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Biosurface and Biotribology 1 (2015) 146–160

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Plasma-synthesised carbon-based coatings for cardiovascular applications

M. Santosa,b,c,n, M.M.M. Bilekb,n, S.G. Wisea,c,d

aThe Heart Research Institute, Applied Materials Group, Sydney 2042, NSW, AustraliabSchool of Physics, Applied and Plasma Physics, University of Sydney, Sydney 2006, NSW, Australia

cSydney Medical School, University of Sydney, Sydney 2006, NSW, AustraliadSchool of Molecular Bioscience, University of Sydney, Sydney 2006, NSW, Australia

Received 29 June 2015; accepted 5 August 2015

Abstract

Current cardiovascular stent platforms interact poorly with the human vasculature and still rely on drug therapy to avoid early failure due to bloodclotting. A drug-free coating technology that could fully integrate an implanted stent through a combination of hemocompatibility and differential regulationof endothelial cells and smooth muscle cells would present a clear advantage over existing clinical approaches. Plasma discharges have been used as acoating technology in a wide range of applications over the last decades. Carbon-based thin films prepared by different plasma deposition methods areusually regarded as biocompatible materials as they are able to prevent the adhesion and activation of platelets and preferentially promote the adsorption ofalbumin over fibrinogen. However, the available literature seldom addresses entirely the aspects of biocompatibility and the challenging mechanicaldemands of such materials for stent coating. Recent advances suggest that plasma enhanced chemical vapour deposition can be used to prepare carbon-based thin films that allow for the linker-free immobilization of bioactive molecules. If successfully applied to a stent these coatings could represent a steptowards stent specific biofunctionalization. This review examines the feasibility of using plasma discharges for the synthesis of carbon-based biocompatiblematerials for cardiovascular implantable devices, particularly stents.& 2015 Southwest Jiaotong University. Production and hosting by Elsevier B.V. This is an open access article under the CC BY-NC-ND license(http://creativecommons.org/licenses/by-nc-nd/4.0/).

Keywords: Plasma deposition; Carbon coating; Functionalization; Biocompatibility; Cardiovascular; Stent

Contents

1. Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1471.1. Current coronary stent platforms . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1471.2. Plasma activated biofunctional stents . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 148

2. Plasma synthesised carbon-based coatings. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1482.1. Diamond-like amorphous carbon and polymer-like carbon-based thin films . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1492.2. Improving the mechanical stability of carbon-based coatings . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1502.3. Biocompatibility. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 151

2.3.1. Adsorption of blood proteins . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1512.3.2. Surface functionalisation with immobilised bioactive molecules . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 152

3. Translation of carbon-based coatings to coronary stents . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1533.1. Challenges . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 153

/10.1016/j.bsbt.2015.08.00115 Southwest Jiaotong University. Production and hosting by Elsevier B.V. This is an open access article under the CC BY-NC-ND licensemmons.org/licenses/by-nc-nd/4.0/).

nce to: Applied Plasma Physics Group, School of Physics, A28 University of Sydney, Sydney, Australia. Tel.: þ61 2 93516079;7725.sses: miguels@physics.usyd.edu.au (M. Santos), marcela.bilek@physics.usyd.edu.au (M.M.M. Bilek).nder responsibility of Southwest Jiaotong University.

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3.2. Plasma-assisted coated stents . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1544. Conclusions and future perspective . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 154Acknowledgements . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 157References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 157

1. Introduction

Every year more than 6 million people undergo surgeryaround the world for the insertion of medical implants such aship and knees prostheses, pacemakers, heart valves and coronarystents [1,2]. In the US alone around 1 million orthopaedicimplants are inserted annually for knee and hip replacements [3]and these numbers are expected to rapidly increase to around 3million by the year 2030 [4]. The increasing life expectancy ofthe population elevates demands for medical implants with asuperior life span that can resist corrosion and wear, especially inload bearing and blood and/or tissue contacting applications.Additionally, insertion of exogenous materials into the bodytriggers immune responses in the areas surrounding the implantwhich can compromise implant performance and even necessi-tate revision surgery for implant replacement or extraction.Novel surface technologies that could mediate interactionsbetween the implant and the host to prevent adverse reactionswould clearly benefit the biomedical devices industry. Inparticular, there is a great demand for a universal technologycapable of producing tailored biofunctional interfaces to simul-taneously address a number of specific tasks depending on theapplication. Modulation of the surface properties such ashemocompatibility, protein adsorption / immobilization, differ-ential attachment, proliferation and differentiation of cells arehighly desirable. This review investigates the potential of usingplasma-based tools as a universal coating technology forimplantable medical devices, with particular focus on coronarystents. The deposition of biofunctional surfaces on coronarystents is a particularly interesting and complex case study as itcombines: (i) physical–chemical, (ii) mechanical, (iii) geometric,(iv) blood clotting and (v) cellular response considerations.

1.1. Current coronary stent platforms

The World Health Organization [5] estimates that cardiovasculardiseases represent 30% of total deaths worldwide, with coronaryheart complications, including atherosclerosis, leading the list witha total of 7.3 million deaths every year. The same organizationpredicts that these numbers will rise up to 8% within the next 15years, exacerbating an already serious problem. In the last threedecades, percutaneous coronary interventions became widely usedin the treatment of cardiovascular diseases and have beenestablished as the first-line technique for revascularization of thecoronary arteries. Among implantable biomedical devices, coron-ary stents are now the dominant vascular implant in percutaneouscoronary interventions [6] with around 800,000 stents implantedannually in the US alone [7]. Stents are expandable cylindricalmeshes used to re-establish the normal blood flow in blockedatherosclerotic coronary arteries, resupplying ischemic tissue.

However, despite their widespread use, current stent platformshave only sub-optimal biocompatibility, interacting poorly withvascular cells and promoting blood clot formation.The first commercially available stent platform was based on

metallic alloys such as stainless steel (SS). Despite the extra-ordinary results in reducing the rate of abrupt vessel closure, oneof the major pitfalls of coronary angioplasty, post-procedurecomplications related to stent implantation started to arise. Themost common was in-stent restenosis, triggered by an inflam-matory response due to blood vessel intima injury during stentdeployment. Damage to the endothelium led to the migration andover-proliferation of underlying smooth muscle cells (SMC),causing vessel re-narrowing in a process called neointimalhyperplasia [8]. In order to overcome the draw-backs of baremetal stents, drug eluting stents (DES) were introduced. DESsare grafted with layers of biodegradable [9–11] or non-biodegradable [12–14] polymers, in which a pharmacologicalagent is loaded and can be locally released within a given periodof time. By incorporating anti-proliferative agents, drug-elutingstents were successful in inhibiting the over-proliferation ofSMC, hence reducing the effects of hyperplasia. However, anunintended consequence of DES was a substantially increasedrisk of long-term stent thrombosis, termed late stent thrombosis(LST) [15]. LST is mainly associated with reduced re-endothelialisation of the stent, leading to a delayed healing ofthe vessel and over-exposure of the stent struts to the blood.Another possible reason for late thrombosis could be associatedwith adverse inflammatory responses to the drugs and the graftedpolymer used in these stents. Extensive research has beenundertaken to address the issues of in-stent restenosis and late-stent thrombosis, either through the development of morebiocompatible polymer-based coatings [16,17] or by integratingspecific antibodies with the ability to recruit and immobilizeendothelial progenitor cells [18–20], thus promoting re-endothelialisation. The incorporation of nitric oxide (NO) donorsin the stent design has also been considered as an encouragingalternative platform to reduce in-stent restenosis and LST. NO isan endogenous signalling molecule that participates in importantbiological processes and possesses several vasculoprotectiveproperties [21]. It has been shown that the administration ofdifferent NO donors inhibits the proliferation and migration ofSMC [22,23], enhances the proliferation of endothelial cells [24],prevents platelet aggregation [25] and adhesion [26] to thevascular endothelium and reduces intimal hyperplasia followingvascular injury [27]. The delivery of NO through eluting stentplatforms has been achieved by incorporating NO donors in apolyurethane polymer which was then coated onto the stents [28]or by incorporating the NO donor with paclitaxel [29] (an anti-proliferative agent extensively used in the prevention of rest-enosis). However, no significant improvements in preventing

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restenosis were shown so far relative to the controls in differentin-vivo animal trials.

A new generation of bioabsorbable stents [30–32] have beenprompting increasing interest from the scientific community. Intheory, these stents can be temporarily used has a treatment foratherosclerosis as they are dissolved in the body after a periodof time, leaving behind a structurally normal and recoveredblood vessel. Although promising, some issues regarding theirmechanical properties, which are paramount for a proper stentdeployment, radio-opacity and control over the rate that theyare absorbed in the body still need to be addressed [33].

Currently used DES platforms still carry an ongoing risk ofthrombosis requiring patients to be on pharmacological dual anti-platelet therapies, which can result in bleeding and is not feasiblefor patients requiring additional surgery. A drug-free coatingtechnology, with all the necessary mechanical and biologicalproperties for enhancement of the clinical performance ofcoronary stents, would provide a clear benefit over current stentplatforms. In particular, the coating should be able to resistcracking and delamination during and after stent deployment inorder to avoid the contact of the struts with the blood stream. Itshould also be multi-functional, promoting the adhesion andproliferation of endothelial progenitor cells, for re-endothelialisa-tion, while also inhibiting the over-proliferation of smoothmuscle cells to avoid hyperplasia.

1.2. Plasma activated biofunctional stents

Waterhouse et al. [34] reported the feasibility and safety ofdelivering a prototype plasma-activated coating (PAC) in vivo.This coating was deposited using hydrocarbon-based radio-frequency plasma discharges and prepared with energetic ionsand a graded interface between the stainless steel surface andthe PAC coating to provide sufficient adhesion. Followingsome initial optimisation of mechanical and biological proper-ties the coating was applied to a stent and implanted in rabbitiliac arteries. Results showed that PAC had sufficient adhesion,undergoing stent crimping and expansion, and performedequivalently in terms of re-endothelialisation when comparedto bare metal stent platforms.

Plasma discharges have been extensively used as a coatingtechnology in a vast number of applications. The extremeversatility of plasma-based systems could be fostered to providea universal coating platform for biocompatible and multi-functional cardiovascular implants. A considerable number ofpublications have reported the viability of using plasma-basedtechnologies for the manufacture of biofunctional coatings forcardiovascular implants. These works were recently reviewed byWise et al. [35]. Efforts have been made to deliver mechanicallyrobust coatings with sufficient adhesion strength to metallicsubstrates. Wear and corrosion resistance due to fluid permeationafter prolonged exposures to body fluids has also been a concern.With regards to coating biocompatibility, it is still not clearwhich physical and/or chemical properties play the mostimportant roles. Bonding configuration, chemical composition,presence of dopants, surface roughness and wettability have beenregarded has important factors predominantly affecting blood

protein adsorption and cellular response such as platelet adhesionand activation.Previous studies [36,37], also showed that carbon-based

polymer-like coatings produced by plasma-enhanced chemicalvapour deposition (PECVD) with enhanced ion-bombardmentpossess superior biological qualities over conventional coatingalternatives (e.g. chemical vapour deposition, physical vapourdeposition and wet techniques such as solvent dipping).Importantly, these coatings also allow the covalent immobili-zation of biomolecules [38]. The possibility to conjugatebioactive molecules to biomedical implants is particularlyinteresting. The development of more biocompatible stentsmay be achievable through a specific and controlled functio-nalization via the stable immobilization of bioactive proteins orpeptides that can interact with the vasculature, especially bypromoting endothelialisation, inhibiting the proliferation ofsmooth muscle cells and reducing the formation of thrombus.

2. Plasma synthesised carbon-based coatings

Among the four fundamental states of matter, (i.e. solid,liquid, gaseous and plasma), plasma is the most abundant inthe Universe. For instance, the stars, interstellar and inter-galactic space, nebulae, lightning and auroras borealis are all inthe form of plasma. The plasma state is achieved when enoughenergy is supplied to a gas, or a gaseous mixture, so that afraction of the atoms (or molecules) that constitutes thatmedium becomes ionised. In contrast to a gaseous medium,formed only by non-charged species (atoms and/or molecules),the plasma medium is comprised not only of non-charged butalso by charged species (ions and electrons) and photons.Therefore, the electrical and thermodynamical propertiesinherent to a plasma medium are far more complex and subtlethan in the gaseous state. The complex interaction betweenpositively and negatively charged particles gives rise tocomplex phenomena and collective behaviours, which areunique to the plasma and may vary depending on theconditions in which the plasma is produced and sustained.Plasmas can also be produced in laboratory conditions.

Humankind has been studying and developing different waysto produce and control plasmas for its own benefit for morethan 130 years. Plasma discharges have found a large numberof applications from illumination, metallurgy, micro-electro-nics, environment, materials engineering, analytical spectro-scopy, medicine, etc [39]. The success of man-made plasmadischarges is a result of the unique physical and chemicalcharacteristics inherent to this medium, as well as the extra-ordinary flexibility achieved so far in controlling a largewindow of working conditions. Plasma can be artificiallyproduced in a wide range of size scales (from micrometre upto several metres), geometries (planar, cylindrical, spherical,etc.), densities (from 107 up to 1032 particles per cubic meter),pressures (from a few mTorr up to atmospheric pressure),temperatures (from negative values up to 108 1C) and lifetimes(from femtoseconds up to years).Depending on the type of application, many different

techniques of plasma generation can be used. For instance,

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in surface modification and deposition applications it isdesirable to achieve high levels of purity in the coatings.The use of a vacuum chamber, where the system is pumpeddown to very low pressures, allows control of the plasmacomposition and, therefore, also the coating composition.Magnetic fields and high voltage sources can also be used togain control over trajectory and energy of the charged speciesthat reach the surfaces to be coated. In the last decades forinstance, different types of plasma discharges have proventheir potential for biomedical applications, for a wide range oftasks such as sterilisation, cell proliferation, treatment of oralpathogens and healing of burn wounds [40]. Low pressure andlow temperature plasmas have become widely used for surfacemodification in different applications, including the depositionof thin-films on implantable biomedical devices. Organosiliconand hydrocarbon precursors are commonly used for thecreation of hydrogenated silicon films and amorphous hydro-genated carbon films (frequently named diamond-like carbonfilms) for a wide range of applications. The rich compositionof the plasma medium, i.e., electrons, ions, atoms, radicals,molecules and photons, is achieved through efficient dissocia-tion and ionization of precursor and carrier gas atoms andmolecules, which takes place via many different kineticmechanisms and energy transfer channels. The resultingprecursor ions and radicals then diffuse towards the substrate,initiating a set of surface chemical reactions and ultimatelyleading to the deposition and growth of a coating on thesubstrate surface.

PECVD, for instance, is a popular and extremely versatileplasma-based technology for the deposition of thin films [41].The plasma is generated at low temperatures (typically ambienttemperature) inside a vacuum chamber and is typicallysustained by capacitively coupling radio-frequency (rf) power,usually at a frequency of 13.56 MHz, through a matchingnetwork so as to minimise power losses within the externalelectrical circuit. The rf power can be coupled to the plasmaeither in continuous or in pulsed regimes [42]. A pulsed powermodulation regime is frequently used in plasma polymerizationrepresenting an extra degree of freedom in the depositionprocess. Samples can also be electrically biased to acceleratethe plasma ions towards the growing film. The applied bias canbe either continuous DC [43], radio-frequency [44] or pulsed[45]. Control over the chemical composition of the coatings isachieved by choosing different gases (or gaseous mixtures) togenerate the plasma. The coatings deposited by PECVD,especially when bias is used to increase the energy of thedepositing species, are characterised by superior adhesion tothe substrates, high cross-linking, improved biocompatibilityand chemically active and functional surfaces. This depositiontechnology also allows a level of control over differentproperties of the films, such as roughness, wettability, hard-ness, density, flexibility and composition without changing theproperties of the underlying substrate. This control is achievedby tuning the different process parameters—a difficult taskgiven the complexity of the plasma and the use of reactivegaseous mixtures. Therefore, the process optimization for thedeposition of thin-films with desired features is time-

consuming and requires further characterization and under-standing of the plasma medium.It is paramount, therefore, to understand how changes in the

process parameters, e.g. power coupled to the plasma, pres-sure, temperature, sample bias and different gas ratios affectthe plasma physical quantities and how sensitive the coatingproperties are to those. Process monitoring via different plasmadiagnostic techniques using electrostatic probes [46], opticalemission spectroscopy (OES) [47] and mass spectroscopy [48]can ultimately assist in modulating key features in the finalcoating. For instance, the radiation emitted by the plasma dueto electronic, vibrational and rotational atomic and moleculartransitions can be analysed by means of OES and infraredspectroscopy. The plasma optical signature retains pertinentinformation about the most populated atomic and molecularstates within the plasma. Such information can be used toderive the relevant channels of creation and destruction ofthose populations. Electron and ion temperature, density andtheir transport in the discharge can also be understood throughelectrical diagnostics, using for instance Langmuir probes andquadrupole mass spectrometers. The electron energy distribu-tion function (EEDF) can also be obtained using the formerdiagnostics. Important macroscopic parameters can be deter-mined by appropriate integrations of the EEDF [49] in energyspace. These include, for instance, electron mobility, diffusioncoefficients, drift velocity, characteristic energy, powerabsorbed by the plasma, plasma conductivity and all theexcitation / de-excitation, dissociation and ionization coeffi-cients related with the plasma kinetic mechanisms involvingcollisions between electrons and other populations. There isalso a great demand for simulation models capable of describ-ing such discharges and their operation. Plasma self-consistentmodelling with the necessary complexity to describe the mainplasma kinetic mechanisms, the transport of species andplasma / surface interactions are seldom available. Complexhybrid simulation tools [50] comprising state-of-the-art colli-sional–radiative models [51], plasma fluid models [52] andplasma / surface interactions could be developed to addressfundamental investigations, reactor design and optimization toultimately provide stabilization and reproducibility of thefabrication process. Research combining plasma modelling,plasma experimental diagnostics and coating characterizationtechniques will yield further important insights on howdischarge parameters affect the coating properties.

2.1. Diamond-like amorphous carbon and polymer-likecarbon-based thin films

The exceptional physical, chemical and tribology propertiesof plasma synthesised diamond-like carbon (DLC) haveprompted an increasing volume of research since they werefirst synthesized in the form of thin films in 1971 [53]. Thepotential use of DLC as a biocompatible coating for medicalimplants in a wide range of applications has been extensivelyinvestigated in the last 2 decades. A considerable number ofreview papers covering these investigations are now available[54–58]. The use of DLC coating for orthopaedic prostheses

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[59] and cardiovascular implants such as heart valves [60],guiding wires and coronary stents [61] are among the mostcommon examples.

The designation DLC is commonly attributed to a variety ofamorphous carbon-based materials that may contain differentsp2 and sp3 carbon bond ratios as well as different hydrogencontents. Depending on the bonding configuration and thehydrogen content, DLC materials are usually sub-classifiedinto: (i) tetrahedral amorphous carbon (ta-C) for higher sp3

contents (typically 470%) and low levels of hydrogen; (ii)amorphous carbon (a-C) if the ratio sp3/sp2 ranges between 0.4and 0.7 and (iii) hydrogenated amorphous carbon (a-C:H) forhydrogen containing materials which are typically charac-terised by higher sp3/sp2 ratios but they are soft because manyof the sp3 bonds are hydrogen terminated. The properties ofDLC thin films, their mechanisms of deposition and theircharacterization techniques were extensively reviewed byRobertson [62] and McKenzie [63].

Highly hydrogenated soft a-C:H thin films are usuallydenominated by hydrocarbon plasma polymers [64–66]. Thesefilms are characterised by a polymer-like structure which isdeposited on the substrate material. The atomic arrangement ofthis structure is strictly dependent on the type of carbonprecursor (monomer) and the process parameters used duringthe deposition. Carbon-based polymer-like deposits are usuallyhighly cross-linked and may contain large concentrations ofimprisoned radicals. The mechanisms of plasma polymeriza-tion are still not fully understood and several models have beenproposed in the last decades and these were also recentlyreviewed [67].

Carbon-based materials can be prepared as a coating in theform of a film using a wide range of plasma-based depositiontechniques. Different properties of the DLC and polymer-likecarbon coatings can be tuned by setting adequate depositionparameters. For instance, ta-C thin-films prepared by arcvapour deposition [68,69] are characterised by extreme hard-ness (up to 80 GPa), great wear and corrosion resistance andchemical inertness. These properties make ta-C potentialcoating candidates for load-bearing biomedical applications[70]. However, the high internal compressive stress character-istic of thick ta-C films still remains the primary source ofcoating failure in orthopaedic implants [69]. The exposure ofthe underlying metallic substrate due to poor adhesion anddelamination presents one of the major obstacles to thecommercial success of DLC coatings.

a-C films can be prepared using a variety of plasma-baseddeposition technologies. For instance, a-C materials are usuallymade by plasma magnetron sputtering (PMS) where a graphitetarget is exposed to argon plasma [71,72]. Carbon atoms andions are sputtered from the graphite target and are then depositedonto the substrates. Substrates to be coated are usually electri-cally biased [73] which allows control over the kinetic energy ofthe carbon ions reaching the samples. Control over the biasvoltage is characteristically used to deliver films with differentphysical structures, surface topographies and sp3/sp2 ratios.

a-C:H are commonly prepared by PMS [74–76] or by plasmaenhanced plasma vapour deposition (PECVD) [77,78]. In the

first case, a reactive mixture of argon and hydrogen is injected inthe PMS reactor. Hydrogen atoms bind with the sputtered carbonatoms leading to the deposition and growth of a hydrogenatedcarbon thin film. The concentration of hydrogen gas during thedeposition process can also be adjusted to yield films withdifferent hydrogen contents, hence with diverse properties aswell. Highly hydrogenated a-C:H have reduced internal stress[55] resulting in improved film adhesion to the underlyingsubstrates [79]. Hydrogen also behaves as an etching elementand increasing the hydrogen concentration in the plasma wasreported to promote surface smoothness of the films [62].Control over the chemical composition of the DLC coating byPECVD is achieved by choosing different gases, or gaseousmixtures, to generate the plasma. While argon is commonly usedas the background gas, different hydrocarbon precursors (acet-ylene, methane, ethylene, ethane, butane, etc.) are used to tailordifferent chemical and physical structures of the a-C:H films. Forinstance, increased hydrogen contents and smoother surfaces canbe created by using methane (CH4) instead of acetylene (C2H2)as the precursor gas [80]. Also, the incorporation of hydrogen inCH4 plasmas was shown to reduce the coating friction and wearwhen deposited on steel substrates [81]. The incorporation ofdopants, such as nitrogen, phosphorous, fluorine or silicon canalso be easily performed by choosing the appropriate gaseousmixture. This feature is particularly interesting since the combi-nation of these elements with DLC was shown to be paramountin tailoring the mechanical and biological performance of the a-C:H coatings. It is now generally accepted that sample bias is afundamental process parameter in the production of DLC coat-ings. Films with variable sp3 fraction[82], residual stress [83],substrate adhesion [84], refractive index [85] and hemocompat-ibility [86] can be prepared by adjusting solely the sample biaswhile maintaining constant the other process parameters. Impor-tantly, sample bias promotes the creation of radicals in thecoating’s surface and bulk, allowing for a further functionaliza-tion of the coated materials with bioactive proteins, peptides andother biomolecules [36].

2.2. Improving the mechanical stability of carbon-basedcoatings

a-C and a-C:H made by PMS and PECVD are generallycharacterised by a reduced internal stress and improvedsubstrate adhesion when compared to ta-C films. However,and despite their excellent biocompatibility, these materialsstill can’t fulfil all the requirements necessary for a permanentand successful translation to real medical implants. Thesecoatings are particularly sensitive to wear and corrosion,especially when deposited on metal substrates and exposedto saline solutions or biological fluids. It has been observedthat the prolonged exposure of the coatings to fluids reducesthe interfacial strength, ultimately leading to delamination andspallation at the coating / substrate interface [61,87,88].Chandra et al. [88] proposed that failure of DLC coatingsafter exposure to fluids, such as water and phosphate bufferedsaline (PBS), was related with the permeation and perforation

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of the solution through porosities and failures localised at thecoating surface.

A great amount of research has been undertaken to addressthe issue of coating failure due to poor: (i) substrate adhesionand (ii) wear and corrosion resistance. A potential solution forthe first would be to increase the interfacial strength betweensubstrates and carbon-based coatings through the addition ofinterlayers [87–93]. The deposition of silicon-based thin layersbetween metallic substrates and DLC coatings, prepared by rf—PECVD, can significantly increase the adhesion strength ofDLC to stainless steel (SS) substrates. SiNx interlayers havebeen reported as particularly effective for the enhancement ofboth adhesion and coating corrosion resistance of samplesincubated in saline solution up to 2 years [87]. The deploymentof an a-Si:H to a-C:H graded interface (also prepared byrf—PECVD) was shown to increase the adhesion strength andcorrosion resistance of DLC on Ti alloys [88]. The addition of90 nm thick Si interlayers prior to DLC deposition on CoCrMosubstrates resulted in a substantial enhancement in the adhe-sion strength when compared to non-interlayered samples [92].

Further enhancement of DLC adhesion strength to metallicsubstrates can be fostered by creating a graded interfacebetween the metal surface and the carbon layer. Gradedmetallic interfaces were previously reported and performedon SS substrates prior to the deposition of carbon-basedplasma polymers (a-C:H ) using a low pressure PMS system[94]. A SS target is bombarded by energetic argon atomscoming from the plasma, which ultimately leads to thesputtering of metallic ions from the target and their subsequentdeposition on the substrate surface. If a hydrocarbon precursor(such as acetylene) is gradually added to the plasma during thesputtering process, it is possible to create a metal-carbongraded interface. This graded interface is characterised by anincreasing carbon to metal content from bottom to top with thetop layer being exclusively composed of DLC or a carbon-based plasma polymer. The adhesion strength of DLC coatingsto SS substrates prepared with graded layers was reported to begreater than 26 MPa. This value represents an improvement of260% when compared to coating deposited directly on thesame substrates [94].

2.3. Biocompatibility

Biocompatible coatings for cardiovascular implants shouldbe inherently non-thrombogenic and able to recruit andpromote the proliferation of endothelium cells while inhibitingthe attachment and overgrowth of smooth muscle cell (SMC)layers. They should fully integrate into the human vasculatureby prompting healing responses to prevent implant rejectionand severe immune responses post-implantation. Implanthemocompatibility can be achieved by preventing the adsorp-tion and further denaturation of undesirable blood proteins thatmediate the pathways of blood clot formation. The attachment,on the coating surface, of bioactive molecules that couldprevent platelet adhesion and activation and that could becapable of further inducing favourable local cellular responseswould represent a great advantage.

Although plasma synthesised carbon-based coatings wereshown to retain good hemocompatibility in various in-vitroessays, in-vivo investigations to access their capacity forendothelialisation, SMC inhibition and hemocompatibility areseldom reported. Good results of these coatings in the greatmajority of such in-vitro tests were explained by the higheraffinity of the surfaces to preferentially adsorb albumin overfibrinogen. However, in-vitro tests poorly simulate the harshconditions of the human vasculature such as the high flowregimes of blood in arteries. High flow regimes can lead todesorption of adsorbed or weakly attached proteins. For thisreason covalent immobilization is preferred over physicaladsorption. Recent advances [36] showed that PECVD pre-pared carbon-based coatings with high concentrations of long-lived unpaired electrons can be used for the covalent attach-ment of various biomolecules in their bioactive form. Theseadvances could establish a new era of truly bio-functional and-compatible cardiovascular implants.

2.3.1. Adsorption of blood proteinsThe biocompatibility and particularly the hemocompatibility

of a blood-contacting surface are strongly affected by theaffinity of the surface to adsorb specific blood proteins. Thetime scale on which blood proteins interact with and areadsorbed on a coating’s surface is several times smaller thanthat of any of the interactions involving other blood compo-nents. For this reason, local cellular response to a foreign bodyimplanted in a blood vessel will be in part governed by thenature of the proteins adsorbed on its surface. For instance, it isknown that adsorbed fibrinogen (Fib) enhances the adhesionand activation of platelets [43], triggering the coagulationpathway. Furthermore, Fib is also a fundamental element in thecoagulation cascade since it is cleaved into fibrin to form anirreversible thrombus. On the other hand, albumin can inhibitthe adhesion and activation of platelets, preventing the forma-tion of a clot.The mechanisms governing the adsorption of different blood

proteins on carbon-based surfaces are not yet fully understood.The subject has driven increasing research in the last twodecades and although some advances have been achieved,important questions remain unanswered. Difficulties arise indetermining exactly which physical and/or chemical propertiesof the surfaces rule protein adsorption. Available data suggestsurface roughness and surface energy, hydrogen and nitrogenfraction, sp3 content and electrical band-gap as possiblecandidates. However, these properties are generally interlinkedand their decoupling is largely limited by the depositiontechnique used to prepare the surfaces.It has been suggested that the adsorption of blood proteins

on a DLC surface is related to the charge (electrons) transferbetween the proteins and the surface [56,95]. Protein adsorp-tion and further denaturation occurs if the Fermi energy levelof the protein is such, that electrons from its valence band aretransferred to the conduction band of the coating material [95].Therefore, hemocompatible coatings should be made withelectron energy band gaps greater than those of proteins thatare responsible for the formation of thrombus—such as

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fibrinogen, which has a band gap of 1.8 eV [96]. On the otherhand, adsorption of albumin would be favourable as it caninhibit the activation of platelets on the surface of the film.

It has also been reported that surface roughness andwettability may influence the adsorption of blood proteinsonto the surfaces [80,97–99]. Albumin for instance wasdescribed to have greater affinity for more hydrophilic androugher ta-C surfaces, made by filtered arc deposition, whencompared to smoother and less hydrophilic a-C:H thin filmsprepared by plasma-enhanced chemical vapour deposition(PECVD) [80]. However, different chemical compositions(e.g. hydrogen content) and bonding configurations betweenthe different samples may have also influenced proteinadsorption. Using spectroscopic ellipsometry (ES) [97] sug-gested that different a-C:H coatings made by plasma magne-tron sputtering can adsorb human serum albumin (HSA) andhuman fibrinogen (Fib) differently. In this work the surfaceconcentration ratio (proportional to the thickness of theadsorbed protein layer) HSA/Fib�1 was higher for samplesprepared with floating potential whereas surfaces made using anegative bias presented a ratio HSA/Fib�0.4. Later the sameauthors [85] showed that a-C:H coatings made under the sameconditions presented a different behaviour with respect to HSAand Fib adsorption. The HSA/Fib thickness ratio was found toincrease on samples prepared with bias, which were identifiedby AFM as being rougher. Berlind et al. [99] studied HSAadsorption on different amorphous, graphitic and fullerene-likesurfaces also using spectroscopic ellipsometry. They found thatamorphous structures, i.e. with smoother surfaces, adsorbedmore HSA when compared with graphitic and fullerene-likethin films. However, the authors could not correlate a higherHSA adsorption exclusively with the roughness of the thinfilms. This was because the different films also presenteddifferent chemical structure, wettability and water-bindingaffinities which are important parameters that also influenceprotein adsorption. Available data suggests that blood proteinadsorption is a complex phenomenon and there is still no clearconsensus on its drivers. Additional investigations are there-fore necessary to understand the basic phenomena of protein/surface interactions and how different features in the surfacescan regulate differential protein absorption.

2.3.2. Surface functionalisation with immobilised bioactivemolecules

Recent studies found that rf-PECVD carbon-containing plasmapolymers prepared with bias have the ability to covalentlyimmobilise proteins in their bioactive form [36,38]. Organicprecursor plasma discharges can be used to deposit hydrophilicplasma activated radical-rich a-C:H-like coatings (PAC) if suffi-cient sample bias, usually up to �1 kV, is provided [100]. Here,the ions formed in the plasma are accelerated with increasedenergy provided by the high voltage bias towards the growingcarbonised film on the substrate surface. Ion impact promotes theformation of regions of carbon extended states such as πconjugated sp2 carbon bonds with a high degree of cross-linking. These regions contain stable unpaired electrons that canmove and diffuse towards the coating surface. Free-radical

content, measured by electron paramagnetic resonance spectro-scopy, was found to be proportional to the bias voltage andthickness of the deposited film [94]. Colorimetric assays usingimmobilised horseradish peroxidase catalase as a probe have beenemployed to confirm the capability of PAC coatings to retain thebioactivity of the immobilised protein layer [38,45,101]. A varietyof polyamino acids have also been found to be covalently attachedto PAC surfaces via chemical bonding with the unpaired electronson the film’s surface [36].Metallic substrates such as stainless steel when coated with

PAC alone are significantly more hemocompatible, a propertyrevealed by its ability to considerably reduce the formation ofhuman thrombus in different in-vitro assays. Blood compat-ibility is an essential characteristic for biocompatible cardio-vascular stents and although necessary, it does not addressendothelialisation and restenosis. A stable immobilization ofbioactive biomolecules on the coating surface could represent aleap forward towards the true biofunctionalization of cardio-vascular implants such as arterial stents. Plasma-made radical-rich coatings for stents can be envisaged, immobilising theappropriate biomolecule or biomolecule cocktails, to impartappropriate local biological responses to achieve full vascularintegration of stents (see Fig. 1).The advances discussed above suggest that chemical conjuga-

tion of proteins through the reactions of unpaired electrons withprotein amino acid side chains is possible while preventing proteindenaturation. It was shown that PAC surfaces could be used forthe immobilization of tropoelastin in a bioactive conformation onflat stainless steel substrates [37]. Tropoelastin, being a precursorof elastin, is an important vascular protein since it participates inthe correct structuring and functionalization of the whole vascularsystem. Elastin provides fundamental rheological properties toblood vessels, facilitates endothelialisation and constitutes(together with collagen) the sub-endothelia layer, the lamina,interfacing the intima and the media. As an essential component ofthe lamina, elastin fibres play an important role in avoidinginfiltration of smooth muscle cells towards inner layers of theblood vessel while promoting and regulating the growth andproliferation of luminal endothelial cells to form a healthyendothelium. Covalently immobilised tropoelastin and specificdomains of tropoelastin were shown to drastically enhance thehemocompatibility of PAC coated SS substrates in in-vitro assaysusing human whole blood [102]. Endothelial cell attachment andproliferation were also improved when compared to uncoatedsamples.More recently, PAC surfaces functionalised with bioactive

plasmin revealed superior human blood compatibility andendothelialisation [103] when compared to SS controls.Plasmin is an enzyme, derived from plasminogen, present inthe blood that cleaves fibrin clots in a process calledfibrinolysis. When conjugated with PAC, plasmin is able toprevent the formation and deposition of fibrin complexesreducing the probability of clot development. Hemocompat-ibility and good in-vitro proliferation rates of endothelial cellssuggest that stents functionalised with PAC immobilisedplasmin could prevent early thrombogenicity while promotingendothelialisation for stent vascular integration.

Fig. 1. Proactive biofunctionalization of coronary stents using plasma discharges in carbon-based atmospheres. The stent is electrically biased to accelerate plasmapositive ions towards the growing coating, creating therefore a radical-rich plasma activated coating (PAC). Biomolecules are then immobilized on the coating viaradical bonding and they maintain a bioactive conformation after immobilization. Following stent implantation, the immobilized biomolecules induce favourablelocal cellular responses promoting endothelial cell migration, adhesion and proliferation on the implant surface to form a healthy endothelium. Smooth muscle cellproliferation and platelet activation is inhibited, preventing neointimal hyperplasia and the formation of thrombi to achieve favourable integration of the implantedstent in the host vasculature.

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PAC conjugated with constituents of high density lipoproteinswere also recently studied for hemocompatibility, endothelializa-tion and smooth muscle cell proliferation [104]. A low concentra-tion of high density lipoproteins in the blood was previouslyassociated with restenosis and thrombosis after stent deploymentand coronary angioplasty. PAC was shown to covalently immo-bilise constituents of HDL while retaining its bioactivity. Stainlesssteel PAC coated substrates functionalised with HDL strikinglyreduced the formation of human blood clots in-vitro. Moreover,HDL coated surfaces encouraged endothelialisation and preventedthe over-proliferation of smooth muscle cells.

3. Translation of carbon-based coatings to coronary stents

Significant efforts have been made to develop biocompatiblecoatings for metallic cardiovascular implants and especially forcoronary stents. Although plasma discharges are perceived as apromising solution to fulfil this task, a great deal of researchstill needs to be carried out. Despite the increasing volume ofinvestment and research, there is still no available plasma-based coating technology for stents proved to successfullyaddress all the aspects of biocompatibility in humans.

3.1. Challenges

Stents are three-dimensional objects with wide-ranging dimen-sions, aspect ratios, complex geometries and strut designs.Consequently, coating uniformity using appropriate plasma dis-charges such as low pressure PECVD or PMS is a non-trivial task.One of the main challenges is to avoid edge and shadowing effectson the stent struts caused either by adjacent struts or by theelectrical connections used to support the stent in the discharge. rf-PECVD discharges can provide stable plasma media that uni-formly surround flat surfaces during the deposition process. Also,when using sample biasing, ions are accelerated uniformlytowards the full length of flat surfaces when a uniform sheathlies between the plasma and the sample surface. However, the

intricate design of a stent and the use of electrical and supportingcontacts give rise to the development of complex plasma sheathsthat are not totally conformal to the stents. For this reason athickness gradient in the coating across the stent outer surface isusually hard to avoid. Moreover, while the outside struts aredirectly coated and bombarded with energetic ions from thesurrounding reactive plasma, the inner struts are exposed only tothe poor plasma and less energetic ion flux inside the stent.Therefore, the physical and chemical properties of the coating aregenerally spatially dependent. Coating thickness, for instance, isusually thinner on the inside struts. Importantly, the bondingconfiguration has also been reported to be different in the outerand inner struts with the latter showing lower sp3/sp2 ratios [105].Strut shadowing is particularly accentuated when using PMS

discharges as the great majority of these platforms are builtwith planar magnetrons providing only unidirectional sputter-ing. When no rotation is applied to the stent, gradients in thethickness of the coating are to be expected as thicker layers aredeposited on the struts facing the magnetron. An enhancementof the coating uniformity could be achieved using, for instance,hollow cathode magnetrons (HCM) instead of planar ones.HCM discharges [106–108] use cylindrical targets that sur-round the samples to be coated. When the plasma is generatedthe sputtered material covers the outer layers of the samplesfacing the target. In addition, rotation can be applied tominimise shadowing of the inner struts. Using this technologycould, however, still face some possible drawbacks. Shadow-ing would not be totally eliminated since a mechanical supportto hold and rotate the stent inside the magnetron would still berequired. Additionally, it is known that coating properties, suchas thickness, are strictly dependent on the aspect-ratio (radiusto length ratio) of the magnetron [108]. Therefore, custom-made HCM reactors would need to be built to fit specific stentdimensions, a demand that could largely decrease the costefficiency and compromise the universality of the process.Stent design imposes a further obstacle with regards to its

surface characterization. An accurate determination of the bulk

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and the surface physical (roughness, wettability, etc.) andchemical (elemental composition, atomic bonding configura-tion, etc.) properties using the standard characterisation andspectroscopic techniques is difficult to achieve. The biocom-patibility of coatings deposited on stents is even morechallenging to assess in-vitro since it is hard to reproduceand quantify cellular attachment and proliferation on non-two-dimensional flat surfaces.

Stents are subjected to extreme radial and longitudinalforces communicated by an expandable catheter balloon duringdeployment. Localised plastic deformation follows stentexpansion and for this reason the coatings should be extremelyrobust to avoid cracking and spallation in the areas subjectedto high strains. Fig. 2 shows examples of carbon-based coatedSS stents that underwent balloon expansion. The coatings wereprepared by PECVD using acetylene as a precursor gas.Micro-sized cracking and delamination occurred in the regionswhere strain was more accentuated which is shown by thedevelopment of plastic deformation in the strut—Fig. 2a and b.Although spallation was not propagated throughout the strut,exposure of such coating to the blood stream could aggravatedelamination leading to further peeling. Fig. 2c–f show peelingand spallation examples of similar coatings following crimpand balloon expansion. Low adhesion strengths of the coatingsresulted in a generalized peeling and major areas of theunderlying stent material were exposed.

Given these reasons it is no surprise that the vast majority of theinvestigations available in the literature are merely preliminary,reporting the deposition and characterization of plasma-synthesisedcoatings on flat substrates. For instance, the adsorption of bloodproteins, the adhesion and activation of platelets and the attachmentand proliferation of smooth muscle cells and endothelial cells aretypically assessed on coatings deposited on Si-wafers and flatmetallic samples. Also, correlations between: (i) plasma depositionparameters, (ii) coating physical, chemical and mechanical proper-ties and ultimately (iii) coating biocompatibility are also performedin similar circumstances. However, one should expect that coatingsgrown on flat samples would perform differently from actualcoated stents in in-vitro and particularly in in-vivo. Only few worksaddressed these issues but not to their full extent.

3.2. Plasma-assisted coated stents

Polyethylene glycol (PEG) conjugated DLC coatings onnitinol flat samples revealed superior in-vitro hemocompatibilitypresenting higher albumin/fibrinogen adsorption rates anddecreased platelet adhesion when compared with bare metaland DLC surfaces [109]. However, PEG/DLC coated nitinolstent performance in canine iliac arteries was substantiallyinferior revealing more developed neointimal hyperplasia areascompared to DLC stents and even to control non-coated nitinolstents [110]. The authors associated restenosis with the accumu-lation of a thick collagen layer formed due to over proliferationof fibroblasts on the PEG surface, a factor that had beenoverlooked in the previous in-vitro assays.

Kim et al. [93] studied the mechanical properties of DLCcoatings deposited on self-expandable nitinol vascular stents using

a focused ion beam system and acetylene gas as a carbonprecursor. Interlayers of Si were also deposited by magnetronsputtering between the stent surface and the DLC layer forassessment of potential enhancement of the coating adhesionstrength. Stents were exposed to high stresses through a mechan-ism of stent contraction and expansion. DLC coatings depositedwithout Si interlayers were found not to withstand stress asdelamination and spallation occurred at the strut / coating inter-face. Cracking and delamination was prevented when a-Siinterlayer thicker than 0.6 nm was used. Unfortunately, there isno detailed information on the chemical, physical and biologicalcharacteristics of these DLC coatings. Similar results were foundelsewhere [61] on PECVD prepared DLC coated stainless steelstents. The addition of an a-Si:H interlayer substantially enhancedthe mechanical stability of the carbon coatings following stentexpansion. However, the adhesion of similar coatings to stainlesssteel flat substrates was significantly compromised after 1 monthof incubation in biofluids. Endothelial cell and platelet attachmentwas also studied on DLC coatings deposited on Si substrates.Superior endothelialisation and reduced platelet adhesion werefound for thermally annealed a-C:H films doped with Si.Unfortunately, these results were not confirmed on stents preparedin the same conditions.PAC coated 316L SS stents have been tested for both

mechanical stability and in-vivo performance [34,111]. The inclu-sion of nitrogen in acetylene/argon PECVD discharges providedimprovements in the mechanical properties by reducing thecompressive stress and in increasing the elasticity of the coatings.Superior ability for tropoelastin covalent immobilization on 316Lflat samples was observed for coatings prepared under suchconditions, which was congruent with an enhancement in endothe-lialisation. PAC coated directly onto stents prepared in suchcircumstances resisted delamination and spallation following aballoon expansion with only nano-sized cracks emerging butwithout exposing the underlying strut. Graded interfaces betweenthe metallic strut surface and PAC coatings can be made to furtherenhance PAC mechanical stability in animal trials [34]. AlthoughPAC alone did not present significant in-vivo improvements instent biocompatibility, further functionalization of PAC surfaceswith specific strongly bounded bioactive biomolecules couldovercome the limitations of current stent platforms.

4. Conclusions and future perspective

Carbon-based coatings, prepared either by plasma magne-tron sputtering (PMS) or by plasma enhanced chemical vapourdeposition (PECVD), are among the most promising plasma-based coatings for cardiovascular implants, particularly stents.These are generally characterised by a reduced internal stressand improved hemocompatibility. Recently, carbon-basedplasma polymers were reported to prevent the adverse responseto metallic substrates through linker-free immobilization ofbioactive molecules via free-radicals embedded in the coating.The radicals present in the coatings, prepared by PECVD withthe aid of high voltage bias, allow for the covalent binding ofvarious types of proteins. Because the proteins maintain theirbioactivity, these coatings can regulate favourable cellular

Fig. 2. (a) SEM image of a balloon expanded 316L stainless steel (SS) stent coated with a carbon thin film deposited by PECVD. Arrows indicate the areas ofaggravated plastic deformation. Strain accumulation in the superior region of the strut is indicated by the square; Scale bar denotes a length of 50 mm. (b) zoom offigure (a) showing micro-cracking and delamination of the carbon coating due to strain accumulation during stent expansion; Scale bar denotes a length of 1 mm.(c) As in figure (a) but now showing an example of coating peeling and spallation after balloon expansion due to inadequate coating adhesion strength. In this case,the underlying metal strut is exposed; scale bar denotes 65 mm. (e) Generalised peeling of the coating following stent crimp and balloon expansion. Major areas ofthe metal substrate are exposed in this example; scale bar denotes 15 mm. (f) Electron X-ray dispersive spectroscopy (EDS) mapping of figure (e) displaying a smallregion where the oxidized a-CN:H coating remained adhered to SS stent; scale bar denotes a length of 3 mm.

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responses and trick the body to recognise the implant as a non-foreign object. The one-step and linker-free covalent immobi-lization of a mixture of biomolecules capable of performingdifferent tasks independently would be extremely advanta-geous, fostering novel and totally drug-free biocompatiblestent platforms.

Despite the advances, considerable improvements must bemade for a successful commercialisation of biocompatible coatedstents using plasma discharges. The majority of the published

investigations in the field are still very preliminary. Plasma-basedcoatings are still frequently deposited and characterised on flatsamples such as Si wafers and/or metallic flat sheets. Thecomplex 3D geometry of a stent sets an additional challenge.Uniform coatings are hard to achieve due to shadowing effectseither caused by the stent itself or by the electrical connections ormounts used to hold the stent in the plasma chamber. Further-more, it is also expected that the physical, chemical andtribological properties of coatings deposited on flat samples

Stent surfaceCoating

Hybrid Tools

Plasma Chemistry

Surface Chemistry

Transport of species

and energy

Mass spectroscopy

Optical Spectroscopy

Infrared spectroscopy

Electrical probing

TOF SIMS XPS

XRD

FTIR

EDX

SEM

TEM

AFMEllipsometry

Thrombogenicity

Endothelialisation

Restenosis

In-vivo

In-vitro

Ex-vivo

C o a t i n g F u n c t i o n a l i z a t i o n

Hardness

Wear

Corrosion

Fig. 3. A proposed strategy for the development of a controllable and stable process for the production of biocompatible coatings for cardiovascular stents usingplasma discharges (red arrows indicate the flow of information). This strategy involves the development of plasma hybrid simulation tools fitted with appropriatemodels capable of describing the main physical quantities of the plasma, plasma chemistry and transport and plasma interaction with surfaces during the deposition.Model results need to be confirmed and validated with experimental measurements obtained by means of plasma diagnostics. Ultimately, such models would beused for a complete scaling and optimization of the plasma reactor. Coated stents must be tested for mechanical stability, wear and corrosion resistance. Also, thephysical and chemical properties of such coatings should be determined using appropriate characterization techniques. Coating functionalization through stablecovalent immobilization of bioactive molecules is then performed. These molecules (or cocktail of molecules) should be chosen to address thrombogenicity,endothelialisation and restenosis and their efficacy should be tested in in-vivo, ex-vivo and in-vitro. Iterative feedback between plasma modelling, plasmadiagnostics, coating characterization (before and after functionalization) and coating biological performance offer a strategy to identify the correct window of plasmaparameters for the production of coatings with the necessary mechanical, physical, chemical and biocompatible properties.

M. Santos et al. / Biosurface and Biotribology 1 (2015) 146–160156

would not be necessarily the same to those deposited on real 3Dsamples. Therefore, further in-vitro and in-vivo essays arecertainly needed to access plasma-coated stent performance.

Additionally, a full understanding of the fundamentalprocesses governing the plasma medium during the depositionprocess is far from being achieved. The use of reactivemixtures containing hydrocarbon precursors gives rise tocomplex plasma chemistries. The use of sample biasing, thethree-dimensional geometry of the samples and the addition ofreactive dopants such as nitrogen or hydrogen represent

additional challenges. The main kinetic mechanisms and thephysical quantities to describe such discharges are not yetclear, demanding therefore the development of advancedplasma modelling tools capable of describing the complexityof the plasma medium. Principally, a connection between theprocess parameters and their effects on the final coatingproperties are not yet elucidated. Not even for simplest gaseousmixtures. The mechanical stability of the coating, its capabilityto be further functionalised and its biocompatibility needs to bethoroughly investigated as a function of changes in the

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physical and chemical properties induced by varying appro-priate process parameters. Combinations of plasma dischargesimulation tools, plasma diagnostics and coating characteriza-tion techniques could be used to identify the relevant windowof plasma parameters for a reproducible and stable process (seeFig. 3) and ultimately offer feedback control strategies. Suchan interdisciplinary endeavour will only be possible throughcombined efforts of scientific communities in plasma physics,materials engineering, biology and medicine.

Acknowledgements

This work was supported by grants from the National Healthand Medical Research Council (APP1033079 andAPP1039072) and Australian Research Council. The authorsalso acknowledge the facilities as well as scientific andtechnical assistance at the Australian Centre for Microscopyand Microanalysis. We have no competing interests.

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