Progress in Materials Science - City University of Hong · PDF fileSurface functionalization of biomaterials by radical polymerization Tian Zhoua, Yizhou Zhua, Xia Lia, Xiangmei Liua,⇑,

Progress in Materials Science 83 (2016) 191–235

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Progress in Materials Science

journal homepage: www.elsevier .com/locate /pmatsc i

Surface functionalization of biomaterials by radicalpolymerization

http://dx.doi.org/10.1016/j.pmatsci.2016.04.0050079-6425/� 2016 Elsevier Ltd. All rights reserved.

⇑ Corresponding authors.E-mail addresses: [email protected] (X. Liu), [email protected], [email protected] (S. Wu).

Tian Zhou a, Yizhou Zhu a, Xia Li a, Xiangmei Liu a,⇑, Kelvin W.K. Yeung b, Shuilin Wu a,⇑,Xianbao Wang a, Zhenduo Cui c, Xianjin Yang c, Paul K. Chu d

aHubei Collaborative Innovation Center for Advanced Organic Chemical Materials, Ministry-of-Education Key Laboratory for the Green Preparation andApplication of Functional Materials, Hubei Key Laboratory of Polymer Materials, School of Materials Science & Engineering, Hubei University, Wuhan 430062, ChinabDivision of Spine Surgery, Department of Orthopaedics & Traumatology, Li Ka Shing Faculty of Medicine, The University of Hong Kong, 999077, Hong Kong, Chinac School of Materials Science and Engineering, Tianjin University, Tianjin 300072, ChinadDepartment of Physics & Materials Science, City University of Hong Kong, Tat Chee Avenue, Kowloon, Hong Kong, China

a r t i c l e i n f o a b s t r a c t

Article history:Received 28 December 2015Received in revised form 21 March 2016Accepted 23 April 2016Available online 23 April 2016

Keywords:Surface functionalizationPolymerizationSurface chemistryBiomaterialsTissue engineeringBiomimetic

One effective strategy in the field of biomaterials is to develop biomimetic interfaces tomodulate the cell behavior and promote tissue regeneration and surface modification isthe best way to obtain biomaterial surfaces with the desired biological functions andproperties. Surface radical polymerization offers many advantages compared to othermethods, for instance, low cost and simplicity, ability to control the surface chemistrywithout changing the properties of the bulk materials by introducing high-density graftchains and precisely controlling the location of the chains grafted to the surface, as wellas long-term chemical stability of the chains introduced by this method due to the covalentbonding. Because of the precise control of the macromolecules and easy preparation, con-trolled/living radical polymerization has been widely used to modify biomaterials. Thereare three main techniques: atom transfer radical polymerization (ATRP), nitroxide-mediated polymerization (NMP), and reversible radical addition-fragmentation chaintransfer (RAFT) polymerization. Some other grafting methods such as plasma-inducedpolymerization, irradiation-induced polymerization, and photo-induced polymerizationalso have great potential pertaining to functionalization of biomaterials and tailoring ofsurface chemistry. This paper summarizes recent advances in the various grafting polymer-ization methods to enhance the surface properties and biological functions of biomaterials.

� 2016 Elsevier Ltd. All rights reserved.

Contents

1. Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 192
1.1. Necessity for surface modification for biomaterials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1921.2. General background on grafting polymerization. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 192
1.2.1. ‘‘Grafting to” and ‘‘grafting from” approaches . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1921.2.2. Controlled/living radical polymerization. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1931.2.3. Brief introduction of other grafting methods . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 194

2. Nitroxide-mediated polymerization . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 194

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192 T. Zhou et al. / Progress in Materials Science 83 (2016) 191–235

2.1. Fundamentals of NMP. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1942.2. Surface grafting of biomaterials by SI-NMP. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 194

2.2.1. Biometals . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1952.2.2. Bioceramics . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1952.2.3. Biopolymers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1982.2.4. Others . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 198

2.3. ‘‘Grafting to” route . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2002.4. Factors influencing the properties of polymerized films by NMP . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 200

3. Atom transfer radical polymerization . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 201
3.1. Fundamentals of ATRP . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2013.2. Surface grafting of biomaterials by ATRP. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 201
3.2.1. Functionalized surface coatings . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2013.2.2. Factors influencing the properties of polymerized films by ATRP. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 206

4. Reversible addition-fragmentation chain transfer polymerization . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 206
4.1. Fundamentals of RAFT polymerization . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2074.2. Surface polymerization of biomaterials by RAFT. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 208
4.2.1. Monomers used in RAFT polymerization . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2084.2.2. Factors influencing the properties of polymerized films by RAFT . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 209

5. Other surface radical polymerization methods . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 211
5.1. Plasma-induced radical polymerization. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2115.2. Radiation-induced radical polymerization . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2135.3. Photo-induced radical polymerization . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 215
6. Biocompatibility of polymerized surfaces. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 223
6.1. Cytotoxicity . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2236.2. Hemocompatibility . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2236.3. Manipulation of cell behaviors . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 224
7. Conclusion and outlook . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 225Acknowledgements . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 226References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 226

1. Introduction

1.1. Necessity for surface modification for biomaterials

Biomaterials are widely used in biomedical devices or biological products in clinical biological diagnosis, treatment/repair, replacement of damaged tissues and organs, and enhancement of functions [1,2]. However, some interactionsbetween the biomaterials and biological tissues may undermine the ability of these materials to carry out their designed bio-logical functions sometimes even producing adverse effects. For instance, the contact between synthetic implants and bloodoften leads to protein adsorption, platelet adhesion, and concomitant activation of the body’s defense system, consequentlyinducing thrombosis and pulmonary thromboembolism [3–6]. Post-operative implant-related bacterial infection has beenreported to result in implant failure in 2%, 5% and 14% of implants for bone fracture fixation, total hip replacement, and kneereplacement, respectively [7–9]. Hence, a suitable surface design is crucial to more successful application of biomaterialsbecause surfaces are in direct contact with the biological media [10]. Before biological events take place in response tothe biomaterials, cells detect the surface affinity through the recognization of the filopodia of cellular transmembrane pro-teins. Many studies confirm that both the surface structure and surface chemistry influence the initial cell behavior such asadhesion, migration, proliferation, and differentiation on biomaterials [11–13].

It is important to perform surface functionalization to endow currently available biomaterials with specific and desirablefunctions. Currently, surface modification can be used to modulate the surface properties of substrates such as adhesion,wettability, biocompatibility, and antifouling [7–9,14]. Radical polymerization as one of surface modification methods is par-ticularly useful because it can easily and controllably introduce high-density graft chains and graft them in a precise mannerwithout affecting the original properties of substrates [15]. Moreover, covalent attachment of graft chains onto the surfacecan minimize chain delamination improving the long-term chemical stability of the introduced chains and it is possible tograft different polymers onto the same substrate [16].

1.2. General background on grafting polymerization

1.2.1. ‘‘Grafting to” and ‘‘grafting from” approachesPolymers can be grafted onto the surface to produce functional polymer brushes to tailor the surface properties and there

are two general ways to prepare polymer brushes: physisorption and covalent attachment [17–19]. Polymer physisorption isa reversible process in which the polymer chains with the sticking segments adsorb onto a suitable substrate but suchpolymer brushes are often unstable. Stamm et al. reported that poly(N-isopropylacrylamide) (NIPAM) brushes were

T. Zhou et al. / Progress in Materials Science 83 (2016) 191–235 193

immobilized with gold nanoparticles by physisorption [20] and covalent attachment could be accomplished by either the‘‘grafting to” or ‘‘grafting from” approaches [17,21].

In the ‘‘grafting to” approach, the polymer chains are bonded by the reactions between those preformed, end-functionalized polymers and the reactive groups on the substrate. Although this method was widely used in the early daysdue to the simplicity [22–25], there is a significant drawback that only a small amount of the polymer can be immobilized onthe surface due to steric constraints and kinetic factors [26,27]. The space constraints around potential reactive sites furtherlimit the grafting density and this barrier becomes greater when the layer thickness increases. Hence, the ‘‘grafting-to”method generally results in brushes with small grafting densities and film thicknesses [26].

The ‘‘grafting from” approach, also referred to as surface-initiated polymerization, utilizes the initiators bound to the sub-strate to initiate polymerization and the polymer evolves from the surface itself. This technique provides a viable alternativeto control the functionality, density, and thickness of the polymer brushes [18]. Owing to the effectiveness in producing thickand large-density functional polymer brushes in a controllable fashion [19], the ‘‘grafting from” approach has attracted muchinterest in recent years [26,28,29]. In fact, the ‘‘grafting from” and ‘‘grafting to” techniques can be implemented together[30]. For instance, Berger reported that the first polymer could be immobilized onto one side of silica particles using the‘‘grafting from” approach and the second polymer was grafted onto the other side of the particle surface via the ‘‘graftingto” approach [30].

1.2.2. Controlled/living radical polymerizationGrafting methods can be classified according to the type of radical bonding in traditional radical polymerization, ionic and

ring-opening polymerization, and controlled/living radical polymerization (CLRP). The major drawback of conventional rad-ical polymerization is the lack of control of the polymer structure [31,32], while ionic polymerization requires demandingconditions such as very low temperature without moisture. Furthermore, only a limited number of monomers can be usedto trigger undesirable side reactions due to the presence of functional groups in the monomers. In contrast, CLRP is a simpleprocess that produces well-defined, low-polydispersity, and multi-functional polymers [33–38] and the polymer architec-ture can be precisely controlled [39–42]. CLRP usually includes three categories: atom transfer radical polymerization(ATRP), nitroxide-mediated polymerization (NMP), and reversible radical addition-fragmentation chain transfer (RAFT)polymerization, which will be discussed individually in details later in this review.

1.2.2.1. Fundamentals of CLRP. The general reaction in CLRP is illustrated in Fig. 1 [41]. The key feature of all CLRP systems isthe establishment of a dynamic equilibrium between propagating radicals P� and various dormant species throughout thepolymerization process to minimize the occurrence of irreversible termination reactions [43]. The dormant (end-capped)chain P–X can be activated to produce the polymer radical P� by thermal, photochemical, and/or chemical stimuli. If themonomer M is present, P� will undergo propagation until it is deactivated back to P–X. Generally, [P�]/[P–X] should be lessthan 10�5 to ensure the success of the reaction, indicating that a living chain spends most of the polymerization time in thedormant state.

1.2.2.2. Architecture of polymers grafted onto the surface. When the surface of biomaterials is grafted with polymers by CLRP,the polymeric architecture is very important. According to the topography of the polymers, the polymeric structures includestar polymers, comb polymers, hyperbranched polymers, network and cyclic polymers. Star polymers possess lower bulk andsolution viscosities as well as many chain-end functionalities compared to linear polymers. Star polymers are usually pre-pared by two different routes: the ‘‘core-first” approach and ‘‘arm-first” approach that are suitable for ATRP, NMP and RAFTpolymerization [40]. Since the ‘‘core-first” approach employs multifunctional initiators from which several arms are grownsimultaneously, ATRP is more useful for polyols that can be converted to an initiating core with more initiating sites [44]. Inthe ‘‘arm-first” approach, the polymer chains are attached to a functional core. RAFT polymerization offers a uniquepossibility since the RAFT reagents can be attached to a core via the Z group [45].

Comb polymers have a main polymer chain and side chains with the same chemical nature. Yu et al. prepared well-defined comb copolymer brushes on poly(tetrafluoroethylene) (PTFE) films by surface-initiated CLRP [46]. The poly(glycidylmethacrylate) (PGMA) brushes were synthesized on the pretreated PTFE film by surface-initiated RAFT polymerization andthe hydrophilic vinyl monomers sodium salt of 4-styrenesulfonic acid (NaSS) was polymerized on the epoxy groups sidechains of the PGMA brushes by surface-initiated ATRP.

Hyperbranched polymers are polymers with a dense and branched structure and a large number of end-groups, which arealso called dendritic polymers. Mori and co-workers prepared hyperbranched polymers from silicon by self-condensing ATRP[47,48] and one-pot hyperbranched polymers was synthesized by RAFT [49].

Fig. 1. The mechanism of controlled/living radical polymerization (CLRP) [41]. Copyright 2007. Adapted with permission from Elsevier Science Ltd.


Polymer networks have attracted significant interest due to the well-defined mechanical and physical properties [50,51].Microgel networks can be formed by preparing well-defined polymers with crosslinkable pendant moieties. Oh et al. pre-pared cross-linked nanogel with a uniformly cross-linked network utilizing ATRP and it was possible to use crosslinkerswhich could be reversibly cleaved to form reversible gels [52]. The traditional way to prepare cyclic structures is basedon intramolecular cyclization of a linear precursor of the cyclic polymer and either two identical or two different functionalgroups react with each other. For example, Lepoittevin et al. prepared macrocyclic polystyrenes with controlled dimensionsand narrow distributions via CLRP [53].

1.2.3. Brief introduction of other grafting methodsIn addition to the aforementioned surface grafting methods, three polymerization techniques have been developed to

improve the surface properties of biomaterials, plasma-induced graft polymerization, radiation-induced graft polymeriza-tion and photo-induced graft polymerization. For example, biomedical stainless steel, titanium alloy, polystyrene (PS),polycaprolactone (PCL), polyethylene (PE), polypropylene (PP), and polymethyl methacrylate (PMMA) are often modifiedby these grafting methods.

2. Nitroxide-mediated polymerization

NMP is the simplest CLRP technique by just adding free nitroxides in the conventional radical polymerization process. Incomparison with other CLRP methods, NMP is purely thermal activation that does not require any catalyst or bimolecularexchange. In addition, it does not require purification after polymerization and aqueous processes can be implementedbecause the reagents in NMP are typically environmentally friendly [32,54].

2.1. Fundamentals of NMP

NMP involves a reversible activation–deactivation equilibrium in which nitroxide acting as a control agent reversiblydeactivates the propagating and growing radicals to an alkoxyamine dormant end-functionality as the predominant species(Fig. 2) [55]. Whenever this equilibrium is turned to the dormant form, the irreversible chain termination is limited and thestationary concentration of the active species is small.

NMP can be initiated in two different ways. The first is the bicomponent initiating system consisting of a conventionalthermal initiator (e.g., 2,20-azobisisobutyronitrile (AIBN) or benzoyl peroxide (BPO)) and a stable free nitroxide (e.g., 2,2,6,6-tetramethylpiperidinyloxy (TEMPO)) [56]. When these systems are used, conventional radical polymerization can beemployed in the NMP process by simply adding free nitroxides. However, this system cannot accurately define the structureand concentration of the initiating species. In contrast, in the monocomponent initiating system, the well-defined unimolec-ular initiator can release both the initiating radical and nitroxide with a 1/1 M ratio by thermolysis to circumvent thisproblem [57–59].

As shown in Fig. 2, it is predicted theoretically that the activation–deactivation equilibrium (equilibrium constant K = kd/kc) and persistent radical effect (PRE) control the polymerization kinetics in NMP [60]. Practically, side reactions significantlyaffect the polymerization rate in some cases. The first one involves the presence of an excess of free nitroxide when thereacted monomers (e.g., acrylates) have a high propagation rate constant (kp). It leads to rapid consumption of the initiatingalkoxyamine resulting in living polymer chains with poor control. The second factor is degradation of nitroxides. Nitroxidesare only highly effective at a relatively low temperature and degrade at temperature over 120 �C in NMP. The potential appli-cation of unstable nitroxides depends on the nature of the degradation products. The third way to influence the polymeriza-tion rate is to use additional long half-life radical initiators to decrease the concentration of the free nitroxide in the medium.The most complex side reaction in NMP is disproportionation or the hydrogen transfer reaction [61,62].

2.2. Surface grafting of biomaterials by SI-NMP

NMP can be used to modify biomaterials by introducing bioactive polymeric groups to the surface regardless of the natureof the substrates. Owing to the living nature of NMP, this method can be applied to surface-initiated polymerization and‘‘grafting to” processes [63]. The former requires a well-defined initiator and controlled chain growth, whereas the latterrequires the ability to end-functionalize the chains. In surface-initiated NMP (SI-NMP), binding of the initiator depends

Fig. 2. The mechanism of nitroxide-mediated polymerization (NMP) [55]. Copyright 2008. Adapted with permission from the American Chemical Society.


on the chemical nature of the inorganic surface and can take place via covalent bonding, electrostatic interaction, or hydro-gen bonding. In the second case, NMP-derived copolymers are grown on the surface of the inorganic substrate by covalentlybonding with reactive groups. However, on account of steric hindrance, the ‘‘grafting to” process generally produces inbrushes with small grafting density and film thickness. Hence, this method is not very popular.

2.2.1. BiometalsOwing to excellent mechanical properties such as high strength, toughness, wear resistance, and corrosion resistance,

metals are often used in biomedical devices and orthopedic implants [64]. Recently, biodegradable metals such as Mg-based alloys, Fe-based alloys, and Zn-based alloys are popular in tissue regeneration applications such as intravascular stentsand bone scaffolds [65]. However, the bioinert nature of some metals and antibacterial performance must be improved.

Ignatova and co-workers utilized a two-step ‘‘grafting-from” method to synthesize copolymer brushes on stainless steel.Electrografting of the poly(2-phenyl-2-(2,2,6,6-tetramethyl-piperidin-1-yloxy)-ethylacrylate) with a large grafting densitywas conduced as the initiator (first step) followed by nitroxide-mediated radical polymerization of 2-(dimethylaminoethyl)acrylate and styrene or n-butyl acrylate (second step) initiated by the electrografted polyacrylate chains [66]. The pro-cess is schematically illustrated in Fig. 3 [66]. The functionalized surface exhibited significant antibacterial activity againstthe Gram-positive bacteria Staphylococcus aureus (S. aureus) and Gram-negative bacteria Escherichia coli (E. coli) and theantibacterial activity decreased when the alkyl chain length increased. Furthermore, the thickness and hydrophilicity ofthe adherent copolymer films could be adjusted by SI-NMP [66]. A similar process was implemented in the synthesis ofcopolymer brushes on a steel surface on which acrylate was first electropolymerized under a cathodic potential and NMPof styrene was initiated from the electrografted polyacrylate chains with formation of polystyrene with a controlledmolecular weight and narrow polydispersity [67].

2.2.2. BioceramicsOwing to the excellent biocompatibility, bioceramics are popular in tissue repair, maintenance, and restoration to

improve the functions of organs and tissues [68]. For example, the excellent properties of silica nanoparticles such as thehigh intrinsic tunable porosity, hydroxyl surface functionality, and specific surface area make the materials suitable forNMP [69], especially SI-NMP in applications including drug delivery, drug discovery and biosensing [68,70–73].

Bartholome and co-workers prepared PS-grafted silica particles by simple condensation and reacted these particles with anitroxide/conventional radical initiator (e.g., AIBN or BPO) bicomponent system [74]. The modified silica particles exhibitedbetter colloidal stability and enhanced dispersibility in toluene. Chevigny et al. proposed a new convenient and efficient SI-NMP method using a covalent grafting strategy involving an alkoxyamine which acted both as the initiator and controlleragent [75]. The synthesis parameters were optimized to obtain good conversion rates and expected molecular weight ofthe grafted chains while maintaining the colloidal stability and avoiding aggregation of silica particles. The synthetic andcharacterization methods represented a robust and reproducible way to design well-defined grafted polymer nanoparticles.Subsequently, this group synthesized well-defined PS-grafted silica nanoparticles by adapting their previous synthetic pro-cess without using a free initiator to tune the grafted chains while controlling polymerization and colloidal stability whileavoiding the formation of free polymer chains [76].

Recently, Zhao et al. prepared poly(tert-butyl acrylate) (PtBA)/PS brushes on silica nanoparticles by sequential ATRP oftert-butyl acrylate (tBA) at 75 �C and NMP of styrene at 120 �C [77–80]. Subsequent hydrolysis of the tert-butyl groupsproduced amphiphilic mixed poly(acrylic acid) (PAA)/PS brushes that underwent chain reorganization in response toenvironmental changes [77]. The mixed PtBA/PS brushes exhibited two distinct glass transition temperatures, suggestingthat the two grafted polymers were microphase-separated in the brush layer (Fig. 4) [78,79]. The grafting density increasedfrom 0.6–0.7 chains nm�2 to 0.9–1.2 chains nm�2 when a triethoxysilane anchoring group was used instead of achlorodimethylsilane group [79]. Later, this group studied the effects of the overall grafting density on the microphase sep-aration of PtBA/PS brushes [80]. A series of mixed PtBA/PS brushes with different overall grafting densities but similar molec-ular weights and comparable individual grafting densities for the two polymers was prepared by changing the mass ratio ofthe Y-initiator to silica particles in the initiator immobilization step (Fig. 5) [80]. When the grafting density was further

Fig. 3. Two-step ‘‘grafting from” strategy by electrochemistry [66]. Copyright 2004. Adapted with permission from the American Chemical Society.

Fig. 4. Schematic illustration of phase morphologies of mixed PtBA/PS brushes with a fixed PtBA molecular weight and varying PS molecular weight [78].Copyright 2010. Adapted with permission from the American Chemical Society.

Fig. 5. Synthesis of mixed homopolymer brushes with various overall grafting densities by changing the mass ratio of Y-initiator to silica particles in theinitiator immobilization step via an ammonia-catalyzed hydrolysis/condensation process [80]. Copyright 2012. Adapted with permission from theAmerican Chemical Society.


decreased, phase separation became weaker and no microphase separation was observed from the sample with an overallgrafting density of 0.122 chain/nm2. The properties of the materials were related to microphase separation. Recent researchhas demonstrated that mesoporous silica particles can be used as NMP substrates [81]. Although SI-NMP of styrene was per-formed from various types of ordered mesoporous silica particles with different morphologies and pore sizes, the pore diam-eter was critical to diffusion of the chemical agents. For instance, a diameter of 2 nm was too small to ensure good diffusionof the reagents and 5 nm was adequate. The pore morphology and connectivity were crucial to diffusion of the chemicalagents when the pores were larger than 5 nm [81].

TiO2-based nanostructured materials play important roles in many biomedical applications such as tissue reconstructionand disease diagnosis due to their excellent biocompatibility, good chemical stability, high catalytic efficiency, excellentmechanical properties, and high strength-to-weight ratio [82]. Some properties can be enhanced by surface graftinghigh-density PS brushes onto the surfaces by SI-NMP [83,84]. For example, PS-grafted-TiO2 hybrid films exhibited high trans-mittance in the visible light region together with the ultraviolet absorption characteristics of TiO2 because PS-grafted-TiO2

nanoparticles formed a stable dispersion in good solvents for the PS chains [83]. In addition, the PS hybrid films withPS-grafted-TiO2 showed enhanced wear resistance compared to the PS film, possibly because the PS-grafted-TiO2

nanoparticles at the interface between the silicon wafer surface and PS thin film withstood the load [83,84].


Fe3O4 nanoparticles were grafted with PS and poly(3-vinylpyridine) (P3VP) by SI-NMP and the resulting PS- andP3VP-grafted magnetite particles were stably dispersed in the appropriate solvents [85–87]. Binder et al. prepared core/shellc-Fe2O3 nanoparticles with a well-defined poly(N-isopropylacrylamides) (PNIPAM)-shell by SI-NMP (Fig. 6) [88]. Robbeset al. presented a new multi-step efficient ‘‘grafting from” method to obtain well-defined magnetic c-Fe2O3 nanoparticlesgrafted with polystyrene (PS) chains [89]. The PS-grafted magnetic nanoparticles are of interest due to applications suchas drug delivery and targeting carrier systems.

The NMP technique can be used to functionalize the surfaces of carbon nanotubes (CNTs) which exhibit remarkable elec-tronic and mechanical properties [90–92]. Datsyuk et al. reported in situ NMP of methyl acrylate (MA) on double-walledCNTs (DWCNTs) to obtain the DWCNT-PAA-PMA or DWCNT-PS-PMA composites [93]. The main advantage of this two-step synthetic route is that it does not involve any pre-treatment or functionalization of the CNTs. However, Dehonoret al. reported direct free-radical functionalization of CNTs with BPO/TEMPO followed by SI-NMP of styrene without an acidtreatment. The peroxide group was first attached to the CNT surface to create a resonant radical that could trap nitroxide[94]. In situ polymerization of PS-grafted CNTs by NMP did not induce extensive damage on the CNT surface [94] and theacid treatment was very useful in surface functionalization. Zhao et al. used the acid treatment to introduce carboxyl groupsto the surface of multi-walled carbon nanotubes (MWNTs) (Fig. 7) to enable further reaction with TEMPO-basedalkoxyamine bearing a hydroxyl group to prepare a MWNT-supported initiator [95,96]. The initiator was used to obtainMWNT-PS [95] or MWNT-poly(4-vinylpyridine) (P4VP) [96], respectively via SI-NMP. The MWNT-PS showed relatively gooddispersibility in various organic solvents and the MWNT-P4VP exhibited good dispersibility in acidic aqueous solutions asthe dispersibility of MWNTs was a reflection of the solubility of the attached polymer [95,96]. The MWNT-PS-P4VP showeddifferent dispersibility characteristics than the other two samples. Water-soluble MWNTs are expected to have applicationsin supramolecular chemistry and biochemistry.

Fig. 6. SI-NMP of NIPAM from the surface of c-Fe2O3 nanoparticles with attached TIPNO-based alkoxyamine [88]. Copyright 2007. Adapted withpermission from the American Chemical Society.

Fig. 7. Synthesis of the MWNT-supported alkoxyamine initiator [96]. Copyright 2006. Adapted with permission from the Royal Society of Chemistry.


2.2.3. BiopolymersCommon biopolymers include polysaccharides, polypeptides, proteins, polyethylene, polycaprolactone and others and

NMP has been applied to polymer-protein/peptide bioconjugates [97]. For example, Chenal designed a protein/peptidePEGylation system based on functional comb-shaped polymethacrylates with poly (ethylene glycol) (PEG) side chains[98]. Möller and co-workers obtained peptide-polymer conjugates with well-defined linear peptide backbones and multiplenumbers of PS or poly PNIPAM side chains via NMP [99]. Because the composition, number, and length of the polymer sidechains vary, the conjugates aggregate to different structures. For example, peptide-PS conjugates may aggregate tohoneycomb structures, whereas peptide-PNIPAM conjugates show a differentiated aggregation behavior. As a naturalpolysaccharide, chitosan has been grafted by NMP to obtain functionalized copolymers with controlled molecular weightsand ‘‘well-defined” structures and good biocompatibility, biodegradability, bioactivity, as well as antibacterial andwound-healing activity have been observed from grafted chitosan [100–102].

2.2.4. OthersBesides the aformentioned traditional biomaterials, other biomaterials can be modified by this method [103–117]. Zhao

prepared mixed polymer brushes on the silicon wafer using a combination of ATRP and NMP (Fig. 8) [103–106]. Owing to theunique structure, these mixed homopolymer brushes exhibited reversible changes in the surface properties in differentsolvents [103,106] and even formed relatively ordered nanoscale domains [105–106]. As shown in Fig. 9 [106,112], exposing

Fig. 8. Synthesis of mixed PMMA/PS brushes from an asymmetric difunctional initiator-terminated SAM (Y-SAM) by combining ATRP and NMP techniques[106]. Copyright 2004. Adapted with permission from the American Chemical Society.

Fig. 9. Solvent-dependent reorganization of the binary polymer brushes [112]. Copyright 2009. Adapted with permission from Wiley-VCH Verlag GmbH &Co. KGaA, Weinheim.


the wafer to the polar solvent glacial acetic acid induced the polar polymer chains of PMMA to move to the outermost layer,while the apolar polymer brushes of PS were buried in the polymer film. The non-selective solvent CHCl3 yielded a statisticalmixture of polar and apolar polymer brushes at the surface. In contrast, in the apolar solvent cyclohexane, the polar PMMAbrushes aggregated in the film while the apolar PS brushes stretched to the outside.

Studer and co-workers demonstrated that PS brushes prepared by SI-NMP on silicon wafers possessed either proteinadsorption or protein repellent properties depending on the thickness of the PS brush [113,116,117]. They found that PSbrushes with thicknesses larger than 45 nm exhibited a surprising degree of protein repellence, whereas protein adsorptionwas observed from PS brushes with thicknesses below 20 nm [116]. PNIPAM brushes have been reported to have goodprotein resistance at room temperature [118]. Cimen and Caykara modified the surface of silicon wafers by grafting biofunc-tional PNIPAM with a high density of amine groups to immobilize a range of antigens or ligands in biosensing applicationsand to study cell-surface interactions. The process of SI-NMP of NIPAM is depicted in Fig. 10 [115]. PEG and PEG-basedmaterials have been used to modify biomaterials by NMP because of the ability of PEG to resist protein and cell adhesionin addition to its non-toxicity and non-immunogenicity [98,107]. By using the same method, poly(carboxybetaine methacry-late) (PCBMA) was grafted onto a silicon wafer to obtain polymer brushes that were able to suppress protein adsorption,while the terminal phosphonate group might promote endothelial adhesion [114].

Owing to their surface chemistry, outstanding optical properties and target specificity, quantum dots (QDs) such as CdSe,CdS, and CdTe are often used in immunolabeling, cell tracking, and in vivo imaging [119–121]. However, the processability,shelf-life stability, and relevant applications of these QDs are usually determined by the nature of the ligand periphery. Theapplicability ultimately depends on the performance of these specific platforms. Sill et al. demonstrated the growth of PS andP(styrene-co-methyl methacrylate (MMA)) copolymers on CdSe nanoparticles by NMP (Fig. 11) [122]. While free radicals canquench the fluorescence of CdSe nanoparticles, NMP allowed fluorescence of the prepared composites to be maintained dueto the low concentration of radicals inherent to this CLRP technique.

In summery, almost all biomaterials including biometals, biopolymers, bioceramics, and nanomaterials can be graftedwith various types of bioactive polymers by NMP to achieve specific biological functions such as hydrophilicity/hydrophobicity, antibacterial performance, bioactivity, and so on (Table 1) [66,67,74–81,83–88,93–96,98,99,101–117,122].

Fig. 10. Surface-initiated-NMP of NIPAM in the presence of CA [115]. Copyright 2012. Adapted with permission from the Royal Society of Chemistry.

Fig. 11. Ligand exchange and polymerization of styrene from nitroxide-functionalized CdSe quantum dots [122]. Copyright 2004. Adapted with permissionfrom the American Chemical Society.

Table 1Typical biomaterial grafted with bioactive polymers through NMP.

Substrate Grafted polymer Ref.

Biometals Stainless steel Poly(n-butyl acrylate) [66,67]

Bioceramics Silica PS [74–81]Titanium oxide (TiO2) nanoparticles PS [83,84]Fe3O4 nanoparticles P3VP [85]Fe2O3 nanoparticles PNIPAM [88]Carbon nanotubes PAA, P4VP [93–96]

Biopolymers Peptide/protein PEGMA [98]Peptide PNIPAM [99]Chitosan PMMA, PSS [101,102]

Others Silicon wafers PEG, PNIPAM, PCBMA [103–117]Quantum dots PS [122]


2.3. ‘‘Grafting to” route

In some cases, new biomaterials like carbon nanotubes can be modified by SI-NMP method by adopting the ‘‘grafting to”approach because a covalent carbon–carbon linkage can be formed when the surface reacts with the radicals produced bythermolysis of the nitroxide-end capped polymer chains. The well-defined poly(2-vinylpyridine) (P2VP) was synthesizedon MWNTs and the resulting materials were easily dispersed in organic solvents as well as acidic water too immobilizethe metal nanoclusters on the surface [91]. P4VP, PAA, and PS could also be grafted onto MWNTs by this technique[123,124]. Similarly, Charleux and co-workers employed this process to graft a variety of polymer chains onto primary-amine functionalized silica nanoparticles under soft conditions and the polymer grafting densities were 0.1–0.2 chain nm�2

[125]. This approach has also often been employed to modify the silicon wafers. A convenient route is to link the NMP-derived polymer by the ‘‘grafting to” process to the silicon wafer, followed by end-grafting the corresponding polymer tothe silanol groups on the surface through the hydroxyl group of 4-hydroxy-TEMPO. Drockenmuller et al. reported anotherapproach for grafting polymer brushes onto passivated silicon surfaces and well-defined PEG, PMMA, and PS brushes witha thickness of 6 nm and density of 0.2 chains nm�2 were prepared [126].

2.4. Factors influencing the properties of polymerized films by NMP

The properties of the polymer brushes such as the grafting density, molecular weight, molecular weight distribution,composition, as well as number and length of polymer chains play important roles in the performance of the biomaterials.The grafting density of the polymer brushes is related to microphase separation. Bao and co-workers studied the effects ofthe grafting density on microphase separation of PtBA/PS brushes by changing the grafting densities but maintaining similarmolecular weights and individual grafting densities for two polymers [80]. Microphase separation became weaker as thegrafting density decreased and no microphase separation was observed when the grafting density was sufficiently small.

The molecular weight of the polymer brushes not only changes the colloidal stability and dispersibility of the biomate-rials, but also influences the solubility of biomaterials in addition to the grafting density of the polymer. Bartholome et al.observed that the colloidal stability and dispersibility of silica particles in toluene was significantly improved when themolecular weight of the grafted PS chains was larger than 20,000 by focusing on alkoxyamine initiators of which the graftingdensity was proportional to the amount of grafted polymer chains [74]. Homenick and co-workers studied the carbon nan-otube solubility by polymer grafting and reported that neither a large polymer grafting density nor polymer alone producedthe maximum solubility, whereas maximum solubility was achieved with a polymer with an intermediate molecular weighthaving a relatively large grafting density [127].

Owing to the difference in the composition, number, and length of the polymer side chains, the conjugates aggregate todifferent structures. The aggregation behavior of peptide-polymer conjugates was studied by Möller et al. [99]. Linearconjugates could formmuch larger aggregates and show sharp phase transitions, but comb type conjugates could precipitateinto particles with a limited size. When the peptide side chains were short and the compositions of the peptide side chainswere different, the peptide-PS conjugates aggregated to the honeycomb structures, whereas peptide-PNIPAM conjugatesaggregated to the ring-shape topology. When the composition of the peptide side chains was the same and the peptide sidechains were different, peptide-PS conjugates bearing small side polymer chains aggregated to the honeycomb structureswith a coral like morphology for systems with longer side chains. Peptide-PNIPAM conjugates showed a different aggrega-tion behavior with the different length of the peptide side chains [99].

The thickness of the polymer brushes affects protein adsorption and repelling [113,116,117,128]. The PS brushesexhibited a surprising degree of protein repellence when the PS brushes were larger than 40 nm, whereas brushes smallerthan 20 nm showed a protein adsorption behavior [113]. Halperin also demonstrated that secondary protein adsorption wasrepressed by increasing the brush thickness while primary protein adsorption was repressed by increasing the graftingthickness [128].


Although NMP has been widely used to modify biomaterials, few studies have been carried out to explore the effects ofthe modification on biological functions. Most studies have focused on surface properties such as antibacterial characteristicsor antifouling, hydrophilicity or hydrophobicity, surface potentials, and surface biocompatibility. The in vitro and in vivo bio-logical reactions between NMP-functionalized biomaterials and cells or tissues should be studied in the near future.

3. Atom transfer radical polymerization

In comparison with other CLRP techniques, ATRP is particularly attractive because this polymerization process can beeasily controlled, enabling the achievement of a low polydispersity index and eliminating the need for involved purificationpost-synthesis [129–131]. Surface-initiated ATRP is more accessible than NMP because it is easier to synthesize thiol andsilane-derivatized surface-bound initiators than the AIBN-silane derivative or the nitroxide silane derivative in NMP.

3.1. Fundamentals of ATRP

Control of radical polymerization depends on two principles [130]. First of all, initiation should be rapid to provide a con-stant concentration of the growing polymer chains. Secondly, the majority of the polymer chains are dormant so that theycan retain the ability to grow based on the dynamic equilibrium between the dormant species and growing radicals. ATRPfollows these principles by involving a transition metal halide and a suitable ligand [132] and the mechanism is described inFig. 12 [133].

Due to the complex catalyst, the equilibrium between the active and dormant species is reversible. In ATRP, the reactionbetween the alkyl halide initiators or dormant species (RX or Pn–X) and activator-metal complexes Mtn/L with a low oxida-tion state, where Mtn represents the transition metal species with the oxidation state n and L is a ligand, reversibly generatesthe propagating radicals (R�/Pn�) and deactivator-higher-oxidation-state metal complexes X–Mtn+1–Y/L. These radicals initi-ate polymerization by forming a double bond with the monomer, propagate, terminate by either coupling or disproportion-ation, or are reversibly deactivated by the transition metal complex with the higher oxidation state [134,135].

When the concentration of the propagating radicals is sufficiently small compared to the dormant species, the proportionof chains terminated by bimolecular recombination can often be neglected (<5%). The small concentrations of propagatingradicals suppresses termination and enables the production of highly functional polymers (>95%). Thus, the concentrationof active species or propagating radicals remains at a very low level throughout polymerization. As ATRP is a catalyticprocess, it can be mediated by redox-active transition metal complexes such as Cu, Fe, Mo, Os, and Ru [130,132,136–155]. However, the rate of ATRP does not depend on the absolute catalyst concentration, but rather the ratio of the activatorto deactivator concentration. This is because the rate constants and the ratio (KATRP = kact/kdeact) determine the radical con-centration and rates of polymerization and termination as well as polydispersity. The key to the success of ATRP is to employa proper catalytic systemwith high activity and selectivity. Recently, initiation systems have been developed to utilize a verysmall amount of active catalyst and to constantly convert the deactivator to an activator by a redox process.

Compared to other radical grafting methods, ATRP minimizes radical–radical coupling at the surface due to site-specificinitiation and ATRP equilibrium, which reduces the radical concentration and termination.

3.2. Surface grafting of biomaterials by ATRP

Because of the controlled nature of the reversible activation–deactivation reaction between the growing polymer chainand a copper-ligand species, ATRP is compatible with various types of monomers. Some monomers with special functionalgroups are often grafted onto substrates such as PMMA, PS, and others by ATRP to achieve specific antifouling, antibacterial,and stimuli-responsive properties (Table 2) [156–242].

3.2.1. Functionalized surface coatings3.2.1.1. Antifouling coatings. Biomaterials with antifouling coatings are resistant to cell adhesion and non-specific proteinadsorption. These are important characteristics to prevent undesirable events such as platelet adhesion, thrombus formation,foreign body reaction, bacterial infection, and adhesion of macrophages [156–160]. Antifouling surfaces are typicallyhydrophilic and prevent binding with many hydrophobic foulants including non-polar solutes, proteins and bacteria

Fig. 12. The mechanism of transition metal-catalyzed ATRP [133]. Copyright 2014. Adapted with permission from Elsevier Science Ltd.

Table 2Bioactive surface prepared via surface-initiated ATRP.

Nature of polymericcoating

Grafted monomers Substrate Ref.

Antifouling coatings PEG-containingmethacrylate

Silicon, gold, stainless steel, silica, Ti, magnetitenanoparticles

[158–160,162,165,172–177]

MPC Silicon, gold [156,157,169,178]SBMA Stainless steel, gold, poly(propylene oxide) [179–182,184]CBMA Gold, glass [180,183,184]HEMA Ti, silicon, protein, peptides [185–189]AAm Silicon, gold, chitosan [190–192]NVP Silicon [193]

Antibacterial coatings DMAEMA PP, silicon [196–199]4VP Stainless steel, microporous polysulfone [200,201]SPMA Gold, Si/SiO2 [202]TBAEMA PP, LDPE, stainless steel [203–205]

Stimuli-responsivecoatings

NIPAM Silicon, PS, gold nanorods, silica [158,215–219]AA Cellulose membrane, porous nylon membrane, silicon [220–222]4VP Gold nanoparticles [224]DMAEMA Silica, gold, PS, glass conical nanopore channel [225–230]

Others MMA Silicon, magnetite nanoparticles, multiwalled carbonnanotubes

[237–239]

Styrene Chitosan, silica, multiwalled carbon nanotubes [240–242]


[161]. Biomaterials can be endowed with antifouling properties by ATRP with PEG-containing polymers, zwitterionic group-containing polymers, and other antifouling polymers.

3.2.1.1.1. PEG-containing polymers. PEG-based materials have emerged as common antifouling materials due to the goodantifouling effects against a wide variety of proteins, suppression of platelet adhesion, and reduction of cell attachmentand growth [162–165]. Dense ‘non-fouling’ polymer brushes have been synthesized by ATRP of various PEG macromono-mers on gold [159,160], silica [158,162,165,172,175], Ti [176,177], stainless steel [177], and magnetic nanoparticles (MNPs)[173,174]. Chilkoti reported that poly(oligo(ethylene glycol) methacrylate) (polyOEGMA) brushes grafted on gold substratesby ATRP prevented non-specific cell adhesion for up to 30 days [159]. Ti substrates coated with PEG brushes approximately100 nm thickness were shown to have excellent resistance against cell adhesion for up to 3 weeks [176] and PEGylated MNPsresisted non-specific protein adsorption better than the pristine nanoparticles [173]. After PEGylation, the uptake of MNPs bythe macrophage cells was reduced to 2 pg/cell while the pristine MNPs were taken up very effectively (Fig. 13) [174]. Themorphology and viability of the macrophage cells cultured in a medium containing the PEGylated MNPs were similar tothose of cells in the control experiment without the nanoparticles (Fig. 13), indicating that this amount of PEGylatedMNP uptake did not induce significant cytotoxicity. However, PEG being a polyether autoxidizes readily in the presenceof oxygen and transition metal ions often resulting in the cleavage of the ethylene oxide units and formation of reactivealdehyde moieties [166–168].

3.2.1.1.2. Zwitterionic group-containing polymers. Zwitterionic surfaces resist non-specific protein adsorption via a hydratedlayer formed by solvation of the charged terminal groups by electrostatic interactions in addition to hydrogen bonding[169–171]. Zwitterionic monomers undergoing ATRP include 2-methacryloyloxyethyl phosphorylcholine (MPC)[156,157,169,178], sulfobetaine methacrylate (SBMA) [179–182,184], and carboxybetaine methacrylate [180,183,184], withMPC being more difficult to synthesize and handle. Recently, PCBMA-grafted surfaces have been shown to have improvedresistance against non-specific protein adsorption in human serum and plasma compared to polyPEGMA- and polySBMA-grafted surfaces, indicating that zwitterionic polymers prepared by ATRP are a viable alternative to PEG-based materialsenabling a variety of applications such as diagnostics and drug delivery [184].

Dopamine and organosilane are often used as a bridge to yield a better combination between the substrate and the poly-mer, while self-assembled monolayers (SAMs) are usually used to synthesize non-fouling coatings with an appropriate graftdensity due to the large surface density and ease of preparation. Sin and co-worker compared the effectiveness of polySBMAbrushes grown from dopamine and silane-assembly layers on stainless steel (SUS) by ATRP in resisting adhesion of proteins,cells, and bacteria [179]. As shown in Fig. 14, both the dopamine and organosilane produced substrates with active aminogroups and allowed the subsequent grafting reactions to continue. As shown in Fig. 15a, fibrinogen from the blood plasmawas used as a model protein to assess the ability of a non-bioadhesive surface to resist protein adsorption, with polystyreneused as a reference polymer due to its 100% relative fibrinogen adsorption. Compared to SUS (approximately 85%), fibrinogenadsorption of SUS-D-pSBMA (1%) was much less than that of SUS-Si-pSBMA (8%). The relative cell adhesion density of SUS-D-pSBMA was also smaller than that of SUS-Si-pSBMA in both MG63 and HT1080 as shown in Fig. 15b, which presents astatistical analysis of relative cell adhesion on the substrate surface. As shown in Fig. 15c, E. coli and Staphylococcusepidermidis (S. epidermidis) commonly related to infection on orthopedic implants were utilized to evaluate the antibacterialactivity [194]. The SUS-D-pSBMA surface resisted bacterial adhesion with only 0.8% E. coli adhesion and 0.02% S. epidermidis

Fig. 13. (a) RAW 264.7 cells in control culture (without any nanoparticles) after 1 day, (b) cells after culturing in medium containing pristine magneticnanoparticles (0.2 mg/mL) for 1 day and (c) for 4 days, and (d) cells after culturing in medium containing P(PEGMA)-immobilized nanoparticles (0.2 mg/mL)for 1 day. The P(PEGMA)-immobilized nanoparticles were obtained after polymerization time of 2 h. Scale bar = 40 lm [174]. Copyright 2006. Adapted withpermission from the American Chemical Society.

Fig. 14. Schematic illustration of the preparation process of zwitterionic polySBMA grafting from stainless steel via ATRP using (a) catechol dopamine and(b) organosilane as respective self-assembly anchoring agent [179]. Copyright 2014. Adapted with permission from the American Chemical Society.


adhesion, whereas significant amounts of bacteria adhered to the SUS-Si-pSBMA surface (1.2% E. coli adhesion and 1.5% S.epidermidis adhesion, respectively). These results showed that the polySBMA grafted from dopamine had better adhesionresistance than polySBMA from silane.

3.2.1.1.3. Other antifouling polymers. In addition to PEG and zwitterionic polymers, some polymers such as poly(2-hydroxyethyl methacrylate) (PHEMA) [185–189], poly(acrylamide) (PAAm) [190–192] and poly(N-vinyl-2-pyrrolidone)(PNVP) [193], also possess excellent antifouling properties. Dense hydrophilic PHEMA brushes exhibited excellent proteinrepellency and well-defined antifouling PHEMA brushes grafted by ATRP also prevented cell adhesion. Consequently,surfaces with a well-defined density gradient of PHEMA brushes could be prepared to tailor cell adhesion [189]. Graftinga densely packed and covalently attached PAAm brush layer from a silicon wafer by ATRP reduced the attractive forcebetween the surface and microorganisms and this approach has great potential in mitigating infection on implants [191].Ultra-low fouling surfaces were prepared by grafting PAAm brushes on gold surfaces by surface-initiated ATRP [190]. The

Fig. 15. Statistical analysis of (a) relative fibrinogen adsorption and water contact angle, (b) adherent human MG63 osteoblast and HT1080 fibroblast cellsdensity, and (c) percentage of occupied area of E. coli and S. epidermidis attachment on various surface-modified stainless steel coupons [179]. Copyright2014. Adapted with permission from the American Chemical Society.


PAAm-grafted surfaces resisted protein adsorption and adhesion of bovine aortic endothelial cells (BAECs) and inhibitedattachment of S. epidermidis and P. aeruginosa.

3.2.1.2. Antibacterial coatings. Infection is a great concern for implants and medical devices resulting in orbidity, mortality,and medical costs [194]. To obtain antimicrobial surfaces to prevent infection, active antimicrobial agents are introducedto biomaterials by covalent interactions. In general, the antibacterial brush coatings can be subdivided into three categories:(1) polymer brushes composed of a bactericidal polymer, (2) polymer brushes functionalized with a bactericidal or bacterio-


static compound covalently linked to the brush, and (3) nonfouling polymer brushes that aim at repelling bacterial adhesionand mitigating biofilm formation [195].

3.2.1.2.1. Bactericidal polymer brushes. Bactericidal polymers generally contain cationic groups such as alkyl pyridinium orquaternary ammonium moieties. Cationic bactericidal polymers including vinyl monomers with tertiary amino groups suchas 2-dimethylaminoethyl methacrylate (DMAEMA) and 4-vinyl pyridine (4VP) can endow biomaterials with self-antibacterial abilities by two steps. The first step is polymerization or copolymerization by ATRP followed by quaternization[196,197]. In this reaction, quaternization of the amino groups of poly(2-dimethylaminoethyl methacrylate) (PDMAEMA) orP4VP brushes by an alkyl halide produces the biocidal functionality on the surface of the polymer-modified inorganic mate-rials [196–200]. However, different alkyl halides are used in the reactions to prepare PDMAEMA or P4VP brushes, namelyethyl bromide and hexyl bromide, respectively [196,198,200]. These modified surfaces exhibit substantial antimicrobialcapacity [196–201]. In addition to cationic antibacterial materials, anionic polyelectrolyte brushes of poly(3-sulfopropylmethacrylate) are synthesized by ATRP of 3-sulfopropylmethacrylate (SPMA) in a controlled manner and then introducingantibacterial silver ions to the polyelectrolyte film [202]. The silver-loaded sulfonate brushes inhibit the growth of bothGram-negative and Gram-positive bacteria. The brushes exhibit slow leaching of silver ions but retention of the silver ionsat the surface. The silver-loaded sulfonate brushes possess the desirable properties of an antibacterial surface [202].Additionally, the neutral polymer poly(2-(tert-butylamino)ethyl methacrylate) (PTBAEMA), a typical representativewater-insoluble biocide, has been used to prepare antibacterial surfaces by ATRP [203–205]. For instance, the modified LDPEexhibits effective antimicrobial activity [203]. Anchoring of the PTBAEMA chains onto PP produces long-lasting antibacterialactivity [205] and similarly, bacterial adhesion on the modified stainless steel diminishes significantly [204]. The major roleof PTBAEMA is believed to be the displacement of Ca2+ and/or Mg2+ ions from the outer bacterial membrane thereby disrupt-ing and compromising the membrane functions.

3.2.1.2.2. Polymer brushes functionalized with antibacterial compounds. In addition to directly grafting antibacterial polymers,a polymer brush without antibacterial properties can first be formed on the substrate by ATRP, followed by covalent linkageof an antimicrobial compound to the polymer brush. Antimicrobial peptides (AMPs) are attractive antibacterial compoundsdue to the broad spectrum activity against both Gram-negative and Gram-positive bacteria and relatively lack of toxicitytowards host cells [206–208]. Recent studies about this type of antibacterial coating were carried out on a Ti surface[208–210]. Copolymer poly-(N,N-dimethylacrylamide-co-N-(3-aminopropyl)-methacrylamide hydrochloride) (poly(DMA-co-APMA)) brushes were synthesized by ATRP and different types of AMPs were conjugated to the copolymer brushes bythe maleimide groups [208–210]. The peptides immobilized on the copolymer brush are shown in Fig. 16 [209]. The peptidedensity on the surface was larger when the brush layer was used for peptide conjugation than when the peptides weredirectly grafted onto the surface [208]. The AMP-immobilized polymer brushes also had excellent broad spectrum antimi-crobial activity and biofilm resistance in vitro, depending on the types of AMPs [209,210].

3.2.1.2.3. Antifouling polymer brushes with antibacterial properties. Antifouling polymer brushes are resistant to bacteria inaddition to preventing cell adhesion and non-specific protein adsorption. For example, as shown in Fig. 15c, the stainlesssteel surface grafted with polySBMA brushes by ATRP shows excellent resistance to bacterial adhesion.

3.2.1.3. Stimuli-responsive coatings. Stimuli-responsive polymers play an important role in biomedical applications such asdrug delivery, biomolecular diagnostics/biosensors, and others [51,211–214]. Stimuli can be divided into physical stimulisuch as temperature, pressure, electrical stimuli or magnetic fields as well as chemical stimuli such as pH, ionic strengthand chemical agents [211]. Temperature-sensitive polymers (also called thermo-responsive polymers) and pH-sensitivepolymers are the most commonly studied stimuli-responsive materials.

Owing to the lower critical solution temperature (LCST) of approximately 32 �C in an aqueous medium, PNIPAM is themost extensively studied thermo-responsive polymer. Well-defined PNIPAM can be prepared by ATRP to control the cellbehavior [158,215,216] and drug delivery [217,218]. Fig. 17 shows the mechanism of temperature-responsive PNIPAM

Fig. 16. Representation of peptide immobilized copolymer brush on surface [209]. Copyright 2011. Adapted with permission from Elsevier Science Ltd.

Fig. 17. The mechanism of temperature responsive PNIAM [133]. Copyright 2014. Adapted with permission from Elsevier Science Ltd.


[133]. Below the LCST, hydrogen bonds form. The PNIPAM chains extend and a stable hydration shell forms around thehydrophobic groups. Above the LCST, hydrogen bonding weakens and the intrinsic affinity of PNIPAM chains dominates.

The pH-responsive polymers consist of ionizable pendants that can accept and donate protons in response to changes inpH, while polymers containing ionizable groups in their backbone can form polyelectrolytes in an aqueous system. There aretwo types of pH-responsive polyelectrolytes: weak polyacids and weak polybases. Weak polyacids such as PAA or poly(methacrylic acid) (PMAA) accept protons at low pH and release them at neutral and high pH, whereas polybases such asP4VP are protonated at high pH and become positively ionized at neutral and low pH [18,211]. These pH-responsive poly-mers can be anchored onto the surface by ATRP to produce polymer brushes [220–224] and applications of pH-responsivepolymers include drug delivery and intracellular delivery in gene therapy [211].

PDMAEMA is both temperature- and pH-sensitive because it has amine pendant groups that can gain protons and can bepositively ionized at neutral and low-pH conditions. It has a LCST of 40–50 �C and a pKa of 7.0–7.5 [226–229]. The LCST ofPDMAEMA is strongly dependent on pH with no LCST at low pH [231,232]. The PDMAEMA-modified nanopore channels pre-pared by ATRP have high gating efficiency and reversibility, which may stimulate future applications in biosensors and actu-ators [229,230]. Furthermore, PDMAEMA can be used to prepare novel controlled drug release systems in the near future.

In addition to the four properties described above, some polymers such as MMA and styrene can be grafted onto bioma-terials by ATRP to facilitate grafting of functionalized polymeric groups onto these materials [237–242].

3.2.2. Factors influencing the properties of polymerized films by ATRPThe properties of surfaces such as protein adsorption and adhesion, cell adhesion behavior, and so on can be affected by

the thickness, grafting density, molecular weight, and chain length of the polymer brushes formed by ATRP. The thickness ofthe polymer brushes affects serum protein adsorption and fibroblast adhesion. Serum protein adsorption and fibroblastadhesion were effectively reduced, when the thickness of the poly(2-methacryloyloxyethyl phosphorylcholine) (PMPC) orPNVP brush layer increased [178,193]. The surface hydrophilicity increased when thick PNIPAM brushes were grafted ontothe surface because of more extended chain conformation with relatively high chain mobility accompanying polymer chainhydration [215]. When the thickness of the polymer brush is taken in account together with the grafting density, peptideadsorption and cell attachment can be regulated [160].

By adjusting the grafting density of the polymer brush, cell adhesion and cell morphology can be controlled [189]. Cellsadhered and spread and fibronectin adsorbed on the low graft density end, whereas there were fibronectin repulsion, andlittle cell adhesion and spreading at the high graft density end [189]. The low-density PNIPAM brush also showed the highestcell adhesion, featuring adherent cells with an elongated morphology [243]. Halperin and Kröger’s research revealed thatprotein adsorption could be effected by grafting density and polymerization degree of the PNIPAM chain [244].

The molecular weight of the polymer brush influences protein adsorption [182]. The copolymers containing polySBMAwith a larger molecular weight had high protein adsorption after direct adsorption onto a hydrophobic surface [182]. Whenthe grafting density was the same, the surface grafted with a relatively large molecular weight PDMAEMA showed almost100% killing efficiency [198]. Both the graft density and chain length affected protein adsorption [156]. The surfaces withlarge graft densities and PMPC chain lengths showed a dramatic reduction in the fibrinogen adsorption and fibrinogenadsorption depended more on the graft density than chain length [156].

4. Reversible addition-fragmentation chain transfer polymerization

RAFT polymerization is a versatile technique and addition-fragmentation transfer agents were directly utilized to controlradical polymerization in the mid 1980 s [245,246]. RAFT polymerization is different from the other CLRP methods in that itis compatible with a wide range of monomers and reaction conditions. This is because the conditions for RAFT polymeriza-


tion are similar to those in conventional radical polymerization except the added RAFT agent. Most monomers that are poly-merizable by conventional radical polymerization can be polymerized by RAFT contrary to other CLRP methods [247–249].RAFT polymerization can be utilized to form narrow polydispersity polymers or copolymers with controllable molecularweights and achieve high conversion and commercially acceptable polymerization rates.

4.1. Fundamentals of RAFT polymerization

Conventional RAFT polymerization contains three primary components: a monomer, a radical initiator, and a chain trans-fer agent. The key to successful RAFT is the appropriate selection of the RAFT agent or RAFT chain transfer agent (CTA). Trans-fer agents including various dithioesters, dithiocarbamates, trithiocarbonates, and xanthates have been used to control themolecular weight, molecular weight distribution, and even molecular architecture and some thiocarbonylthio compoundsare the most effective transfer agents [41,247,248,250,251].

The mechanism of exchange between dormant and active species involves a series of linked equilibria and so thesequence of addition-fragmentation equilibria is important to RAFT polymerization, as shown in Fig. 18 [252]. The polydis-persity and molecular weight control depend on the nature of the groups Z and R in the thiocarbonylthio compounds (S@C(Z)SR) in which the reactive C@S double bond and weak SAR single bond are key structures. When selecting Z and R, it is criticalto determine both the addition and fragmentation rates of the RAFT agent as they can alter the effectiveness of the RAFTagent [253]. Z, a group that modifies the reactivity of the thiocarbonylthio compound and derived adduct radical, should acti-vate the C@S double bond towards radical addition in order to ensure a high transfer constant, and this is a major factoraffecting the lifetime of the intermediate radical resulting from the addition of a radical species across the C@S bond. TheR group allows for fine tuning of the overall reactivity, thus enabling these species to effectively mediate polymerizationin a controlled manner. It has a profound effect on the polymerization kinetics and overall degree of control. The R groupplays two roles. First of all, it should be a good free radical leaving group and secondly, the radial (R�) generated from homo-lytic dissociation must efficiently reinitiate polymerization to induce chain transfer. The structures of the Z and R groups areclosely related and not all combinations of Z and R groups produce the effective/efficient RAFT agents.

While initiation and radical–radical termination can occur in conventional radical polymerization, RAFT is simply aconventional free radical polymerization technique that utilizes the suitable thiocarbonylthio species. In the early stagesof polymerization, fragmentation of the intermediate radical provides a polymeric thiocarbonylthio compound [PnS(Z)C@S(3)] and new radical (R�) by adding a propagating radical (Pn�) to the thiocarbonylthio compound [RSC(Z)@S (1)]. The radical(R�) reacts with a monomer to form a new propagating radical (Pm�). Chain extension of the polymeric thiocarbonylthio

Fig. 18. The mechanism of reversible addition-fragmentation chain transfer [252]. Copyright 2009. Reproduced from Moad et al. (2009), with permissionfrom CSIRO Publishing.


compound [PnS(Z)C@S (3)] involves essentially the same process, while rapid equilibrium between the active propagatingradicals (Pn� and Pm�) and dormant polymeric thiocarbonylthio compounds (3) results in an equal probability for all thechains to evolve into polymers with narrow dispersity. When polymerization is complete, most of these chains retain thethiocarbonylthio end-group and are stable.

At least four factors determine the effectiveness of the thiocarbonylthio compound (1) on the basis of the addition-fragmentation mechanism: (a) the rate constant of the reaction of the thiocarbonylthio compound (1) with the propagating(or initiating) radicals (kadd), (b) the absolute rate constant of the reaction with the intermediate radical (2) fragments (kb),(c) partitioning of the adduct (2) between the starting materials and products (determined by the relative magnitude of kaddand kb), and (d) the rate and efficiency at which the expelled radicals (R�) reinitiate polymerization.

RAFT polymerization features a different mechanism to achieve control than ATRP and NMP. In ATRP and NMP, by rever-sible deactivation of the propagating radicals, radicals can also be derived from the dormant species. In contrast, in RAFTpolymerization, an external source of free radicals is required to initiate and maintain polymerization because free radicalsare neither formed nor destroyed in the process and the deactivation-activation equilibrium are chain-transfer reactions. Therate of polymerization is determined by the combined effects of the deactivation-activation equilibrium and ‘‘persistentradical effect” [248].

4.2. Surface polymerization of biomaterials by RAFT

Surface RAFT polymerization can be divided into three categories according to the type of species anchored to the sub-strate: (1) initiator, (2) CTA via the Z-group, and (3) CTA via the R-group. The initiator approach yields low grafting densitiesdue to radical recombination. Because the radical chains have to diffuse through the polymer layer to react with the RAFTagent immobilized on the surface, the polymer chain does not grow from the surface. The Z-group approach is not asurface-initiated polymerization/grafting-from method and has the same limitations as the typical grafting-to methodknown as the autophobic effect for polymer brushes [254,255]. The R-group approach is the most effective way to obtainthick polymer brushes with a narrowmolecular weight distribution. Polymerization is initiated in a solution with or withouta sacrificial CTA. Fig. 19 compares the R-group and Z-group approaches with regard to CTA-anchored surface-initiated RAFTpolymerization [255].

4.2.1. Monomers used in RAFT polymerizationMonomers with functional groups such as hydroxyl groups, carboxylic acids, carboxylic acid salts, amides, and tertiary

amines can be used in RAFT polymerization. The typical monomers used to modify biomaterial surfaces via RAFT polymer-ization include styrene derivatives, acrylates and acrylamides, methacrylates, and methacrylamides and vinyl esters, asshown in Table 3 [256–285]. RAFT polymerization can be used to immobilize functional polymers such as PS, PAA, andPNIPAM in the maleimidated surfaces [257]. Fig. 20 depicts the process of immobilization of thiol-terminated polymerson the maleimidated surfaces. The trithiocarbonate-terminated polymers are first synthesized by RAFT polymerization,followed by conversion of the trithiocarbonate group to a thiol group by NaBH4 reduction [257]. Fig. 21 schematicallyillustrates the process of grafting MMA onto the surface of biomedical stainless steel by RAFT polymerization [271]. Thismethod can be combined with other techniques to functionalize biomaterials. For example, RAFT and the ‘‘graft from”strategy were employed to integrate the copolymers of positive charged quaternary amines and PEG units to modify the sur-face of mesoporous silica nanoparticles (MSNs). The resulting PEG coating with a large positive zeta potential showed nearly

Fig. 19. The comparison of R-group and Z-group approaches for CTA anchored surface-initiated RAFT polymerizations [255]. Copyright 2011. Adapted withpermission from Elsevier Science Ltd.

Table 3Biomaterial grafted with bioactive polymers through RAFT.

Grafted monomer Substrate Ref.

Styrene derivatives Styrene Gold nanorod and nanoparticle, siliceous, silicon wafer, silica particles, cellulose [256–263]4VP SiO2 nanoparticles [264]

Acrylates AA Gold nanorod, siliceous [256,257]Methyl acrylate Silica particles, silicon, gold nanoparticle [259,262,263]PEGA Protein [265,266]HEA Protein [267]

Acrylamides NIPAM Siliceous, protein, silica particles, magnetic nanoparticle, chitosan [257,259,260,267–270]DMA Silicon wafer, silica particles [258,259]

Methacrylates MMA Stainless steel, Au, silicon wafer, Silica Particles [258–260,262,271,272]DMAEMA Gold nanorod, cellulose, Silicon [256,262,273–276]AHMA Silica Nanoparticles [277]HEMA Silicon wafers [278]PEGMA Gold, silicon wafer, magnetic nanoparticle [269,279,280]PEGMEMA Au [272]

Methacrylamides AEMA Fe3O4 NP [276,281–283]APMA Antimicrobial peptides [281,284]HPMA Au nanoparticle, gold nanoparticle [263,284]DMAPMA Au nanoparticle, Antimicrobial peptides [284,285]

Fig. 20. Preparation of maleimidated surface and immobilization of thiol-terminated polymer via the Michael addition reaction [257]. Copyright 2012.Adapted with permission from the American Chemical Society.


2-fold enhanced permeability and retention (EPR) effect in addition to a longer blood half-life compared to coating with PEGonly. The polymerized MSNs showed higher efficacy in doxorubicin (DOX) drug delivery and suppression of side effects com-pared to the free drug, as confirmed by the in vivo assessment indicating inhibition of tumor growth in tumor-bearing nudemice after intravenous injection of DOX-loaded MSN-PEG+ (shown in Fig. 22) [269].

4.2.2. Factors influencing the properties of polymerized films by RAFTThe biocompatibility and antibacterial activity are important to biomaterials, particularly implants. Antimicrobial PEG

functionalized polymers synthesized from methacrylate derivatives by RAFT polymerization had shorter hydrophobic chainlengths (i.e., methyl and ethyl) and better antimicrobial properties [286]. DeGrado and Kuroda synthesized antimicrobialpoly(methacrylate) derivatives that could be tailored by tuning the molecular size and number of hydrophobic side chains[287]. When the hydrophobic side chains were shorter, the antibacterial activity was higher [274]. The effect of the pendantamine structure on antimicrobial activity in amphiphilic polymethacrylates was investigated [288,289]. The primary aminesincreased the antimicrobial activity. Polymers containing quaternary ammonium groups increased the antimicrobial potencyby adding extra hydrophobic functional groups. However, the homopolymer of the primary amine-containing methacry-lamide monomer was found to deliver significantly greater antimicrobial activity than poly(methacrylate) due to improvedhydration of the polymer backbone [285]. These results have increased the interest in surface grafting of poly(methacry-lamide) via RAFT polymerization to improve the antibacterial activity of biomaterials.

Ahmed and co-worker studied the effects of the polymer architecture, composition, and molecular weight on theproperties of glycopolymer-based non-viral gene delivery systems [281]. When the architectures (block versus random)were different, the statistical copolymers showed lower toxicity and higher gene expression than the corresponding diblockcopolymers in the presence and absence of serum [281]. When the arbohydrate residues in the copolymers increased, thetransfection efficiency of the polymers decreased because the critical composition of the carbohydrate segment in thecopolymers enhanced gene delivery and had low toxicity [281]. The large molecular weight statistical copolymers showedgreater cell viability and gene expression in the presence and absence of serum [281]. Boyer found that a lower critical solu-

Fig. 21. General procedure for the preparation of CTA-modified stainless steel surfaces and subsequent formation of PMMA brushes by reversible addition-fragmentation chain transfer polymerization [271]. Copyright 2013. Adapted with permission from the American Chemical Society.

Fig. 22. In vivo inhibition of tumor growth in tumor-bearing nude mice after intravenous injection of DOX-loaded MSN-PEG+. (A) Time dependent tumorgrowth curves at 3 d, 6 d, 9 d and 12 d after the injections of saline, empty MSN-PEG+, free DOX and DOX-loaded MSN-PEG+. The dosage of the DOX drugwas 100 mg per mouse for each administration every six days. ⁄⁄ and ⁄⁄⁄ represent significant differences in tumor growth inhibition found by comparingthe four groups at P 6 0.005 and P 6 0.0005, respectively. (B) The photographs of nude mice before and at the end of treatment showing the different tumorgrowth-inhibition effects. (C) The TUNEL staining assay showing enhanced apoptosis and cell death by DOX-loaded MSN-PEG+ compared to free DOX. Theupper blue fluorescence images are of the cell nucleus after DAPT staining. The lower green fluorescence images (200�) represent of the TUNEL-positivecells [269]. Copyright 2014. Adapted with permission from the Royal Society of Chemistry.



tion temperature of the conjugates increased with decreasing molecular weight of the PNIPAM block conjugated to BSA[267].

5. Other surface radical polymerization methods

5.1. Plasma-induced radical polymerization

Plasma, a quasi-neutral gas considered as the fourth state of matter, consists of a collection of neutrons, electrons, ions,and other atomic and molecular species [290]. Plasma surface modification is a rapid, dry, economical, effective, and envi-ronmentally friendly materials processing technique which is becoming increasingly popular in the biomedical community.This method can enhance the surface properties while preserving the bulk properties [291]. In addition to creating activesurface species, it is used to remove contaminants from the polymer surface, an additional benefit that enables the repro-ducible preparation of a homogenously treated polymer surface [292].

Plasma-induced graft polymerization is a two-step process in which the functional groups and reactive sites are incorpo-rated onto the surface. This process is different from activation in that it adds the materials to the surface instead of func-tionally modifying the surface. It results in a high surface density of the polymer chains which are initiated and polymerizeddirectly from the substrate to enhance the stability with under chemical and thermal stress conditions as well as minimizingthe growth of polydispersed chains [293]. Several studies have confirmed that plasma-induced graft polymerization is adesirable method to functionalize interfaces for immobilization of biomolecules [294,295].

In low-pressure plasma methods, the non-thermodynamic-equilibrium processes in weakly ionized gases not only enablethe design of a novel class of materials generally in the form of surfaces, but also allow for the modification of the surfaceproperties of any solid substrates compatible with medium–low vacuum conditions [296]. Plasma-induced graft polymer-ization has hitherto involved low-pressure plasma initiation and surface grafting onto polymeric materials but there havebeen few studies on inorganic oxides [297–299]. It has been reported that the plasma treatment time and radio frequency(RF) power of the plasma generator can control and optimize the surface radical density [298–300]. However, the radicaldensity on the surface of inorganic oxide substrates modified by the low pressure plasma approach may be limited bythe native oxide surface chemistry [301]. These findings indicate that this approach is not sufficient to achieve high-density surface activation and graft polymerization/modification of inorganic substrates with a large surface area. Recently,atmospheric pressure plasma-induced graft polymerization (APP) has been adopted in surface functionalization [294,302–

Fig. 23. Illustration of multistep process plasma-induced graft polymerization [304]. Copyright 2007. Adapted with permission from the AmericanChemical Society.


305] and the typical APP process is described in Fig. 23 [304]. A high throughput platform that allows for the synthesis,screening, and selection of the desirable monomers in a relatively short period of time has been combined with atmosphericpressure plasma to synthesize, test, and select the best protein-resistant membranes [302,305].

Different polymers can be grafted on various substrates as shown in Table 4 [292–325]. For instance, 1-vinyl-2-pyrrolidone monomer was polymerized on the surface of silicon by plasma-induced graft polymerization [304]. The surfacemorphology of the grafted polymer layer depended on the density of the surface-active sites generated by the plasma treat-ment as well as polymerization conditions suggesting commercial potential of large-scale nanostructured surface treatment.

Ti and its alloys have been widely used in biomedical devices but the bioinert nature and susceptibility to bacterial infec-tion restrict wider application to orthopedic implants in vivo. Plasma polymerization using allylamine, acrylic acid (AA) orother monomers as precursors has been applied to modify these metals [307–310]. The modified titanium surface can befurther functionalized by attaching the desired molecules such as RGD, PEG, collagen I, or antimicrobial peptides to enhancethe osteoinductivity, osteoconductivity, and antibacterial activity. To improve the thromboresistance of the titanium alloy,Ye carried out plasma-induced graft polymerization of MPC and after modification, platelet deposition and bulk phaseplatelet activation diminished significantly relative to the unmodified titanium alloy [312].

Plasma treatment is also effective for biopolymers. For example, polycarbonate (PC) membranes are good candidates indrug delivery applications due to their unique structure, excellent biocompatibility, and wide commercial availability [313].Baumann utilized plasma-induced graft polymerization to tune the hydrophilicity and permeability resistance of PC mem-branes in drug delivery systems [292]. Poly(ethylene terephthalate) (PET) was used to treat the pathology of large-diameterarteries [314,315]. To overcome the chemical and biological inertness of the PET surface, bioactive biomolecules such as AAand gelatin were grafted onto the PET films by plasma-induced polymerization for protein immobilization and smooth mus-cle cell seeding. The results showed that the collagen immobilized surface provided an excellent substrate for the growth ofhuman smooth muscle cells [294,314].

Some common biocompatible polymers such as PTFE, polyurethanes (PU), and poly(l-lactic acid) (PLLA) can be modifiedby this method to enhance the biological functions [316–322]. Wang and Chen grafted AA onto a PTFE film using peroxidesto initiate polymerization during which PTFE reacted with AA. The remote argon plasma was used to etch the surface to cre-ate peroxides after exposure to oxygen [317]. The water contact angles decreased from 108� to 41�, indicating enhancedhydrophilicity after treatment as a result of the permanent AA grafted layer on the surface [317]. In contrast, plasmapre-treatment followed by UV-induced grafting of glycidyl methacrylate (GMA) to PTFE enhanced the hydrophobicity asmanifested by the larger contact angle compared to the pristine surface [316]. In addition to varying the hydrophilicity/hydrophobicity, plasma induced polymerization can be employed to modulate the cellular behavior on biopolymers. Hsuand Chen demonstrated that the lactide-grafted PU surfaces after the plasma-induced graft reactions favored adhesion of3T3 fibroblasts and human umbilical vein endothelial cells. Platlet activation also decreased demonstrating potentialapplications of PU vascular grafts [318]. In addition to regulation of the cell adhesion behavior, the cell morphology canbe modulated by this technique. Ding et al. immobilized chitosan on PLLA films by plasma graft polymerization [319] anddespite no obvious difference in the mouse fibroblasts densities among the three types of surfaces, the cell morphologywas quite different. The fibroblasts hardly spread and tended to become round on the modified surface but freely extendedon the untreated surface [319]. A similar phenomenon was observed from human hepatocytes, indicating that plasma graftpolymerization could be used to control the cell morphology [319].

The plasma-induced polymerized surface can be further functionalized by other polymerization techniques. For example,with regard to stainless steel, pre-treatment with the silane coupling agent was required to provide a suitable graftingenvironment and the substrate was subsequently activated by an argon plasma prior to undergoing UV-induced graftpolymerization with PEGMA (Fig. 24) [306]. The graft concentration was demonstrated to rely on the macromonomer con-centration and UV graft polymerization time. The PEGMA graft-polymerized stainless-steel with a large graft concentrationwas effective in preventing protein adsorption [306]. Plasma-induced graft polymerization can be combined with ATRP to

Table 4Biomaterial grafted with bioactive polymers through plasma-induced graft polymerization.

Substrate Monomer Ref.

Silicon VP [304]Stainless steel PEGMA [306]Titanium AA, allylamine, styrene [307–311]Titanium alloy MPC [312]PC AEMA [292]PET AA [294,314,315]PTFE GMA, AA [316,317]PU L-lactide [318]

PLLA chitosan [319]PS NVP, peptide [323,324]PES methacrylamide derivatives [305]PP AA [325]

Fig. 24. Schematic illustration of the processes of surface modification of stainless steel by silane treatment, Ar plasma treatment and UV-induced surfacegraft polymerization [306]. Copyright 2001. Adapted with permission from Elsevier Science Ltd.


modify biomaterials. The quartz crystal microbalance surfaces were first grafted with poly(allyl alcohol) by plasma-inducedgraft polymerization to create a surface-grafted polymerization initiator and ATRP was then performed to graft poly(tert-butyl acrylate) that was eventually converted into poly(acrylic acid) by hydrolysis [326]. The tethered poly(acrylic acid)grafted quartz crystal microbalance synthesized by this method increased the frequency response to the pH, immobilizationamount of the antibody, and immunosensor response [326].

5.2. Radiation-induced radical polymerization

Radiation-induced radical polymerization is well known for its advantages and potential in modifying the physical andchemical properties of pre-existing polymeric materials without altering their inherent properties. Catalysts and additivesare not required during this process and the products maintain the purity and molecular weights of the products. The degreeof grafting can be controlled. Another merit of the radiation techniques is that the processed materials are not only non-toxic, but also sterilized. Radiation can be used to combine two highly incompatible polymers in a unique way to fabricategraft copolymers with new properties.

Common radiation energy sources include gamma ray, UV, and electrons [327]. High-energy irradiation ofmacromolecules can form polymeric ions and free radicals on the polymer as the result of homolytic fission [15].Radiation-induced graft polymerization can be divided into ionic grafting and free-radical grafting. Ionic grafting may beof two different types: cationic or anionic and the potential advantage of ionic grafting is the large reaction rate. MeV protonbeams have been successfully applied as ionizing radiation to induce graft polymerization on PP or PE substrates [328–330].

Free-radical grafting is a type of radiation-induced graft polymerization. The basic process of radiation-induced free rad-ical graft polymerization is described in Fig. 25 [331]. When the polymeric substrate ‘A’ is exposed to ionizing radiation asmacro-initiators, the active sites formed randomly along the polymer chain initiate free radical polymerization of monomer‘B’. Unlike chemical initiation, there is no contamination from the initiators. Generally, in radiation initiation, absorption ofenergy by the backbone polymer initiates a free radical process, while the initiator decomposes into fragments which thenattack the base polymer, leading to the generation of free radicals as in chemical initiation.

Fig. 25. The use of ionizing radiation to graft monomer ‘B’ to polymer ‘A’ to form a graft copolymer [331]. Copyright 2003. Adapted with permission fromElsevier Science Ltd.


Radiation-induced graft polymerization is typically carried out by three routes: (1) simultaneous irradiation, (2) pre-irradiation method, and (3) peroxide-irradiation [332,333]. The simultaneous irradiation method is the simplest one forthe preparation of graft copolymers. In this method, the polymer substrate is immersed in a monomer which may be vapor,liquid, or in a bulk solution. When the substrate and monomer are exposed to radiation, free radicals are formed to producethe grafted polymer. The pre-irradiation method consists of two steps. Firstly the polymer substrate is irradiated in vacuumor under an inert atmosphere to generate free radicals and secondly, the irradiated polymer substrate reacts with the mono-mer after air has been removed. Similarly to the pre-irradiation method, the peroxide-irradiation method only exposes thepolymer substrate to the ionizing radiation in the presence of air or oxygen, leading to the formation of peroxides orhydroperoxides on the polymer substrate. When heated in the presence of monomer, the peroxy products decompose toradicals, depending on the nature of the polymeric backbone and irradiation conditions, and these radicals initiate grafting.Each method has its own advantages and disadvantages. The pre-irradiation and peroxidation methods are convenient inthat the polymeric substrate can be irradiated and stored for a long period of time before performing the grafting step. Itis useful when access to a radiation source is limited. Furthermore, little homopolymer formation occurs and grafting canbe performed at any time away from the radiation sources, making the pre-irradiation method an attractive option[334–336]. The graft yields of the simultaneous method are generally higher than those of the pre-irradiation methodsdue to the smaller radical loss through the decomposition reactions in the latter method. However, the simultaneousmethod is limited by the high degree of homopolymer formation. The mechanisms of the three processes are presentedin Fig. 26 [337].

Similar to plasma polymerization, radiation-induced graft polymerization is a promising technique for the surface mod-ification of biomaterials because it can enhance the biocompatibility or provide the desired surface chemistry by immobi-lizing certain biomolecules onto the surface of biomaterials that come in contact with blood [338,339]. Casimiro andcolleagues grafted 2-hydroxyethyl methacrylate (HEMA) to chitosan using the three methods of chemical initiation,photo-induction and radiation-induced polymerization and radiation-induced polymerization produced a higher graft yieldthan the other two methods [340]. The monomers commonly used in radiation initiation are MMA, AA, acrylamide (AAm),

Fig. 26. The mechanisms of three different radiation processes [337]. Copyright 2000. Adapted with permission from Elsevier Science Ltd.

Fig. 27. Water contact angle and platelet adhesion morphology of (A) pristine-SS, (B) PDA-SS and (C) PHEMA-g-PDA-SS [341]. Copyright 2014. Adapted withpermission from Elsevier Science Ltd.


NIPAM, HEMA, vinyl alcohol, vinyl pyrrolidone, glycidyl methacrylate, and styrene [333,337,341–348]. Jin et al. graftedHEMA onto dopamine-coated stainless steel by c irradiation and the HEMA grafted stainless steel surface exhibitedenhanced hydrophilicity, excellent anti-platelet adhesion, and high hemocompatibility as shown in Fig. 27 [341].

Another important biomedical application of radiation-induced grafting is modification of fabrics such as cotton fabricsand non-woven polypropylene fabrics to obtain specific biological functions [349–352]. Cotton fabrics possess desirableproperties such as good folding endurance, breathability, and good bio-degradability, but their major shortcoming is thatthey are prone to bacterial and fungal growth under moist conditions. Therefore, development of antibacterial fabrics forhospitals is important. Goel prepared antibacterial cotton fabrics by radiation-induced grafting of [2-(methacryloyloxy)ethyl]trimethylammonium chloride onto cotton. The modified cotton exhibited good antibacterial activity against bothgram-positive bacteria such as S. aureus and Bacillus cereus and gram-negative bacteria such as E. coli and Pseudomonasfluorescens [349,350]. Although polypropylene non-woven fabrics (PPNWF) have some excellent properties such as non-toxicity, chemical stability and low cost, the poor hydrophilicity and biocompatibility of unmodified PPNWF can often inducethrombus formation due to blood protein adsorption and platelet adhesion [351,353,354]. These shortcomings could beovercome by introducing bioactive groups onto the surface by radiation-induced grafting [351,352]. For example,N-vinyl-2-pyrrolidone could be grafted onto PPNWF using a simultaneous irradiation-induced graft polymerization tech-nique to enhance the surface hemocompatibility of PPNWF [351]. Because some diseases induce changes in the body tem-perature and/or pH [355–358], Kuamari et al. used a temperature- and pH-responsive monomer (NIPAM, AA) to modifyPPNWF, and the morphology of the modified surface was significantly affected by the grafting process [352].

RAFT polymerization could be applied to radiation-induced graft polymerization of styrene from cellulose to preparehighly hydrophobic surfaces [359]. Radiation-induced graft polymerization in combination with NMP is a useful tool inthe production of proton exchange membranes for fuel cell applications [360]. The membrane prepared by bothradiation-induced graft polymerization and ATRP had better chemical stability than the membrane prepared by conventionalradiation-induced grafting [361].

5.3. Photo-induced radical polymerization

Photografting polymerization was first reported in the 1950s. The monomers were photo-grafted by UV radiation ontothe polymer substrate blended with a photoinitiator [362]. By adopting high energy irradiation initiated surface polymeriza-tion, photografting has been used in surface modification [330]. UV irradiation is used to extract active hydrogen atoms fromthe outer substrate surface and the surface free radicals initiate the in situ polymerization of the monomers attached to thesubstrate to form the surface grafted chains [330,363–365]. In comparison with other surface grafting approaches, photo-induced grafting has some advantages such as the low cost, simple equipment, mild reaction conditions, ease of industrial-ization, weak penetrability, and long-term stability of the grafted chains without affecting the bulk materials [322,366–370].

The photoinitiators are important in most photo-induced polymerization processes and benzophenone (BP) and itsderivatives are most frequently used because they produce the highest grafting efficiency. Using this initiator, Ma et al. pro-posed a novel photo-induced living graft polymerization process consisting of two steps (Fig. 28) [364]. In the first step, inthe absence of monomer solutions, the surface photoinitiators were produced by the combination of surface radicals andsemipinacol radicals generated by BP abstracting hydrogen from the substrate. In the second step, graft polymerizationwas initiated by the surface initiators under UV irradiation after the monomer solution was added to the active substrate.During this process, initiator formation and graft polymerization occurred in different steps ensuring that they were inde-pendent and enabling independent control over the graft density and chain length of the graft polymer. This process alsoeliminated the formation of undesired homopolymers and crosslinked or branched polymers [364].

Although most photografting polymerization techniques require the addition of photoinitiators, some monomers haveunique self-initiating properties, for instance, the ability to undergo photografting polymerization without photoinitiators[371]. The typical monomers include styrene, maleic anhydride, acrylic acid, methacrylic acid, 2-hydroxyethyl acrylate,

Fig. 28. Schematic diagram of a novel photo-induced living graft polymerization method [364]. Copyright 2000. Adapted with permission from theAmerican Chemical Society.


and 2-hydroxyethyl methacrylate [366,372]. In addition, some polymeric substrates can produce semibenzopinacol-containing BP radicals upon photo-irradiation. For example, Kyomoto et al. conducted photo-initiated polymerization ofMPC onto poly(ether-ether-ketone) (PEEK) in the absence of photoinitiators to improve the surface properties [373]. BecauseMPC is a highly hydrophilic compound, the water wettability and lubricity of the PMPC-grafted PEEK surface were consid-erably enhanced compared to the untreated PEEK surface. It was further improved as the photo-irradiation increased asshown in Figs. 29 and 30, respectively [373]. During photo-irradiation, a nanometer-scale PMPC layer formed on the surfaceand became thicker with time (Fig. 31) and the amount of BSA adsorbed on the PMPC-grafted PEEK surface decreased withtime extends compared to the untreated PEEK (Fig. 32).

As shown in Table 5 [363,364,369,371,373–394], different monomers are selected for different substrates during photo-induced grafting polymerization. For example, MMA, N-vinyl-2-pyrrolidone (NVP), AA, and AAm are frequently used asmonomers in photo grafting modification of PP because the poor biocompatibility and bad hydrophilicity restrict biomedicalapplications despite the high void volumes, well-controlled porosity, and chemical inertness as microfiltration and ultrafil-tration membranes [363,364,374–377]. Yang and co-workers utilized the sugar-containing monomer D-gluconamidoethylmethacrylate (GAMA) to modify the surface of PP in UV-induced photografting polymerization to improve the surfacehydrophilicity and hemocompatibility [374]. As shown in Fig. 33 [374], the static, advancing, and receding contact anglesdecreased with the degree of GAMA grafting, indicating the improvement of the surface hydrophilicity by grafting GAMA.The number of adhered platelets decreased as the degree of GAMA grafting increased and in the case of the higher degreeof GAMA grafting (4.72 and 6.28 wt%), the adhered platelets had almost the spherical shape, meaning that the surface didnot activate the platelets so that the hemocompatibility was improved (Fig. 34) [374]. In addition to biocompatibility, theanti-bacterial polyvinyl-pyrrolidone-iodine complexes could be grafted onto the surface of the PP film to improve its

Fig. 29. Static-water contact angle (n = 15) of PMPC-grafted PEEK as a function of the photo-irradiation time with MPC concentrations of 0.25 and 0.50 mol/L. Bar: Standard deviation [373]. Copyright 2010. Adapted with permission from Elsevier Science Ltd.

Fig. 30. Coefficient of dynamic friction (n = 5) of PMPC-grafted PEEK as a function of the photo-irradiation time with MPC concentrations of 0.25 and0.50 mol/L. Bar: Standard deviation [373]. Copyright 2010. Adapted with permission from Elsevier Science Ltd.

Fig. 31. Cross-sectional TEM images of PMPC-grafted PEEK obtained with a MPC concentration of 0.5 mol/L and various photo-irradiation times. Bar:100 nm [373]. Copyright 2010. Adapted with permission from Elsevier Science Ltd.


antibacterial performance and the products exhibited clear anti-bacterial activity with a broad spectrum and high efficiency(Fig. 35) [375].

Polydimethylsiloxane (PDMS)-based materials are promising materials in ophthalmologic applications, microfluidicdevices, artificial lungs and artificial finger joints due to the high oxygen permeability, convenient processability, goodmechanical properties, optical transparency, self-sealing properties, and chemical stability [380,381]. In spite of these advan-

Fig. 32. Amount of adsorbed BSA (n = 3) of PMPC-grafted PEEK as a function of the photo-irradiation time with MPC concentrations of 0.25 and 0.50 mol/L.Bar: Standard deviation [373]. Copyright 2010. Adapted with permission from Elsevier Science Ltd.

Table 5Biomaterial grafted with bioactive polymers through photo-induced graft polymerization.

Substrate Monomer Ref.

PP MMA, NVP, AA, AAm, GAMA [363,364,374–377]PEEK MPC [373]PDMS AA, PEGMA, PEGDA, MPC [378–380]PE MPC, AA [382–384]PCL NVP, MAA [385,386]PLLA NVP [385]PLGA NVP [385]PLA AA, AAm [369]SEBS NVP [371]PES AA, HEMA, EDA [387]PVDF AA, HEMA. EDA [388]PU AAm [393]PCU PEGMA [394]Silicon wafer MMA, St, MA, NIPAM, HEMA, AAm [389–391]Porous polymer monolith PEGMA [392]

Fig. 33. Comparison of advancing contact angle (AC), receding contact angle (RC), and statistic contact angle (SC) on the nascent and modified PPMMs: (1)–(7) 0, 1.52, 2.58, 3.79, 4.72, 5.76, and 6.28 wt% GAMA grafted PPMMs, respectively [374]. Copyright 2005. Adapted with permission from the AmericanChemical Society.


Fig. 34. Scanning electron micrographs of adhered platelets on the nascent and modified PPMMs: (a)–(d) 0, 1.52, 4.72, and 6.28 wt% GAMA-grafted PPMMs,respectively (left 2500�, right 10,000�) [374]. Copyright 2005. Adapted with permission from the American Chemical Society.


tages, PDMS still requires further modification because of its inherent hydrophobicity and susceptibility to biofouling.Sugiura et al. applied photo-induced graft polymerization to micropatterned surface modification of PDMS with polyethy-lene glycol diacrylate (PEGDA) [379]. As shown in Fig. 36, the green fluorescence intensity was proportional to the amountof adsorbed BSA, indicating that the untreated PDMS area adsorbed more protein than the PEGDA-grafted area. That is, the

Fig. 35. Change of the viable cell number of E. coli, S. aureus, and C. albicans after contact with the virgin and modified film surfaces for different times. E.coli⁄, S. aureus⁄, C. albicans⁄”: the microorganisms contacting with the virgin films [375]. Copyright 2005. Adapted with permission from Wiley Periodicals,Inc.

Fig. 36. FITC-BSA adsorption onto the micropatterned-PEGDA-grafted PDMS [379]. Copyright 2008. Adapted with permission from Elsevier B.V.


PEGDA-grafted layer prevented non-specific protein adsorption on PDMS [379]. PDMS could also be imparted with biologicalfunctions by photo grafting other monomers such as MPC [380]. The cell adhesion behavior could be tailored by this methodand as shown in Fig. 37 the HepG2 cells were observed to attach to the gelatin-adsorbed area and not to the PEG-coveredarea.

The wear resistance of some biopolymers can be improved by photo grafting techniques. As the important component inthe bearing couple of artificial joints, ultra-high molecular weight polyethylene (UHMWPE) releases wear debris at the inter-face between the implant and surrounding tissues during contact with the metal femoral component under a load usuallyinducing osteolysis. Minimizing the formation of wear particles from UHMWPE is the best way to prevent osteolysis. Denget al. grafted AA onto UHMWPE by photo-induced radical graft polymerization. The wear rate of the untreated UHMWPEwashigher than that of modified UHMWPE in saline, distilled water, and calf serum lubricant and the modified UHMWPE graftedwith 3.5% showed the smallest wear rate which was just quarter that of the untreated sample (Fig. 38) [382]. Anothercommon acetabular material is cross-linked polyethylene (CLPE) because of its excellent anti-wear properties [383–384].Kyomoto and co-workers proposed an artificial hip joint system with the MPC polymer grafted onto the surface of CLPE(CLPE-g-MPC) by UV irradiation at 5 mW/cm2 for 10–360 min at 60 �C [383]. As shown in Table 6 [383], the physical andmechanical properties of the CLPE and CLPE-g-MPC changed little but the friction coefficient of CLPE-g-MPC decreased shar-ply by 88% compared to the untreated CLPE [383].

Generally, tissue regeneration or reconstruction requires high affinity between the materials and surrounding tissues orcells. Next-generation materials must be ‘‘smart” to induce cell growth and differentiation. Photo-induced surface graftingcan play this role in some polymeric materials. For example, biodegradable PCL can be used in drug delivery systems,tissue-engineered skin, and scaffolds to foster the growth of fibroblasts and osteoblasts. However, PCL is hydrophobic and

Fig. 37. Optical micrographs of HepG2 cells cultured on the micropatterned- PEGDA-grafted PDMS observed through a 4 � objective lens (a) and a10 � objective lens (b) [379]. Copyright 2008. Adapted with permission from Elsevier B.V.

Fig. 38. Wear rate of untreated and modified UHMWPE in different lubricant [382]. Copyright 2013. Adapted with permission from Springer.


unfavorable to cell growth. Zhu et al. grafted PMAA and gelatin onto PCL under UV irradiation and as shown in Fig. 39, thePCL membrane modified with PMAA or gelatin exhibited improved cytocompatibility with human ECs [386]. Similarly, NVP

Table 6Properties of CLPE and CLPE-g-MPC [383]. Copyright 2013. Adapted with permission from Springer.

Properties CLPE CLPE-g-MPC

Physical properties Density (g/cm3) 0.944 (0.002) 0.943 (0.001)Swelling ratio 2.99 (0.11) 2.94 (0.10)Network chain density (�103�mol/ml) 0.437 (0.043) 0.459 (0.044)M.W. between Cross-links (g/mol) 2165 (214) 2069 (186)Cross-link density (mol%) 0.65 (0.06) 0.68 (0.07)

Mechanical properties Yield strength (MPa) 23.2 (0.4) 23.1 (0.5)Impact strength (kJ/m2) 75.0 (1.4) 77.0 (1.9)Creep deformation (%) 0.89 (0.17) 0.63 (0.40)Hardness (shore D) 68.2 (0.9) 68.4 (0.5)

Tribological properties Friction coefficient 0.0075 0.0009Wear rate (mg/106 cycles) 3.12 �1.43

The standard deviation is in parenthese.

Fig. 39. The cell morphology after cultured for 4 d on TCPS (a), PCL-g-PMAA with WCA 69.91 (b) and PCL-g-PMAA-gelatin with WCA 62.61 (c), respectively.(1) refers to higher magnification while (2) to an overview under SEM with the same sample. Cell seeding density 15 � 104 cm�2. (TCPS: tissue culturepolystyrene, WCA: water contact angle) [386]. Copyright 2002. Adapted with permission from Elsevier Science Ltd.


was covalently grafted onto PCL, PLLA, and poly(lactide-co-glycolide) (PLGA) under irradiation to produce a bioactive surfacefavoring adhesion, growth, and proliferation of normal human cells [278].

In addition to the aforementioned polymer substrates, there are other polymer substrates such as poly(lactic acid) (PLA)[369], poly(styrene-b-(ethylene-co-butylene)-b-styrene) (SEBS) [371], polyethersulfone (PES) [387], poly(vinylidene fluo-ride) (PVDF) [388], PU [393], polycarbonateurethane (PCU) [394], and inorganic substrates such as silicon wafer [389–391]. Photo-induced graft polymerization can be combined with ATRP [395] and NMP [396]. The combined method providesa straightforward way to prepare block or comb-like copolymers with a controlled architecture [395,396].

Despite the aforementioned advantages, photo-induced graft polymerization should be further improved [366]. There isstill no direct technique to characterize parameters such as the length of the grafted chains and their distribution, nature ofthe branched or superbranched graft polymer chains, and crosslinking structure. These parameters are important to theeffectiveness of the process. There are other problems associated with the reproducibility of grafting, degradation, and rear-rangement. Owing to high-energy ion bombardment or UV radiation, most polymers undergo inevitable degradation. Rear-rangement of the grafted functional groups can lead to substrate degradation common to both plasma- and photo-inducedgrafting polymerization. The last problem is the microenvironmental effect. Because the polymer chains covalently graftedonto the substrate should behave differently than the free polymer chains in the solution or bulk, the performances of thegrafted layers should be experimentally examined and compared with the corresponding free polymer chains.


6. Biocompatibility of polymerized surfaces

Since biomaterials are in direct contact with the biological media, the surface must be biocompatible, that is, not inducingadverse response from the body. Williams proposed the definition of biocompatibility as the ability of a material to performwith an appropriate host response in a specific application [397]. Therefore, the suitability of these polymers after surfacepolymerization must be assessed in terms of the potential toxicity and detrimental effects in the physiological environmentbesides the biological performance of the biomaterials. In general, implanted biomaterials often induce a foreign body reac-tion consisting of three steps starting with adsorption of nonspecific proteins onto the surface of implants, followed byattachment of different types of cells onto the surface inducing up-regulation of cytokines and subsequent pro-inflammatory processes, and finally formation of giant cells and cytokine release [398]. The biocompatibility of polymerizedsurfaces is often evaluated by cytotoxicity and blood compatibility closely related to protein adsorption and platelets adhe-sion in vitro and in vivo [399].

6.1. Cytotoxicity

The cytotoxicity of polymerized surfaces is determined by the nature of the polymers including molecular weight, chargedensity, type of cationic functionalities, structure, and sequence (i.e. block, random, linear, branched, and conformationalflexibility) [400]. Chen et al. reported that cationic dendrimers were more cytotoxic than anionic or PEGylated dendrimers[401] and Li et al. disclosed that quaternized polyamidoamine (PAMAM)-OH derivatives displayed lower cytotoxicity thanPAMAM-NH2 due to the shielding of the internal cationic charges by surface hydroxyl groups [402]. By means of in vitro incu-bation with HEP G2 cells or 293 human embryonic kidney cells, it was found that the increased charge density and molecularweight increased the toxicity of PAMAM after modifying the surface of PAMAM with lysine or arginine [403]. Scarioti et al.found that the surface charge could be controlled by in situ polymerization of AA on cationic liposomes by changing the con-centration of PAA and adjusting the cytotoxicity of cationic liposomes. In their experiments, completely coated liposomesexhibited better mouse fibroblasts viability [404]. On the polymerized surface, direct contact with negatively-charged cellmembrane proteins and phospholipids with cationic polymers disturbed the membrane structure and functions due to ery-throcyte lysis [405,406].

Polymerization of PLLA, PAA or PHEMA on biomaterials not only modulated the surface hydrophilicity but also providedfunctional groups for surface bioconjugation consequently offering a more favorable biological environment for cell adhesionand growth on the surface on inorganic [407] and organic substrates [408]. Multihydroxyl hyperbranched polyglycerol (HPG)had good biocompatibility. It could decrease the cytotoxicity of CdTe quantum dots (QDs) indicating that surface polymer-ization of biocompatible dendritic polymers could be employed to obtain robust nontoxic functionalized QDs [409]. Otherpolycations polymers like diethylaminoethyl (DEAE)-dextran [400], PLL [405], dendrimers [410] and PEI [411] also exhibitedincreasing cytotoxicity as a function of molecular weights but it was applicable to polymers with different structures. Forinstance, Fischer et al. showed that poly(l-lysine hydro bromide (PLL) with 36.6 kDa influenced the cell viability more thancationized human serum albumin (cHSA) with 67 kDa and the 500 kDa DEAE-dextran [400].

6.2. Hemocompatibility

Protein adsorption plays a crucial role in vivo pertaining to the initial formation of thrombus at the blood-material inter-face because it is the initial contact between blood and biomaterials that determines subsequent platelet adhesion and acti-vation [412]. Hence, control of protein adsorption is crucial to the design of surface polymerization. Yuan et al. employedATRP to obtain block copolymers with various compositions comprising of PMPC and poly(3-methacryloxypropyltrimethoxysilane) (PMTSi) segments as coatings on cellulose membranes (CMs) [413]. The surface properties such as the sur-face contact angle, surface morphology, and number of functional phosphorylcholine groups were changed by controllingthe density of the functional MPC by adjusting the ratio of the two monomers (MPC and MTSi) and their results disclosedthat protein adsorption depended on the MPC density and PMPC chain length [413]. The surface grafted polyethylene (g-PE) films with different types of water soluble polymers such as PMPC, poly[2-(glucosyloxy)ethyl methacrylate] (PGEMA),poly(N-isopropylacrylamide) (PNIPAM) and poly[N-(2-hydroxypropyl) methacrylamide] (PHPMA) exhibited different hemo-compatibility, and PMPC-g-PE film showed the best biocompatibility among the g-PE films because its surface adsorbed lessprotein than the untreated PE and other g-PE films. Consequently, it showed the highest inhibitory effect of thrombin activityamong the water soluble polymers due to the formation of a hydrogel structure by the PMPC segments [412]. The MPC moi-ety was believed to have a weak interaction with water molecules and a small surface active ability [412]. The chain length,molecular weight, and density of the grafted polymers on the biomaterials also affect platelet adhesion. For insistence, Fenget al. conducted ultraviolet initiated photopolymerization to immobilize poly(ethylene glycol) monoacrylates (PEGMAs)with molecular weights between 400 and 1000 g mol�1 onto PCU to increase the hydrophilicity and improve the hemocom-patibility [394]. The surface-grafted films were hydrophilic and the brush-like structure of the PEG chains resisted plateletadsorption more than the unmodified film. The PCU-g-PEGMAwith 800 g mol�1 exhibited the lowest platelet adsorption andbest hemocompatibility. It might be ascribed to the optimal balance between the PEGMA immobilization density and PEGchain length as well as the well distributed brush structure with a more homogenous topography [394]. In the case of


the polypropylene non-woven (NWFPP) membrane modified by 2-acrylamido-2-methylpropane sulfonic acid (AMPS) by theO2 plasma pretreatment and UV-induced photografting, the hydrophilicity of the membranes increased with grafting den-sities of the poly(AMPS) from 0 to 884.2 g cm�2 and protein adsorption decreased with grafting densities of poly(AMPS).There is reduction in the number of platelets adhering to the modified membrane compared to the unmodified membraneattributable to incorporation of anionic sulfonic acid groups into the membrane surface due to repulsion with the negatively-charged platelets [414]. Similarly, polymers such as hyaluronic acid (HA) can be used as the polymerization source to resistplatelet adhesion and improve cell-substrate interactions [415].

In practice, biocompatibility depends on many reasons. With regard to surface polymerized biomaterials, the physiolog-ical environment typically induces abrasion, corrosion, or degradation of the polymers in turn influencing the host response.The biomaterials should be sterilized before clinical use by methods such as high-temperature sterilization, radiation bygamma ray and electron beams, chemical sterilization, and so on. In the biological environment, adsorption of water canchange the properties of surface polymers. During sterilization, the high temperature and radiation can alter the perfor-mance of the biopolymers. For example, polyesters easily degrade when exposed to water due to the formation of oligomerswhich may invoke inflammation in the body [398]. Another problem is that the excess heat produced during polymerizationfrom exothermic reactions may influence the behavior of cells and tissues [411]. In addition, the volume expansion or shrink-age as a result of the heating may contribute to interfacial failure between the biopolymers and tissues.

6.3. Manipulation of cell behaviors

Apart from the aformentioned cytocompatibility and hemocompatibility, next-generation biomaterials have been endow-ing some ‘‘smart” biological functions, i.e., precise manipulation of cell behaviors according to the desired setting to accel-erate the tissue repair or reconstruction or through the specific surface design, which will be required and incorporated intothe concept of biocompatibility in near future. Due to their similar molecule structures to the basic organization of naturaltissues, some peptide ligands such as arginine-glycine-aspartic acid (RGD) [416] and Arg-Glu-Asp-Val (REDV) [417] havebeen employed to modify the surface of biomaterials to mimic the natural extracellular matrix (ECM), and thus to regulatethe cell adhesion [418].

It has been demonstrated that the density and conformational changes of peptides, and as well as the coupling strengthbetween the peptide and the substrate can regulate cell behaviors [419–421]. Chollet et al. investigated the behaviors of dif-ferent types of cells on the PET substrates modified by the covalent grafting of different densities of RGD containing peptides(varying between 0.6 and 2.4 pmol/mm2), and their cell attachment assays showed that a minimal RGDC peptide density was1 pmol/mm2 for improving the osteoblastic cells (MC3T3) and endothelial cells (HSVEC) responses [419]. The number of focalcontacts ciould be enhanced by increasing the grafted RGDC density on PET, and the largest number of adherent MC3T3 andHSVEC couldbe observedonRGDC-modified PETwith aRGDdensity of 2.4 and1.7 pmol/mm2, respectively [419]. These resultsdisclosed that the density of peptides immobilized on the surface of biomaterials was an important parameter influencing thecell behaviors. Recently, an electrically switchable surfacewas fabricated, which relied on electrically-induced conformationalchangeswithin surface-grafted RGD oligopeptides as themeans ofmodulating cell adhesion (Fig. 40) [420]. The charged back-bone of the oligopeptide has been proved that it can be potentially harnessed to induce its folding on the surface upon an appli-cation of an electrical potential [422,423]. The RGD modified surface was comprised of a mixed self-assembled monolayer(SAM) consisting of two components: an oligopeptide with three lysine residues and a glycine-arginine-glycine-aspartate-serine (GRGDS), and an ethylene glycol-terminated thiol (C11TEG) [420]. The results showed that conformational changes coulddynamically regulate cell adhesion by selectively exposing under open circuit (OC) conditions (bio-active state) or concealingundernegativepotential (bio-inactive state) theRGDto the cell [420]. Bianand co-workersdevelopeda facile platformto inves-

Fig. 40. Schematic of the dynamic RDG oligopeptide SAM utilized for controlling specific cellular interactions. The electrically switchable SAM exposes theRGD peptide and supports cell adhesion under open circuit (OC) conditions (no applied potential), while under an applied negative potential the RGD isconcealed, inhibiting cell adhesion [420]. Copyright 2014. Adapted with permission from the Royal Society of Chemistry.

Fig. 41. Representative fluorescence micrographs of the hMSCs cultured on LCS0.5, LCS12.5, HCS25, and HCS50 for 24 h. Coverslips treated with piranhaonly is used as the control for all the cell experiments. High-magnification images are taken to show the morphological feature of cells grown on thefabricated surfaces. Scale bar is 100 lm [421]. Copyright 2015. Adapted with permission from the American Chemical Society.


tigate the effects of varied coupling strengthbetween theRGDpeptide and the glass substrate on stemcell behaviors,where thecoupling strength of an integrin-binding ligand RGDwas controlled solely by electrostatic interactions between the positivelycharged aminated glass surface and the negatively charged citrate-capped gold nanoparticles (cit-AuNPs) [421]. A total of 0.5%and 12.5% of (3-aminopropyl) triethoxysilane (APTES) were used to fabricate low-coupling-strength surfaces (namely, LCS0.5and LCS12.5), whereas 25% and 50% of APTES were used to fabricate high-coupling-strength surfaces (namely, HCS25 andHCS50). It revealed that humanmesenchymal stemcells (hMSCs) grownonboth LCS0.5 and LCS12.5 surfaceswerehighly elon-gated and spindle like in shape, while hMSCs were widely spread on both HCS surfaces, especially for cells grown on LCS50(Fig. 41), which suggested that the coupling strength of the RGDpeptide on the substrate couldmodulate the adhesion, spread-ing, and differentiation of hMSCs [421].

Besides peptides, there are some other polymers that can be grafted onto the surface of biomaterials to control cellbehaviors, such as PAA [424], PAAm [425], PNIPAAm [426], and poly(3-sulfopropylmethacrylate) (PSPMA) [427]. The graftdensity of polymers can also regulate cell behaviors [424,425]. For example, different density of ACOOH was introducedon PEEK surface using plasma polymerization of AA, and the results exhibited that the adhesion and proliferation of pre-osteoblasts decreased with increasing ACOOH surface density, whereas the spreading of pre-osteoblasts increased withincreasing ACOOH surface density [424]. Lilge and Schönherr investigated the effects of varied graft density of PAAm pre-pared by surface-initiated atom transfer radical polymerization (SI-ATRP) on cell attachment and spreading, and their resultsdisplayed that cells attached only on low density PAAm brushes, while polymer brushes with higher grafting densityremained cell resistant [425]. The nanoarchitectonics of PNIPAAm brushes could also influence cell adhesion behavior[426]. Amin and co-workers reported the exploitation of directly photografting polymer on polydopamine (PDA) nanosheetsfor controlling cell adhesion [427]. Cells grew and spread over the as-grafted PDA layer, while about 44% more cells grew,attached and spread on a PDA nanosheet compared to as-grafted PDA layer, which indicated that PDA nanosheet promotedmore cells growth and attachment [427]. Furthermore, after grafting negatively charged polymer brush such as PSPMA on aPDA nanosheet, significantly fewer cells grown on PSPMA carpet, which suggested that the PSPMA carpet inhibited cellgrowth and attachment [427].

7. Conclusion and outlook

Recent advances in surface engineering and biomaterials enable functionalization by polymer grafting techniques toimprove the biological functions at the interface between tissues/cells and biomaterials. This review summarizes thecommon grafting techniques with a focus on controlled/living radical polymerization (CLRP), a rapidly developing techniquethat yields polymers with controlled molecular weight and low polydispersities. Of the three main approaches, NMP, ATRPand RAFT polymerization, NMP is developed first but ATRP and RAFT are the most common methods in surface polymeriza-


tion of biomaterials. Recently developed methods for surface functionalization of biomaterials such as plasma-induced graftpolymerization, radiation-induced graft polymerization, and photo-induced graft polymerization are described. Althoughgrafting polymerization methods are promising tools to enhancing surface biological functions such as antibacterial activityand bioactivity, more investigations are needed in order to fathom the mechanisms and realize their full potential. For exam-ple, future work should focus on achieving precise control of grafting onto different types of substrates with different shape,understanding the relationship between specific biological functions and grafted molecular architectures on the surface, andachieving multi-functionalities by simultaneous grafting of different molecules.

Acknowledgements

This work is jointly supported by Special Prophase Program for Key Basic Research of the Ministry of Science and Tech-nology of China (973 Program) No. 2014CB660809, the National Natural Science Foundation of China, Nos. 51422102 and81271715, Hong Kong Research Grants Council (RGC) General Research Funds (GRF) Nos. 112212 and 11301215, as wellas Hubei Provincial Natural Science Foundation (China) Nos. 2013CFA018 and 2014CFB551.

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