Download - Swamy Science
-
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