    Bonan et al. [218] were the first to use heparin coated

    zigzag stents in canine coronary arteries. Several other

    animal studies [219,220] and clinical trials [221225]

    investigated the efficacy of heparin-coated stents and

    strongly confirmed the absence of thrombosis. A reduction

    in neointimal formation was also reported in few animal

    studies [163,226].

    6.2.2. Sirolimus-eluting stents (SES)

    Sirolimus, an immunosuppressive agent, binds to an

    intracellular receptor protein and ultimately induces cell-cycle arrest [227]. It inhibits vascular smooth cell migra-

    tion, proliferation and growth [227,228].

    A variety of biodegradable and non-biodegradable

    polymers have been used as coatings for SES [229]. The

    FDA approved sirolimus-coated BXTM

    Velocity balloon

    expandable stent (CYPHERTM

    ) [230] is made from

    electropolished laser-cut 316L SS. The stent is coated with

    a three layers of polymer coating: parylene C, an inert,

    hydrophobic and biocompatible polymer is initially coated

    on the metallic stent. Then, a mixture of polyethylene-co-

    vinyl acetate (PEVA) and poly n-butyl methacrylate

    (PBMA) in a ratio of 67:33 is mixed with sirolimus and

    coated on the parylene C coating. Finally, a mixture of

    PEVA and PBMA is then applied as the third layer without

    sirolimus. The main purpose of this final coating is to

    prevent the fast leaching of drugs from the polymer coating

    during the initial period post-implantation. Recently Chen

    et al. [113] spray-coated collagen and sirolimus layer-by-

    layer alternatively on to a SS stent. The collagen matrices

    were used for releasing the sirolimus in a controlled fashion

    without having any burst effect. As the collagen is already

    known for its better blood compatibility, this kind of

    biocompatible coatings without the use of polymers seems

    promising. Several clinical studies have addressed the

    usefulness of SES [231235].

    ARTICLE IN PRESS

    Fig. 1. Chemical structure of (A) heparin, (B) sirolimus and (C) paclitaxel.

    G. Mani et al. / Biomaterials 28 (2007) 168917101700

  • 7/30/2019 Swamy Science

    13/22

    6.2.3. Paclitaxel-eluting stents (PES)

    Paclitaxel is a drug used in the treatment of cancer. This

    drug binds to the tubulin protein of microtubules, which

    are the components of cells that provide structural frame-

    work and enable cells to divide and grow. The abnormality,

    paclitaxel/microtubule complex, in vascular smooth muscle

    cells inhibits cellular replication and ultimately causescellular death [236].

    The coating of paclitaxel on stents can be broadly

    classified into two types: polymer-based and non-polymer-

    based coatings. Heldman et al. [111] coated paclitaxel

    directly on the stent surface by dipping the stent in the

    ethanolic solution of paclitaxel followed by evaporating the

    alcohol. This technique is particularly advantageous in that

    there are no concerns over inflammatory reactions induced

    by the polymer-based drug delivery systems. Also, the

    tissue can be in direct touch with the drug coating.

    However, the disadvantage of dip coating is that a

    significant amount of drug is lost during the stent

    placement and expansion similar to the problems that

    were encountered during catheter-based drug elution

    attempts [111]. This shows the significance of having a

    carrier, which can hold the drugs and release them in a

    controlled fashion. In one study, three different doses (0.2,

    15, and 187mg/stent) of paclitaxel were coated directly on

    to the stent [111]. The higher dose showed significant

    reduction in neointimal hyperplasia compared to the lower

    doses in a pig model after 28 days. The amount of drug that

    is loaded on the stent is a very crucial factor. In the case of

    dip coating, the amount of drug that can be coated depends

    on the surface area of the stent. The surface area is

    determined by the basic design of the stent and thickness ofthe struts. A low profile design with thinner struts is usually

    preferred for successful implantation. However this sig-

    nificantly limits the amount of drug that can be coated on

    the stent. A multicenter study evaluated the ability of

    Supra-G (Cook Inc), a 316L SS stent coated with a

    polymer-free formulation of paclitaxel, to inhibit restenosis

    [237,238]. This study showed that a paclitaxel-coated stent

    could significantly reduce restenosis at 6 months after

    intervention. The dose of 3.1 mg/mm2 was more effective

    than the dose of 1.3 mg/mm2. Intravascular-ultrasound

    evaluation demonstrated a dose-dependent reduction in the

    volume of intimal hyperplasia. In another clinical study,

    researchers evaluated the V-Flex Plus (Cook Inc), a 316L

    SS stent, coated with increasing doses of paclitaxel (0.2,

    0.7, 1.4, and 2.7 mg/mm2 stent surface area) [238]. The drug

    was applied directly to the albuminal surface of the stent

    with polymer-free formulation. The authors concluded that

    the in-stent restenosis was significantly reduced for a

    paclitaxel dose density of 2.7mg/mm2 without short- or

    medium-term side effects. These studies clearly demon-

    strate the importance of the amount of drug loading in

    preventing restenosis. Polymer coatings, in general, can

    carry higher loads of drug compared to direct drug

    adsorption on the metal surface. A vascular compatible

    poly(styrene-b-isobutylene-b-styrene) triblock copolymer

    (SIBS) is used as a paclitaxel carrier in the FDA approved

    TAXUSTM

    Express2TM

    paclitaxel-eluting coronary stent

    [239]. Unlike CYPHERTM

    , this stent does not have

    additional outer coating (polymer coating without drug)

    to prevent the burst effect. This may be one of the reasons

    why the drug elution profiles for TAXUSTM

    stents are for

    30-day periods while compared to the 60-day periods forCYPHER

    TMstents. Ranade et al. [240] found that the

    paclitaxel solubility in SIBS matrix is extremely low and it

    exists as nanoparticles in the polymer matrix. There was no

    observable change in the surface morphology of the

    polymer matrix after the incorporation of paclitaxel.

    AFM images confirmed the morphology changes only

    during the drug elution period and supported a burst effect

    of greater than 8.8% of the total amount of the loaded

    drug. This confirmed that the paclitaxel release is directly

    dependent on the paclitaxel loading. In order to increase

    the miscibility of paclitaxel in SIBS, Sipos et al. [241]

    chemically modified the styrenic portion of the SIBS

    polymer system and thereby modulating the drug release

    profile. It was reported that the modulation is due to the

    improved hydrophilicity and polarity of the polymer

    systems [241]. The efficacy of NIRxTM

    (Boston Scientific)

    PES were evaluated by several clinical trials [242244].

    A recent meta-analysis of 6 randomized trials comparing

    PES with SES was performed [245]. Superiority of SES was

    observed with angiographic restenosis rates of 9.3% when

    compared to 13.1% for PES. The difference between the

    two DES is multifactorial, and may be related to the

    underlying stent design, polymeric coating, mechanism of

    drug action, drug-release kinetics, and drug distribution

    across the vessel wall [245]. The hydrophobic/hydrophilicnature of drug molecules is also a vital parameter in this

    application [246]. The stent surface is exposed to blood

    from the time it is mounted on a balloon catheter till it is

    expanded at the tissue in the target area. If the drug is

    hydrophilic, it can possibly be lost in the blood due to its

    high solubility [246]. The loss of hydrophobic paclitaxel is

    less than 5% for a 30 s exposure time because of its low

    solubility in blood [111]. Though the material aspects

    contribute the most to the success of DES in reducing the

    restenosis percentage, the physicochemical characteristics

    of drugs have their own contribution [246248].

    Though the sirolimus and paclitaxel eluting stents have

    shown promising short-to-medium term clinical results,

    recently published reports on the occurrence of LST

    (occurs after 30 days) with adverse clinical events have

    raised concerns. A clinical trial that compared the efficacy

    of BMS and DES reported the occurrence of early stent

    thrombosis (occurs within 30 days) as 11.5% for both

    stent categories and no difference was observed in the

    occurrence of early stent thrombosis [249]. The occurrence

    of late thrombosis in patients treated with BMS and DES

    vary from 0.65% to 0.76% [250,251] and 0.35% to 0.7%

    [10,11], respectively. Though the reason(s) for the occur-

    rence of LST in DES is still unknown, the factors that

    could contribute may be the following: (a) delayed

    ARTICLE IN PRESS

    G. Mani et al. / Biomaterials 28 (2007) 16891710 1701

  • 7/30/2019 Swamy Science

    14/22

    endothelialization [10,252,253]the reasons for the delay

    in vessel wall healing after the implantation of DES are not

    yet clearly known. However, there is a concern that the

    nature of therapeutics used and their concentration/

    distribution across the vessel wall may affect the healing

    [252,253]; (b) adverse effects of the polymer coatings

    [254]polymer coatings like parylene, PEVA and PBMAhave been traditionally used for coating blood-contacting

    devices for their quality of adherence to the metal surfaces

    and/or their hydrophobicity. It is puzzling that such

    coating materials did not provide optimal results. This

    exposes the lack of knowledge about diseased tissuebio-

    material interactions; (c) discontinuation of antiplatelet

    therapyLST was observed among the patients treated

    with DES when they stop taking antiplatelet medications

    [252]; (d) neointimal growth for a longer periodthe

    growth of neointima reaches the peak at 6 months for BMS

    and regresses after that [255]. On the contrary, the

    neointima grows up to 4 years in DES [255]; (f) increased

    length of DES [256258]long DES were implanted to

    treat the entire diseased portion of the artery. This was

    reported to cause problems during deployment and

    positioning of stents in the arteries, which may eventually

    result in abnormal shear stress and cause thrombosis [259].

    Also, the longer stents increase the area of polymer

    coatingtissue interactions. These studies clearly show the

    need for further research in this area and that the currently

    available DES are far from optimal.

    7. Conclusion

    From a review of the literature it is evident that thematerial used for making stents has to have appropriate

    mechanical properties, suitable surface characteristics,

    excellent haemocompatibility, good biocompatibility, and

    drug delivery capacity. Every material has its own pros and

    cons. Table 3 provides a list of materials which posses the

    ideal for a specific material property (Table 3). It may not

    be possible for a single material to posses all the desired

    requirements. So, the success lies in choosing the optimum

    combination of materials and properties for the coronary

    stent applications.

    Though DES emerged recently, they appear to be the

    future of coronary stents. Ever since the FDA approved

    DES, the commercial availability of these stents hasincreased rapidly. However, it will take several years for

    this approach to become optimized once the long-term

    outcomes of the clinical trials are reported. The occurrence

    of late stent thrombosis in the patients treated with DES

    has raised concerns about these stents. Additionally,

    several cases have been reported recently on hypersensitiv-

    ity reactions to DES [254,260262]. In a pathological study

    of stent-related hypersensitivity reactions, it was noted that

    the polymer-coated stents released polymer fragments

    which were surrounded by giant cells and eosinophils

    [254]. Stents were also found to induce inflammatory

    reactions predominantly consisted of T lymphocytes and

    eosinophils with extensive inflammation of the arterial wall

    [254,260,262]. The FDA has posted a cautionary view

    about the adverse and hypersensitive reactions following

    deployment of sirolimus-eluting CYPHER stents

    [261,263,264].

    In conclusion, in its present form, percutaneous trans-

    luminal coronary angioplasty cannot be performed without

    damaging blood vessels and eliciting restenosis. Drug

    elution at the target site is a clear solution to this problem.

    However, the present methods for drug elution are still

    plagued with problems. Most commercially available DES

    use polymer matrices for coating and releasing the drugs.

    Increasing evidence suggests that some adverse reactionsmay be caused by these polymers. Hence, research should

    be carried out in designing and developing new polymer

    materials and should include essential features like

    hemocompatibility, hydrophobicity, anti-inflammatory,

    conformability to the stent surface, flaking resistance,

    sterilizability, and biodegradability. Other approaches such

    ARTICLE IN PRESS

    Table 3

    Materials with ideal characteristics for coronary stent applications

    Properties Materials Rationale

    Elongation modulus 316L stainless steel Optimal value for a balloon expandable stent

    Tensile strength CoCr Higher value

    Yield strength CoCr Much lesser when compared to its own tensile strength

    Surface energy PTFE Lower value

    Biocompatibility Ti Extensive literature

    Presence of stable oxide layer

    Surface potential Ta Stability of surface oxide layer

    Surface texture Electropolishing Best polishing technique to-date

    Stability of surface oxide layer Ta/Ti Excellent stability among the implant materials

    Therapeutics Paclitaxel Hydrophobicity

    Radiopacity Gold High density

    MRI compatibility Ta/Ti/Nitinol No Fe content

    Preferred way of drug loading Polymer based Amount of drug can be increased to the need just by increasing the thickness

    of the coating

    Preferred way of drug elution Biodegradable No polymer material will be present once the process is finished

    Pref erre d cat egor y of po lymers B iop olyme rs Mini ma l in flammatory a nd h yper sen siti ve reac tio ns

    G. Mani et al. / Biomaterials 28 (2007) 168917101702

  • 7/30/2019 Swamy Science

    15/22

    as nanoporous coatings and using self-assembled mono-

    layers [265268] for drug delivery also have potential

    applications in the next generation of stents.

    References

    [1] Shepherd RFJ, Vlietstra RE. The history of balloon angioplasty. In:

    Vlietstra RE, Holmes DR, editors. Percutaneous transluminal

    coronary angioplasty. Philadelphia: F.A. Davis Company; 1987.

    p. 117.

    [2] Myler RK, Stertzer SH. Coronary and peripheral angioplasty:

    historical perspective. In: Topol EJ, editor. Textbook of interven-

    tional cardiology. 2nd ed. Philadelphia: W.B. Saunders Company;

    1994. p. 17185.

    [3] Serruys PW, Jaegere PD, Kiemeneij F, Macaya C, Rutsch W,

    Heyndrickx G, et al. A comparison of balloon-expandable-stent

    implantation with balloon angioplasty in patients with coronary

    artery disease. New Engl J Med 1994;331(8):48995.

    [4] Fischman DL, Leon MB, Baim DS, Schatz RA, Savage MP, Penn I,

    et al. A randomized comparison of coronary-stent placement and

    balloon angioplasty in the treatment of coronary artery disease.New Engl J Med 1994;331(8):496501.

    [5] Holmes J. State of the art in coronary intervention. Am J Cardiol

    2003;91(3A):50A3A.

    [6] Wolf MG, Moliterno D, Lincoff A, Topol E. Restenosis: an open

    file. Clin Cardiol 1


Recommended