Responsive polymers in controlled drug delivery

Progress in Polymer Science 33 (2008) 1088–1118

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Progress in Polymer Science journal homepage: www.elsevier.com/locate/ppolysci

Responsive polymers in controlled drug delivery A.K. Bajpai ∗ , Sandeep K. Shukla, Smitha Bhanu, Sanjana Kankane Bose Memorial Research Laboratory, Department of Chemistry, Government Autonomous Science College, Jabalpur 482 001, M.P., India

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Article history: Received 1 November 2007 Received in revised form 1 July 2008 Accepted 30 July 2008 Available online 16 September 2008 Keywords: Responsive polymers Hydrogels Drug delivery

a b s t r a c t This article reviews the state-of-the art in responsive polymer systems for controlled drug delivery applications. The paper describes different types of stimuli-sensitive systems and gives an account of their synthesis through methods such as group transfer polymerization, atom transfer radical polymerization and reversible addition-fragmentation chain transfer polymerization. The article also discusses classification of various drug delivery systems: diffusion controlled systems, chemically controlled systems, swelling-controlled systems and modulated release systems. A survey of the recent literature on various stimuliresponsive polymer hydrogels in controlled drug delivery is also included. © 2008 Elsevier Ltd. All rights reserved.

Contents 1. 2.

Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Responsive stimuli-sensitive materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.1. Polymer blends . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.2. Interpolymer complexes . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.3. Interpenetrating polymer networks . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.3.1. Classifications of interpenetrating networks . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.3.2. Supramolecular morphology: polymer brushes . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.3.3. Supramolecular morphology: nanoparticles . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.4. Block copolymers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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Abbreviations: ␣CDS, ␣-cyclodextrins; AA, acrylic acid; AMPSA, 2-acrylamide-2-methyl-1-propane sulfonic acid; ATRP, atom transfer radical polymerization; CBDS, 2-cyano-2-butyl dithiobenzoate; CDB, cumyl dithiobenzoate; CDDS, controlled drug delivery system; CPDB, 2-cyano-prop-2-yl dithibenzoate; CP, carboxy vinyl polymer; CPS, conducting polymers; CS, chitosan; DEA, 2-(dimethylamino) ethyl methacryl; DMA, 2-(dimethylamino) ethyl methacrylate; DSC, differential scanning calorimetry; EGDMA, ethylene glycol dimethacrylate; ETPFU, tetrahydrophthalimide-5-fluorouracil; FTIR, Fourier transmission infrared; GMA, glicidyl methacrylate; GTP, group transfer polymerization; HA, hyaluranate; HEA, hydroxylethyl acrylate; HEMA, 2-hydroxy ethyl methacrylate; HFA, heptadecafluorodecylacrylate; IPC, Iinterpolymer complexes; IPN, interpenetrating polymer networks; LCST, lower critical solution temperature; MA-Inulin, methacrylated inulin; MBA, N,N -methylene bisacrylamide; MC, methyl cellulose; MDIC, macrodiisocyanate; ODIC, oligodiisocyanate; PAA, poly(acrylic acid); PAAm, polyacrylamide; PAMS, poly(acrylamide-co-styrene); PCL, poly(␧-caprolactone); PCLA, poly(␧-caprolacone-co-lactide); PDMAM, poly(N,N-dimethylacrylamide); PEG, polyethylene glycol; PEO, polyethylene oxide; PFA, pentafluoropropylacrylate; PFS, pentafluorostyrene; PHE, antipyretic phenacetin; PHEMA, poly(2-hydroxyethyl methacrylate; PLA, poly(d,l-lactic acid); PLGA, poly(d,l-lactic acid-co-glycolic acid); PLLA, poly(llactide); PMAA, poly(methacrylic acid); PMMA, poly(methylmethacrylate); PNIPAAm, poly-N-isoproply acrylamide; PnBuA, poly(n-butyl acrylate); PST, polystyrene; PVA, poly(vinyl alcohol); PVP, polyvinyl pyrrolidone; RAFTP, reversible addition-fragmentation transfer polymerization; SCK, shell crosslinked knedel; (SEVA-C), corn starch/ethylene-co-vinyl alcohol; SPH, superporous hydrogel; SSM, stimuli-sensitive material raft; tBuA, tert-butylacrylate; tBuMA, tert-butyl methacrylate; TC, tetracycline; THF, tetrahydrofuran; TFA, trifluoroethylacrylate; TMSMA, trimethylsilyl methacrylate; UCST, upper critical solution temperature. ∗ Corresponding author. E-mail addresses: [email protected], [email protected] (A.K. Bajpai). 0079-6700/$ – see front matter © 2008 Elsevier Ltd. All rights reserved. doi:10.1016/j.progpolymsci.2008.07.005

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2.4.1. General considerations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.4.2. Preparative methods of polymerization for block copolymers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Responsive polymers in controlled drug delivery . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.1. Drug delivery profiles and systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.2. Diffusion-controlled systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.2.1. System types . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.2.2. Reservoir (membrane) systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.2.3. Matrix (monolithic) systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.3. Chemically controlled systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.3.1. Bioerodible and biodegradable systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.3.2. Pendant chain systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.4. Swelling-controlled systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.5. Modulated-release systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.6. Temperature-sensitive release systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.6.1. Negative thermosensitive hydrogels . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.6.2. Positive thermosensitive hydrogels . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.7. pH sensitive release systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.7.1. Cationic hydrogels . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.7.2. Anionic hydrogels . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.8. Magnetic-sensitive release systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.9. Electric-sensitive release systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Challenges and prospects . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Conclusions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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1. Introduction

2. Responsive stimuli-sensitive materials

Hydrogels are three-dimensional high-molecular weight networks composed of a polymer backbone, water and a crosslinking agent. They are gaining tremendous importance in a wide variety of applications in medical, pharmaceutical and related fields, e.g. wound dressings [1], contact lenses [2], artificial organs and drug delivery systems [3]. Hydrogels are polymeric materials that do not dissolve in water at physiological temperature and pH. They swell considerably in an aqueous medium [4] and demonstrate extraordinary capacity (>20%) for imbibing water into the network structure. Gels exhibiting a phase transition in response to change in external conditions such as pH, ionic strength, temperature and electric currents are known as “stimuli-responsive” or “smart” gels [5]. Being insoluble, these three-dimensional hydrophilic networks can retain a large amount of water that not only contributes to their good blood compatibility but also maintains a certain degree of structural integrity and elasticity [6]. Hydrophilic functional groups such as –OH, –COOH, –CONH2 , and –SO3 H present in the hydrogel are capable of absorbing water without undergoing dissolution. Hydrogels can be prepared from natural or synthetic polymers [7]. Although hydrogels made from natural polymers may not provide sufficient mechanical strength and may contain pathogens or evoke immune/inflammatory responses, they do offer several advantageous properties such as inherent biocompatibility, biodegradability and biologically recognizable moieties that support cellular activities. Synthetic hydrogels, on the other hand, do not possess these inherent bioactive properties. Fortunately, synthetic polymers usually have well-defined structures that can be modified to yield tailored degradability and functionality [8].

Hydrogels have been developed as stimuli-responsive materials, which can undergo abrupt volume change in response to small changes in environmental parameters: temperature, pH, ionic strength, etc. (Fig. 1). These unique characteristics of hydrogels are of great interest in drug delivery, cell encapsulation and tissue engineering [9–12]. Stimuli-responsive polymers play an important role in the development of novel smart hydrogels [13]. The most important systems from a biomedical point of view are those sensitive to temperature and/or pH of the surroundings. The human body exhibits variations of pH along the gastrointestinal tract, and also in some specific areas like certain tissues (and tumoral areas) and subcellular compartments. Polymer–polymer and polymer–solvent interactions show an abrupt readjustment in small ranges of pH or temperature. This is attributed to a chain transition between extended and compacted coil states. In the case of pH sensitive polymers, the key element of the system is the presence of ionizable weak acidic or basic moieties attached to a hydrophobic backbone. Upon ionization, the coiled chains

Fig. 1. Intelligent, stimuli-responsive hydrogels, modulated release of drug (circles).

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extend dramatically, responding to the electrostatic repulsions of the generated charges (anions or cations). Thermosensitive polymers, like pH responsive systems, offer many possibilities in biomedicine. They present a fine hydrophobic–hydrophilic balance in their structure; and small temperature changes around a critical solution temperature (CST) make the chains collapse or extend, responding to adjustments of the hydrophobic and hydrophilic interactions between the polymer chains and the aqueous medium [14–16]. A critical solution temperature may be defined as a temperature at which the polymer solution undergoes separation from one phase to two phases. Thus, temperature sensitive polymers undergo an abrupt change in volume as the temperature of the medium is varied above or below the CST [17]. These unique characteristics make hydrogels especially useful in biomedical applications such as controlled release of drugs and in tissue engineering [18–22]. Stimuli-responsive sensitive polymer gels offer potential economic alternatives to conventional separation processes for industrial applications [23]. Controlled permeability variations of responsive gels have also been used to achieve a variety of size- or charge-selective separations. In addition to pH and temperature, other stimuli-responsive hydrogels have been produced that exhibit dramatic changes in their swelling behavior, network structure, permeability and mechanical strength in response to a number of external stimuli, including the presence of specific solutes and applied electrical or magnetic fields [24]. pH sensitive polymers are normally produced by adding pendant acidic or basic functional groups to the polymer backbone; these either accept or release protons in response to appropriate pH and ionic strength changes in aqueous media [25]. The network porosity of these hydrogels changes with electrostatic repulsion. Ionic hydrogels containing carboxylic or sulfonic acid groups show either sudden or gradual changes in their dynamic and equilibrium swelling behavior as a result of changing the external pH. The degree of ionization of these hydrogels depends on the number of pendant acidic groups in the hydrogel, which results in increased electrostatic repulsions between negatively charged carboxyl groups on different chains. This, in turn, results in increased hydrophilicity of the network and greater swelling ratio at high pH. Conversely, hydrogels containing basic pendant groups, such as amines, ionize and show electrostatic repulsion at low pH [26]. Another novel type of responsive polymers results from surface modification of a polymer matrix by attachment of responsive chains to produce responsive interfaces showing different behavior in response to small changes in environmental parameters. Surfaces may change from hydrophobic to hydrophilic [27] or show a variation in pore size [28]. The recent past has witnessed increasing interest in the use of polymer hydrogels in biotechnological applications. Appropriate electric stimulus systems may modify the swelling capacity of hydrogels, thus, inducing specific properties. Moreover, this also affects their mechanical properties and morphology—recognized parameters that correlate hydrogel behavior and end applications. For these

reasons, particular consideration has been paid to polymers that respond to appropriate stimuli [29]. Hydrogels, in conjunction with conducting polymers (CP), form materials capable of undergoing chemical and/or physical transition in response to appropriate electrical field stimuli [30]. It is possible to produce environmentally sensitive or “responsive” hydrogels, whose physical–chemical properties such as surface topography and rheological properties, vary in response to specific environmental stimuli [31]. 2.1. Polymer blends The blending of hydrophilic–hydrophobic polymers produces phase-separated composite hydrogels. Blending different polymers and yet conserving their individual properties in the final mixture is an extremely attractive inexpensive and advantageous way of obtaining new structural materials [32]. The resulting polymer blends show synergistic properties. The advantages of polymer blend systems for controlled release applications may include easy fabrication of devices, manipulation of device properties (hydration, degradation rate and mechanical strength), drug loading and utilization of the dispersed phase domains as microreservoirs for enhanced release properties. Some natural biodegradable polymers, such as carbohydrates, can be used in several applications, such as films for packaging and agriculture. Unfortunately, in the majority of cases, the properties of a natural polymer do not fit the needs of a specific application and blending with synthetic polymer is used to gain the desired properties [33]. For these specific applications, natural polymers such as agar, starches, alginates, pectins and cellulose derivatives are used along with synthetic biodegradable polymers such as polycaprolactone, polylactide and poly(vinyl alcohol) [34,35]. Binary polymeric blends [36] of crosslinked starch and gelatin have been prepared, characterized and examined for enzymatic degradation, using ␣-amylase as a degrading enzyme. Blends of conductive polymers with conventional non-conductive ones have also been produced to achieve good electric conductivity and good mechanical properties for electronic and industrial applications. Blends of PVA (poly(vinyl alcohol)) and sodium polystyrene sulfonate were prepared through aqueous solution casting. The prepared blend films were flexible and transparent with conductivity 10−5 S/cm [37]. Preparation and characterization studies of binary grafted polymeric blends of poly(vinyl alcohol) and gelatin were performed and the resulting hydrogels were assessed for water sorption characteristics. The influence of various factors such as composition of the blend, pH and temperature of the swelling bath and presence of electrolyte on the water uptake potential of blended hydrogel [38] was investigated. More recently, there has been increasing interest in starch based biodegradable blends [39,40], offering a method to modify both properties and degradation rates. A hypothetical scheme showing the formation of a binary blend from its constituent polymers is shown in Fig. 2. The two common methods used to prepare blend polymers are melt blending and solution casting (or solvent casting).

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Fig. 2. Hypothetical scheme depicting preparation of polymer blend.

Melt blending [41] is done at high temperature in the melt or near melt state by using shearing force. Melt blending can be achieved by extrusion with blowing agents and compression molding in the extrusion process. Solution blending [42] can be carried out by dissolving the two components in a common solvent and precipitating out the blend by addition of a suitable precipitant—or by removal of the solvent by evaporation, for example, in cast films. 2.2. Interpolymer complexes Interpolymer complexes, also known as polycomplexes, are insoluble macromolecular structures formed by the non-covalent association of polymers having affinity for one another. The complexes are formed by association of repeating units on different chains (interpolymer complexes) or on separate regions of the same chain (intrapolymer complexes). Polymer complexes are classified by the nature of the association. The major classes of polymer complexes are stereocomplexes, polyelectrolyte complexes, and hydrogen bonded complexes [43]. One of the most widely studied pairs of complexing polymers is PMAA and PEG. In this system, complexation occurs by the formation of hydrogen bonds between the PMAA carboxyl protons and the PEG ether group. These complexes are highly sensitive to pH of the environment, and exist in solutions at pH low enough to allow for substantial protonation of the PMAA acid groups. The formation of hydrogen-bonded interpolymer complexes between weak polyacids, such as a poly(acrylic acid) (PAA) or poly(methacrylic acid) (PMAA), and proton acceptor polymers, such as polyethyleneoxide [44], polyacrylamides [45], poly(vinyl esters) [46–48], etc. in aqueous solutions, has been widely studied. In general, in aqueous solution of mixtures of such complementary polymers, interpolymer association due to successive hydrogen-bonding between the carboxylic groups of the polyacid and the proton acceptor groups of the polybase leads to the formation of compact interpolymer complexes, soluble only within a narrow pH range. At pH higher than 4.5–5.5 (where the degree of ionization of the weak polyacid is higher than 10–15%) dissociation occurs, while at pH below 3–4 the complexes precipitate.

It is of significance to extend the solubility of the hydrogen-bonding interpolymer complexes, particularly in the low pH region. An anionically charged graft copolymer [poly(acrylic acid-co-2-acrylamide-2-methyl-1-propane sulfonic acid)-g-poly(N,N-dimethylacrylamide)] (P(AAco-AMPSA)-g-PDMAM) has been synthesized by grafting [poly(N,N-dimethylacrylamide)] (PDMAM) chains onto an acrylic acid-co-2-acrylamide-2-methyl-1-propane sulfonic acid copolymer (P(AA-co-AMPSA)) backbone [49]. PDMAM, a water-soluble polymer with important proton acceptor properties, forms hydrogen-bonding interpolymer complexes with PAA [50] that precipitate out even at pH as high as 3.75. The presence of negatively charged AMPSA units in the graft copolymer backbone provides these hydrogen-bonded interpolymer complexes with sufficient hydrophilicity, to assure their solubility, even at low pH, as a result of the hydrogen-bonding interpolymer complexes formed between PAA and the PDMAM side chains of the graft copolymer [51]. The controlled release of PHE, an antipyretic phenacetin, has been studied from a solid dispersion of an interpolymer complex between methylcellulose (MC) and carboxyvinyl polymer (CP). The effect of the MC/CP ratio and molecular weight of MC on the PHE release was studied [52]. 2.3. Interpenetrating polymer networks 2.3.1. Classifications of interpenetrating networks Interpenetrating polymer networks (IPNs) are a class of polymer blends in network form in which at least one component is polymerized and/or crosslinked in the presence of the other [53]. IPNs consist of two or more polymers in network form, held together by permanent entanglements, with no more than a few accidental covalent bonds between chains of the two different types of polymers. IPNs mostly possess enhanced physical properties as compared to the normal polymer blends of their components [54]. In this way, unique properties can be achieved with these biomaterials [55]. For example, combination of pH- and temperature-sensitive polymers [56] and polymers forming polyelectrolyte complexes [57] have been investigated. There are two classification schemes used to describe IPNs: (i) by chemistry and (ii) by structure.

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Fig. 3. Schematic representation of: (a) full IPN, and (b) semi-IPN.

By chemistry: (a) Sequential IPNs: in this type of IPN a network is initially formed by polymerizing a mixture of monomer, crosslinking agent, and initiator (or catalyst) to form a network which is then swollen with the second combination of monomer and crosslinking agent. The resulting swollen network on polymerization gives rise to an IPN. (b) Simultaneous interpenetrating networks (SINs): an IPN is formed by polymerizing two different monomers and crosslinking agent pairs together in one step. The key to the success of this process is that the two components must polymerize by reactions that will not interfere with one another. Fig. 3(a) is a schematic representation of an SIN. By structure: (a) Full IPNs: this type of IPN comprises two networks that are ideally juxtaposed, which generates many entanglements and interactions between the networks. (b) Homo-IPNs: these IPNs are a special type of full-IPN, where the two polymers used in the networks are the same. (c) Semi- or pseudo-IPNs: one of the components of the IPN has a linear instead of a network structure. A molecular depiction of this structure is shown in Fig. 3(b). (d) Latex IPNs: these IPNs are formed by emulsion polymerization. The morphology of the IPN depends upon how the IPN components are polymerized. (e) Thermoplastic IPNs: these IPNs are moldable, can be extruded and recycled. At least one component is usually a block copolymer. IPN hydrogels may have more favorable mechanical properties due to the physical entanglements as compared with individual crosslinked networks [58]. Generally, most methacrylate polymers are hard and brittle and have only a limited field of application. One successful method of improving their toughness and mechanical properties is to incorporate judicious qualities of a second

phase. Constituting the dispersed elastomeric particles, into the hard polymethacrylate matrix. This aim is best achieved by the synthesis of IPNs based on two components. The addition of such an elastomeric component into the hard polymer matrix has been shown to enhance toughness, mechanical properties and thermal stability while at the same time lowering the glass transition temperature. As a part of free radical polymerization studies, the synthesis of sequential IPNs from glycerol-modified castor oil polyurethane and poly(2-hydroxyethyl methacrylate) was reported. Super-porous IPNs, having enhanced mechanical properties were synthesized [59]. The main approach used in this study was to form an interpenetrating polymer network by incorporating a second polymer network inside a super-porous hydrogel (SPH) structure. Polyacroylonitrile was used as a second network inside an SPH. Mechanical properties including compressive strength and elasticity were significantly improved up to 50-fold times as compared with a control SPH. The enhanced mechanical properties were a result of the scaffold fiber network structures formed inside the cell walls of SPHs. The fast swelling property of SPHs was not affected by the incorporation of the second polymer network because the interconnected pore structure was maintained. Gastric retention devices based on superporous IPN with the improved mechanical properties are expected to withstand compression pressure and mechanical friction in the stomach better than control SPHs. Bajpai studied the adsorption of fibrinogen on the surface of semi-interpenetrating polymer networks (IPNs) of PEG and PHEMA [60]. Bajpai and Mishra prepared and characterized tetracycline (TC)-loaded ionic (IPNs) of carboxymethyl cellulose (CMC) and poly(acrylic acid). The entrapped drug was monitored by antibacterial activity and structural integrity [61] and release profiles were investigated. Kato et al. [62] synthesized magnetically activated gels showing expansion in an alternating magnetic field. They prepared IPN gels by a sequential IPN synthesis in which ␥-Fe2 O3 -loaded poly-AAm gel was crosslinked by MBA to form an initial polymer-network. Then, a polyacrylic acid

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network was synthesized within the polyAAm gel and the resulting ␥-Fe2 O3 -loaded IPN gel was heated magnetically. Junyan et al. [63] synthesized a Cd (II) imprinted interpenetrating polymer network (IPN) containing epoxy resin, triethylenetetramine and cadmium methacrylateacrylamide-N,N -methylene-bis-(acrylamide) by in situ sequential polymerization. The ion-imprinted IPN adsorbent was successfully applied to the analysis of two natural water samples. 2.3.2. Supramolecular morphology: polymer brushes Biomaterials may be differentiated according to supramolecular structures that have been developed for application in medical devices. Examples are polymer brushes and nanoparticle structures. Polymer brushes are polymer monolayers attached to a solid support [64]. Surfaces formed from polymer brushes have potential applications as low friction biological surfaces [65], nanoparticles organizers [66], and nanoparticles stabilizers for NMR contrast agents [67], biological mimetic and antifouling surfaces [68] and high dielectric polymer coatings [69]. At high grafting density, adsorbed polymers change their conformation from a globule to an extended chain as shown in Fig. 4. This change in conformation is induced by polymer–polymer repulsions when polymer chains are confined to a small area. This situation was first described by Alexander in 1977 and deGennes in 1980, but it was not until 1986 that Tirrell [70] showed experimental evidence for the extended chain conformation. Tirrell’s synthesis is typical of the “grafting to” approach where polymer chains are first synthesized followed by surface adsorption. The attached chains cover the surface at high density and prevent further chain adsorption. Another approach is the “grafting from” method [71]. Here, polymer initiators are attached to a surface, and polymerization proceeds from the surface with the addition of monomer.

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The reactive sites move with the growing chain, allowing complete access of monomer to the initiator site. Polymer brushes can interact with the surrounding solvent and expand and contract, depending on solvent quality. This allows brushes to act as nanoactuators, opening and closing “valves” in response to various conditions. Polymer brushes have been used to synthesize porous membranes that can open and close in response to pH [72], light [73] and temperature [74]. Sato et al. [75] prepared a series of cationic comb-type copolymers (CCCs) possessing a polycationic backbone and abundant water-soluble side chains as delivery systems for small interfering RNA (siRNA) with prolonged blood circulation time. Markedly, the CCC with higher side chain content (10 wt% PLL and 90 wt% PEG) showed stronger interaction with siRNA than CCC with lower content (30 wt% PLL and 70 wt% PEG)—showing that a very dense PEG brush reinforces the interpolyelectrolyte complex between the PLL backbone end siRNA. Peptide nanospheres with PEG brushes possess both stealth properties and chemical functionality for drug delivery [76] systems and biological diagnosis. They are prepared by one-step polymerization of lphenylalanine (Phe) N-carboxyanhydride with the dual initiators hydrophobic n-butylamine and hydrophilic NH2 -monoterminated PEG (NH2 -PEG). The high density PEG brush conformation of the peptide nanospheres was confirmed by 1 H NMR. Target proteins can be covalently immobilized onto the surface of PEG-peptide nanospheres by amide bond formation. Very stable polymer micelles, namely core-surfacecrosslinked nanoparticles (SCNPs) made from amphiphilic brush copolymers, were evaluated as carriers of cisplatin [77], a potent anticancer drug with low solubility in water. Cisplatin encapsulated in SCNPs have much enhanced cytotoxicity to the cancer cells compared to free cisplatin.

Fig. 4. Polymer brushes at three different surface graft densities.

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Poly(N-isopropylamide) (pNIPAAM) alone or as a copolymer is stimulus responsive and has potential applications in molecular motors, drug delivery, sensors and actuation devices. It also undergoes an inverse phase transition triggered by changes in the solvent quality, such as temperature, ionic strength, pH or co-solvent concentration. Clark et al. [78] have shown that microcantilevers with an end-grafted pNIPAAM brush or poly(Nisopropylacrylamide-co-N-vinylimidazole) (pNIPAAM-VI) brush can be used to detect and transduce this phase transition behavior. Changes in the conformational state of the brush induced by phase transition or changes in osmotic pressure cause significant changes in the surface stress in the brush that leads to detectable changes in cantilever deflection. 2.3.3. Supramolecular morphology: nanoparticles Nanoparticles are solid, colloidal particles consisting of macromolecular substances that vary in size from 10 nm to 100 nm. Compared with other colloidal carriers, polymer nanoparticles offer higher stability when in contact with biological fluids, and their polymeric nature allows the desired controlled and sustained drug release to be obtained. They have attractive physiochemical properties such as size, surface potential, hydrophilic–hydrophobic balance, etc., and for this reason they have been recognized as potential drug carriers for bioactive ingredients such as anticancer drugs [79], vaccines, oligonucleotides [80], peptides, etc. Bajpai and Choubey [81] studied swelling-controlled release of sulphamethoxazole from gelatin nanoparticles and investigated the influence of various factors on the release profiles of the drug. Biodegradable polymers often have low toxicity and are not tissue reactive, and thus not require surgical removal from the host. Biodegradable polymer nanoparticles are highly useful because they can be administered at a variety of locations in vivo through a syringe needle [82,83]. A variety of drugs, regardless of their molecular weights and water solubility, can be loaded into biodegradable nanoparticles using different manufacturing techniques [84,85]. Some examples of biodegradable polymers used in nanoparticle preparation include polyester [86], polyanhydrides [87], poly(ortho-esters) [88], polyphosphazenes [89] and polysaccharides. Polymer nanoparticles have shown promise in addressing the challenges encountered in release of macromolecular drugs such as protein by providing a means to encapsulate the protein and release it over an extended period, thereby, potentially reducing the dosage frequency. In addition, it has been demonstrated that protein encapsulation within certain polymers provides protection from proteolytic enzymes to facilitate oral delivery. One immediate challenge is the development of synthetic methodologies that allow preparation of welldefined nanostructures having surface chemistry and internal structure that can be controlled and modified accurately. The synthetic polymers, polylactic acid and polylactic-co-glycolic acid (PLGA) are the two most commonly used. The preparation of well-defined SCK (shell crosslinked Knedel) nanostructured materials that present

biologically active moieties emanating from their surfaces have been accomplished by several synthetic approaches [90–93]. Fig. 5 depicts a chemical route for the synthesis of SCK nanoparticles through polymerization of tert-butyl acrylate and methyl acrylate employing a biotinylated ATRP initiator. 2.4. Block copolymers 2.4.1. General considerations Block copolymers have been studied extensively in recent decades. Traditional amphiphilic block copolymers containing chemically connected hydrophilic and hydrophobic segments provide a great variety of morphologies, both in the solid state and in selective solvents [94,95]. Zwitter-ionic polymers or polyampholytes [96] in which at least two of the blocks are of opposite charge are interesting synthetic analogs for proteins and, therefore, form a novel class of block copolymers [97] of rapidly increasing importance with unique and fascinating properties. In addition to traditional AB- and ABA-type block copolymers; ABC triblock copolymers (or block terpolymers) have also been prepared and are of current research interest because of their characteristic bulk morphologies. Various biodegradable aliphatic polyesters, such as poly(d,l-lactic acid) PLA, poly(d,l-lactic acid-co-glycolic acid) (PLGA), and poly(␧-caprolactone) (PCL), have been chemically modified with hydrophilic poly(ethylene glycol) (PEG) segments to produce thermosensitive, A–B type, diblock or A–B–A triblock copolymers [98,99]. These thermosensitive block copolymers are typically free flowing sols at room temperature, and gels at body temperature, making them promising candidates as injectable hydrogels for drug delivery system and cell therapy. However, the sol–gel transition of thermosensitive block copolymers is not suitable for the injection of hydrogels into deep anatomical sites in the body because of premature gelation inside the microcatheter used to deliver the hydrogel [100]. Shim et al. [101] prepared a novel pH-sensitive and thermosensitive block copolymer by adding the pH-sensitive sulfamethazine oligomer (SMO) to the thermosensitive block copolymer of poly(␧-caprolactoneco-lactide)–PEG–poly(␧-caprolactone-co-lactide) (PCLA–PEG–PCLA). The introduction of the pH-sensitive SMO into the thermosensitive PCLA–PEG–PCLA block provided unique characteristics useful to prevent premature gelation and rapid degradation of thermosensitive aliphatic hydrogels. The sulfonamide modified block copolymer exerted a distinctive sol–gel transition in response to small changes in pH at body temperature [102]. At pH 8.0, the sulfonamide-modified block copolymer solution maintained its sol phase for about 2 h at body temperature, but rapidly formed a gel in physiological conditions (pH 7.4 and 37 ◦ C) within 5 min. Triblock copolymers such as poly(ethylene glycol)– poly(l-lactic acid)–poly(ethylene glycol) (PEG–PLLA–PEG) and block copolymers such as PEG–PLLA block copolymers exhibit a sol–gel transition with decreasing temperature which is influenced by the length of the PLLA block when the PEG block length is constant. These systems have been

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Fig. 5. Synthesis of SCK nanoparticles through ATRP polymerization of tert-butyl acrylate and methyl methacrylate initiated by biotinylated ATRP initiator.

evaluated for the release of hydrophilic or hydrophobic drugs, the release of the hydrophilic one lasting about 2 weeks whereas the hydrophobic one, over 2 months. Degradation of the polymer matrix was slowed by the incorporation of PLGA blocks [103]. Similar systems that include biodegradable segments and involve adjustment of the sol–gel transition are based on poly(ethylene oxide)–PLGA (PEO–PLGA) triblock copolymers. They exhibit sol–gel transitions in aqueous solutions at about 30 ◦ C resulting in the formation of an in situ transparent gel, which maintained structural integrity and mechanical strength [104]. Other systems based on aqueous solutions of PEO-g-PLGA and PLGA-g-PEO form soft gels at 37 ◦ C that are biodegradable and can be applied in tissue engineering. A diblock copolymer containing a polysilane, poly(1,1dimethyl-2,2-dihexyldisilene)-b-PMAA, was prepared by anionic polymerization of disilene and trimethylsilyl methacrylate, followed by hydrolysis of the trimethylsilyl. The crosslinking reaction of the PMAA block with 1,10-diazo-4,7-dioxadecane and 1-ethyl-3(3dimethylaminopropyl) carbodiimide hydrochloride formed shell crosslinked micelles of polysilane. Nanometer-size hollow particles derived from polysilane that will crosslink micelles could be obtained since the polysilane core can undergo photochemical degradation, and it was shown that the hollow particles can undergo reversible uptake of guest molecules [105].

Poly(1,1-diethylsilacyclobutane)-b-PMAA was synthesized by anionic polymerization [106], and the nanostructure of monolayers spread on water was investigated in comparison with poly(␣-methyl-styrene)-b-PMAA [107]. Amphiphilic diblock copolymer micelles [108] have been established as effective carriers for drug delivery applications. Two different approaches can be used to improve the effectiveness of such block copolymer-based drug delivery devices. The first approach involves chemical modification of the amphiphilic block copolymer building blocks. The second approach is the addition of auxiliary agents. Channel proteins and metal (nano) particles have been discussed to improve temporal control over the drug release process under the influence of external stimuli such as IR radiation, light or magnetic fields [109,110]. A range of block copolymers [111] have been developed for various applications including amphiphilic micelles for passive targeting of chemotherapeutic agents and environment-sensitive micelles for the oral delivery of poorly bioavailable compounds. Most amphiphilic copolymers employed for drug delivery contain either a polyester or a poly(amino acid) derivative as the hydrophobic segment. Amino acid-based block copolymers are being studied extensively for drug delivery because of their biodegradability, biocompatibility and structural versatility. By varying the chemical structure of poly(amino acids) (PAA), it is possible to tailor their enzymatic degradability and immunogenicity [112].

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Fig. 6. A scheme of reactions showing atom transfer radical polymerization (ATRP). Poly(acrylic acid) (PAA) was attached on both termini of Pluronic P85 cop.

2.4.2. Preparative methods of polymerization for block copolymers 2.4.2.1. Group transfer polymerization. Group transfer polymerization (GTP) was announced 20 years ago by DuPont as a method for synthesis of acrylic block polymer. GTP utilized 1-methoxy-1-(trimethylsiloxy)-2-methylprop-l-ene (MTS) as initiator and a carboxylic acid salt as catalyst. The number of growing polymer chains corresponds to the amount of MTS used. Chain growth stops when the monomer is depleted. Addition of a new monomer at this point starts chain growth again to produce a block polymer [113]. Butun et al. [114] using GTP, synthesized branched statistical copolymers by copolymerizing either 2-(dimethylamino) ethyl methacrylate (DMA) or 2(diethylamino) ethyl methacrylate (DEA) with ethylene glycol dimethacrylate (EGDMA) in THF at 20 ◦ C. GTP permits good control over both the primary chain length and the molecular weight distribution as compared with branched vinyl polymers synthesized by conventional radical polymerization. Branched diblock copolymers of DMA and DEA were prepared by sequential monomer addition with EGDMA being used to achieve branches in either the DMA block or DEA block or in both blocks. The branched diblock copolymers were characterized using 1 H NMR and their molecular weight distributions were assessed in THF by light scattering. Yamasaki and Patrickios [115] employed GTP to synthesize homopolymers and randomly cross-linked homopolymer networks of 2-(dimethylamino)ethylmethacrylate of various molecular weights The initiator used for polymerization was 1-methoxy-1-(trimethylsiloxy)2-methylpropene while tetrabutylammonium bibenzoate served as the catalyst. Ethylene glycol dimethacrylate was the crosslinker used for the network synthesis at an 8-fold molar excess with respect to the initiator. The authors used gel permeation chromatography THF to characterize the molecular weight distribution of the linear homopolymers. Kadokawa et al. [116] reported the GTP of a cyclic trimethylsilyl dienolate, 2-(trimethyl silyloxy) furan initiated with benzaldehyde. The polymerization was carried out in the presence of a tetrabutylammonium salt as a catalyst THF at 0–50 ◦ C. The product was isolated as an ethyl acetate insoluble fraction. Its structure was further determined by means of 1 H NMR, 13 C NMR and IR spectroscopy. 2.4.2.2. Atom transfer radical polymerization (ATRP). ATRP is a result of the formation of radicals that can grow, but are reversibly deactivated to form dormant species. Reacti-

vation of the dormant species allows the polymer chains to grow again, only to be deactivated later. The radical formation occurs by action of a transition metal catalyst that activates the organic initiator or dormant species by obstructing the halide at the chain end. By ATRP, polymers with controlled molar masses and small polydispersities can be obtained. The mechanism of ATRP is depicted in Fig. 6. Satturwar et al. [117] synthesized block copolymers of poly(ethylene glycol) and t-butyl methacrylate, iso-butyl acrylate, n-butyl acrylate or propyl methacrylate by ATRP. They obtained pH-sensitive micelles by hydrolysis of tbutyl groups. The poorly water soluble drug candersartan cilexetil (CDN) was incorporated in the micelles and the in vitro drug release was studied as a function of pH. The entrapment efficiency of CDN was found to be above 90%. The release of CDN from pH-sensitive micelles increased with increase in pH from 1.2 to 7.2. Thus, this block copolymer synthesized by ATRP can self-assemble to form micelles which exhibit high loading capacities of CDN and release of drug in pH-dependent fashion. Poly(acrylic acid) (PAA) was attached on both termini of Pluronic P85 copolymers via ATRP [118] to produce a novel block copolymer PAA-b-P85-b-PAA (P85PAA). Doxorubicin (DOX) is one of the most common chemotherapeutic drugs with high anti-tumor activity. Physically encapsulated DOX in pluronic micelles enhanced drug delivery to solid tumors and reduced side-effects. The P85PAA-DOX complex formation and drug loading were strongly dependent on the length of the PAA segment and pH. Protonation of carboxyl groups in the PAA segment at pH < 7.3 reduced the binding sites of DOX onto P85PAA chains resulting in a diminished DOX uptake at low pH. Thus, DOX release from the complex is a pH-responsive process where the protonation of carboxyl groups in mildly acidic conditions resulted in a faster dissociation of copolymer-DOX complex, leading to increased release of DOX at pH 5.0. pH-sensitive diblock copolymers [119] of polyethylene glycol and t-butyl methacrylate (tBMA), ethyl acrylate (EA) or n-butylacrylate (nBA) were synthesized by ATRP. The poorly water-soluble drug progesterone (PRG) was incorporated by dialysis or oil-in-water (O/W) emulsion methods. PRG release was evaluated in vitro as a function of pH. The PRG release from supramolecular assemblies increased when the pH of the release medium was raised from 1.2 to 7.2. The results suggest that these supramolecular assemblies with high drug loadings and pH-dependent release kinetics can potentially enhance the oral bioavailability of hydrophobic drugs.

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Controlled polymerization of 2-hydroxyethyl methacrylate (HEMA) was effected by ATRP using activators generated by electron transfer (AGET) [120] in a protic solvent. The AGET process enabled ATRP to be started with an air-stable Cu (II) complex. The level of control in AGET ATRP was similar to that in normal ATRP in protic solvents, and this resulted in a linear increase in the molecular weight with the conversion and a narrow molecular weight distribution. The initiation system AGET [121] is used in ATRP in the presence of a limited amount of air. ATRP of butyl acrylate is successfully carried out in miniemulsion and in the presence of air. Ascorbic acid and tin (II) 2-ethylhexanoate are used as reducing agents. An excess of reducing agent consumes the oxygen present in the system and, therefore, provides a deoxygenated environment for ATRP. The radical concentration remains constant during polymerization. 2.4.2.3. Reversible addition-fragmentation chain transfer polymerization (RAFT). Reversible addition-fragmentation chain transfer radical polymerization allows the synthesis of well-defined macromolecular architectures with relatively low poly-dispersity indices [122,123]. RAFT polymerization achieve their controlled character due to a reversible chain transfer to reduce the amount of radicals and thus, to reduce the occurrence of termination reactions. A conventional RAFT polymerization contains three primary components—a monomer, a radical initiator and a chain transfer (RAFT) agent. In summary, the initiator generates radicals, which propagate into polymeric chains. The RAFT agent “caps” the radicals on the ends of these chains, temporarily stopping propagation. This capping process is reversible. When the RAFT agent is removed from the chains, growth resumes. Since the total amount of time spent capping and uncapping is miniscule relative to that spent propagating, all chains essentially grow incrementally at the same rate [124]. Synthesis of reducible poly(2-dimethylaminoethyl methacrylate) (PDMAEMA) represents promising carriers of therapeutic nucleic acids. Oligomers of (DMAEMA) containing terminal thiol groups were synthesized by RAFT polymerization using a difunctional chain transfer agent [125]. Reducible poly(DMAEMA) (rPDMAEMA) was synthesized by oxidation of terminal thiol groups forming a polymer with a disulfide bond in the backbone. Physicochemical properties of DNA polyplexes of rPDMAEMA were evaluated by dynamic and static light scattering. Cytotoxicity and transfection activity of rPDMAEMA-based DNA polyplexes were evaluated in vitro. Transfection activity was tested in mouse melanoma and six human pancreatic cancer cell lines. rPDMAEMA polyplexes showed better activity than control PDMAEMA polyplexes. Acid-cleavable [126] core–shell-like polymeric colloidal systems for the delivery of hydrophobic drugs at slightly acidic sites were prepared using the acid-labile microgel approach and RAFT-mediated seeded dispersion polymerization. Bisacrylate acetal crosslinker was copolymerized with n-butyl acrylate (BA) in the presence of a RAFT agent, which yielded crosslinked spherical particles with diameters in the range of 150–500 nm. The cleavage of the particles depended on the pH of the medium. In order to

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mask the hydrophobic surface of the particles, polyethylene glycol acrylate (PEG-A) was grafted to P(BA) via a RAFT agent on the particle surface. The acidic-site selective delivery potential of the poly(BA)-g-poly(PEG-A) was assessed in vitro using a lipophilic fluorescent dye as a model drug. The physico-chemical and functional features support the potential value of the acid cleavable poly(BA) core–poly(PEG-A) shell particles as carriers for delivery of hydrophobic drugs at acidic sites. 2.4.2.4. Metathesis polymerization. The olefin metathesis reaction can be thought of as one in which all the carbon–carbon double bonds in an olefin (alkene) are cut and then rearranged in a statistical fashion as shown in Fig. 7(a). If one of the product alkenes is volatile (such as ethylene) or easily removed, then the reaction can be driven completely to the right. Similarly, by using a high pressure of ethylene, internal olefins can be converted to terminal olefins. Ring-opening metathesis polymerization (ROMP) and ring-closing metathesis polymerization (RCM) are the two common types of olefin metathesis. The ROMP reaction uses strained cyclic olefins to produce stereoregular and monodisperse polymers and copolymers. As the reaction involves an alkylidene catalyst and a cyclic olefin, the product olefin that is generated remains attached to the catalyst as a part of the growing polymer chain—as is shown in Fig. 7(b) with a generic strained cyclic olefin. Moreover, the driving force for the ROMP reaction is the relief of ring strain; and therefore, the second step is irreversible. The polymers produced in the ROMP reaction have a very narrow range of molecular weights. Since polydispersities are typically in the range of 1.03–1.10, these polymers are said to be effectively monodisperse. Moreover, ROMP is a superior method for making diblock and triblock copolymers. Ring-closing metathesis is the reverse of ROMP reaction. The rings being formed do not have appreciable ring strain. The RCM reaction sometimes involves running the experiment at high dilution so that most of the reactions are intramolecular rather than intermolecular. Removal of the volatile by-product drives the equilibrium to the ringclosed product. The catalysts are selected to have good reactivity with terminal olefins, but low reactivity with internal ones. Ruthenium-catalyzed olefin metathesis [127] copolymerization was used to synthesize materials structurally equivalent to copolymers of ethylene and vinyl alcohol, vinyl acetate, methyl acrylate and acrylic acid. These polymers serve as superior models for chain-addition functional polyethylenes. The ester group-substituted polymers display rapid decrease in melting temperature and heat of fusion with increasing comonomer contents. The alcohol-substituted polymers, however, show higher melting temperature and a weaker property dependence on comonomer content. Acyclic diene metathesis (ADMET) was used to prepare a series of “N-terminus” amino acid and peptide branched, chiral polyolefins termed bio-olefins [128]. The carboxylic

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Fig. 7. (a) Statistical rearrangement of carbon–carbon double bonds of an olefin in olefin-metathesis reaction. (b) A hypothetical scheme showing the mechanism of ring-opening metathesis polymerization (ROMP).

acid functional groups were protected with methyl, benzyl and tert-butyl esters to enhance both the polymerizability of the monomers themselves and the solubility of the resulting polymers. Many of these polymers are semicrystalline, exhibiting melt transitions up to 132 ◦ C and 74 ◦ C for the amino acid and dipeptide branched polymers. Adding a succinic acid spacer between the amino acid entity and the polymer backbone raised the melting temperature of polymer by 27 ◦ C. Baughman et al. [129] prepared linear ethylene-coacrylic acid (EAA) by a two-mode olefin metathesis polymerization. They used ADMET to prepare three highmolecular weight, high strength EAA materials bearing pendant carboxylic acid groups along the copolymer backbone and used ROMP to create EAA materials of equimolar acid concentrations with irregularly distributed pendant groups along the linear co-polymer backbone. Structure characterization was done by FT-IR, NMR, DSC and X-ray scattering. Wager et al. [130] synthesized precisely defined amphiphilic polyethylene-g-poly(ethylene glycol) copolymers using ADMET polycondensation. The polymers can be either semi-crystalline or completely amorphous. The monomer structures and the corresponding polymer structures were confirmed by 1 H and 13 C NMR, high-resolution mass spectrometry. 3. Responsive polymers in controlled drug delivery

waste of drug and frequent dosing; (5) optimized therapy and better patient compliance; and (6) solution of the drug stability problem. Five controlled release profiles that are highly desirable are shown in Fig. 8 [131]. Profile I: conventional delayed but not constant release. Profile II: constant or zero-order release. Synthetic polymers or pumps deliver drugs at a constant rate so that the drug concentration in the blood stream is maintained at an optimal level of therapeutic effectiveness. These are often referred to as zeroorder drug delivery systems and many have been or are being commercialized to deliver a number of drugs. Profiles I and II are now common in commercial systems. Profile III: substantial delayed release followed by a constant release of active agent. Such systems will be most useful for the delivery of active agents commencing at some period during the night. Profile IV: delay followed by a tight pulse of drug release. This again allows for nocturnal delivery or for the delivery of a hormone, which often requires pulsed rather than constant delivery. Profile V: multiple pulses at specified periods. For example, the World Health Organization requested contract submissions for such systems to deliver pulses of estradiols over 3 days each month for a succession of months.

3.1. Drug delivery profiles and systems Controlled drug delivery can be used to achieve: (1) sustained constant concentration of therapeutically active compounds in the blood with minimum fluctuations; (2) predictable and reproducible release rates over a long period of time; (3) protection of bioactive compounds having a very short half-life; (4) elimination of side-effects,

Controlled (zero-order) drug release is schematically illustrated in Fig. 9. Currently, this may be achieved by various systems. It is known that the drug concentration levels in the blood plasma depend on the quantity of drug released from the device because drug absorption is determined by its solubility in tissues and availability of local blood flow in tissue. Even if the drug concentration in plasma were to

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Fig. 8. Different types of release profiles for drugs (Type I) release rate falling off exponentially with time, (Type II) constant release rate producing a zero order kinetics, (Type III) zero-order release with substantial delay, (Type IV) pulsatile release with delay, (Type V) Multiple release with constant delay in between.

remain reasonably constant, small short-term fluctuation would always be seen due to factors such as physical activity, emotional stimulation (stress), eating and sleeping, etc. [132]. The administration routes of hydrogel-based formulations include transdermal, oral, nasal or parenteral. Drug delivery systems can be classified according to the mechanism controlling the drug release [133]: (1) Diffusion-controlled systems (a) Reservoir (membrane systems) (b) Matrix (monolithic systems) (2) Chemically controlled systems (a) Bioerodible and biodegradable systems (b) Pendent chain systems

(3) Solvent-activated systems (a) Osmotically controlled systems (b) Swelling-controlled systems (4) Modulated-release systems The systems, as classified above, are discussed below; and all these systems are depicted in Fig. 10. Most drug delivery devices act by a combination mechanisms. 3.2. Diffusion-controlled systems 3.2.1. System types In diffusion systems, drugs diffuse through polymer; the polymer may undergo subsequent biodegradation on exhaustion of the drug. The simplest example of a dif-

Fig. 9. A schematic drawing illustrating the controlled (zero-order) drug release.

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Fig. 10. A schematic drawing illustrating the three mechanisms for controlled drug release from a polymer matrix.

fusion system is one in which the drug is enclosed in a reservoir (core) from which it can diffuse. As long as the drug is maintained in the core in a saturated state, release will follow zero-order kinetics until it is nearly exhausted. Two types of diffusion-controlled devices have been used in drug delivery. These are reservoir devices (or laminated matrix devices) and matrix devices as shown in Fig. 11. 3.2.2. Reservoir (membrane) systems Reservoir systems are hollow devices in which an inner core of dissolved, suspended or neat drug is surrounded by a polymer membrane. A schematic representation is shown in Fig. 12.

In this device, the drug core is encapsulated in a polymeric membrane. Drug diffusion through the membrane is rate limiting and controls the overall drug release rate. A saturated concentration of reservoir of the drug inside the reservoir is essential to maintain a constant concentration gradient across the membrane. The drug transport mechanism through the membrane is usually a solution-diffusion mechanism. Drug transport occurs first by dissolution of the drug in the membrane on one side followed by diffusion through the membrane and desorption from the other side of the membrane. A series of membranes prepared by air-drying thin films and composed of PVA blended with chitosan (CS) in differ-

Fig. 11. Schematic drawings of three types of polymer-based diffusion controlled drug delivery devices.

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Fig. 12. Reservoir controlled release device with an inner drug core enclosed by a membrane of thickness ı (slab) and ro –ri (cyc. and sphere).

ent ratios were developed by Sharma and Chandy [134]. These novel membranes demonstrated good permeability properties for small molecules and showed a dramatic reduction in platelet attachment. The prostagladin E1 immobilized substrate also indicated an increase in albumin surface attachment and a reduction in fibrinogen binding. This may be one of the parameters for a reduced platelet-surface attachment, which may also improve the blood compatibility of the substrate. Kim et al. [135] investigated permeation of riboflavin and insulin through PVA and CS membranes. The DSC thermograms of these membranes indicated that the content of free water and the amount of “frozen” bound water increased with the water content in the membrane. A greater permeation rate of solutes in acidic solution was due to an increase in both water content and the amount of free water and bound water. 3.2.3. Matrix (monolithic) systems In matrix systems, the drug is uniformly dissolved or dispersed. An inherent drawback of the matrix systems is their

first-order release behavior with continuously decreasing release rate. This is due to the increasing diffusion path length and the decreasing area at the penetrating diffusion front as the matrix release proceeds. A matrix (or monolith) device is easy to formulate and gives a higher initial release rate than a reservoir device and can be made to release at a nearly constant rate. A monolithic solution device contains drug solution within the polymer, whereas a monolithic dispersion contains dispersed solid drug in a rate-limiting polymer matrix, referred to as a matrix system (Fig. 13). Khan and Zhu prepared controlled-release (CR) matrix tablets of ibuprofen (IBF) and carbopol (R) 934P, and blended mixtures of Carbopol (R) 934P and 971P resins, at different drug to polymer ratios by direct compression [136]. Their investigation focused on the influences of the proportion of matrix material, and several coexpients (lactose, microcrystalline cellulose and starch) on the mechanism and release rate of the drug from the tablets. The release kinetics were modified when blended mixtures of carbopol (R) 934P and 971P resins were used as the matrix materials.

Fig. 13. Matrix delivery device with drug particles dispersed throughout a polymer matrix.

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3.3. Chemically controlled systems In chemically controlled drug delivery systems, the release of a pharmacologically active agent usually takes place in the aqueous environment by one or more of the following mechanisms: (i) Gradual biodegradation of a drug containing polymer system. (ii) Biodegradation of unstable bonds by which the drug is coupled to the polymer system. (iii) Diffusion of a drug from injectable and biodegradable microbeads. 3.3.1. Bioerodible and biodegradable systems In these systems, the polymer erodes because of the presence of hydrolytically or enzymatically labile bonds. As the polymer erodes, the drug is released to the surrounding medium. Erosion may be either surface or bulk erosion. The main advantages of such biodegradable systems are the elimination of the need for surgical removal, their small size and potential low cost. On the other hand, all biodegradable products as well as their metabolites must be non-toxic, non-carcinogenic and non-teratologic. These requirements are not easily met and must be subject to careful scrutiny. The design and preparation of novel bioerodible hydrogels (Fig. 14) developed by free radical polymerization of acrylamide and acrylic acid and some formulations with bis-acrylamide, in the presence of a corn starch/ethyleneco-vinyl alcohol copolymer blend (SEVA-C), is reported for drug delivery [137]. Elvira et al. [137] performed swelling studies as a function of pH in different buffer solutions to determine the water-transport mechanism that governs the swelling behavior. They performed degradation studies of the hydrogels in simulated physiological solutions for times up to 90 days, determining the respective weight-loss, and analyzing the solution residue by 1 H NMR. They characterized mechanical properties of the xerogels by tensile and compressive tests, as well as by dynamicmechanical analysis (DMA). Lu and Anseth [138] developed water-soluble PEGbased macromers containing oligomeric blocks of a hydrolyzable ␣-hydroxy acid photopolymerized to degradable hydrogels to control the release of high-molecular weight, hydrophilic solutes. Network mesh size and volume swelling ratio during the degradation process were characterized to study the effect of hydrogel degradation kinetics

Fig. 14. Chemical reactions of the prepared hydrogel formulations.

and subsequent changes in network structure on solute release. The release behavior was examined and related to the swelling and structural properties of hydrogels. Finally, the ability to tailor the release from multilaminated biodegradable hydrogels to obtain an approximate zero-order release profile was studied theoretically. Choi et al. [139] synthesized a variation of A–B–A triblock copolymers consisting of PLLA and PEG (Fig. 15) and examined it for complexation with ␣-CDS. The copolymer, confined to the CD channels, lost its original crystalline properties but formed a channel-type hydrophobic crystalline structure with CDs due to the long-chain nature of the copolymers. Such a polymeric inclusion complex can have an important role for supramolecular architectures such as polyrotaxanes and molecular tubes for the use of bio-active agent delivery systems. Einerson et al. [140] developed gelatin-based hydrogels for drug delivery by modifying gelatin with PEG-dialdehyde and/or EDTAD (Fig. 16). Swelling/degradation studies showed that the modified hydrogel increased the time to reach the maximum swelling weight ratio Tmax and the time Tfail to failure by hydrolysis (but had little effect on the maximum swelling weight ratio Rmax and the weight ratio Rfail at failure. In vitro drug release studies showed that the hydrogel generally decreased the maximum mass ratio Dmax of drug released and time Tdmax to reach Dmax . Percent glutaraldehyde fixation did not significantly affect Dmax to Tdmax (except for EDTAD-modified gelatin hydrogels). In vivo studies showed that gelatin-based hydrogels elicited comparable levels of acute and chronic inflammatory response to that of the empty-cage control by 21 days. 3.3.2. Pendant chain systems In these systems, the drug molecule is chemically bonded to a polymer backbone and the drug is released by hydrolytic or enzymatic cleavage. The rate of drug release is controlled by the rate of hydrolysis. This approach provides an opportunity to target the drug to a particular cell type or tissue. Natural polymers, e.g. starch, chitosan, etc. as well as synthetic polymers, e.g. polylysine and co-polymers of 2-hydroxypropylamide have been used as drug carriers in such systems. The structure of these polymers can be modified by the incorporation of sugar residues or sulfonyl units to obtain a specific tissue affinity. Shah et al. [141] developed copolymers of NIPAM and Nacryloxy succinimide (AS) by radical polymerization of monomers. They measured the kinetics of hydrolysis of succinimide side chains at temperatures below, at, and above the LCST of the copolymers. The rate of hydrolysis accelerated on dissolution. As a result, the kinetics exhibited either a zero-order or an S-shaped profile. In the latter case, the initial slow phase was dependent on the difference between the experimental temperature and the LCST (Fig. 17). The matrix of poly(methyl methacrylate-co-maleic anhydride) with a surface containing anhydride group at different percentages was prepared by solution polymerization and characterized [142]. A macromolecular prodrug of ampicillin was synthesized by linking the amine group of ampicillin to anhydride group of matrix via an amide bond. The amount of ampicillin covalently bound to the

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Fig. 15. Synthetic route for the copolymers of PLLA (a) and PLLA–PEG–PLLA (b) and proposed structure of the polypseudorotaxane between PLLA–PEG–PLLA triblock copolymers, and ␣-CDS (c).

matrix was spectroscopically characterized and the in vitro release rate in a weakly basic medium was studied with its antimicrobiological activity. The prodrug allows a prolonged release (7–8 days) of the drug. Many anticancer drugs are cytotoxic and, therefore, have to be released to the affected site only, so as to protect normal cells from possible side effects. The attachment

of anti-cancer agents to polymers could be a promising approach towards reducing the overall toxicity arising from poor delivery and targeting [143]. A new tetrahydrophthalimide monomer containing 5-fluorouracil (ETPFU) and its homopolymer and copolymer with AA and with vinyl acetate (VAc) have been synthesized and spectroscopically characterized. The in vivo antitumor activities of all the

Fig. 16. Gelatin modified by PEG dialdehyde and EDTAD.

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Fig. 17. Rate of hydrolysis at different reaction temperatures.

polymers in Bolb/C mice bearing the sarcoma 180 tumor cell line were greater than those of 5-FU and monomer at the highest dose (800 mg kg−1 ). 3.4. Swelling-controlled systems Hydrogels consist of macromolecular chains crosslinked to create a tangled mesh structure, providing a matrix for the entrapment of drugs. When such hydrogels come in contact with a thermodynamically compatible solvent, of polymer chains relax [144]. This happens when the characteristic glass–rubber transition temperature of the polymer is below the temperature of experiments. Swelling is the macroscopic evidence of this transition. The dissolved drug diffuses into the external receiving medium, crossing the swollen polymeric layer formed around the hydrogel. When the hydrogel contacts the release medium, the penetrant water molecules invade the hydrogel surface and thus a moving front is observed that clearly separates the unsolvated glassy polymer region ahead of the front from the swollen and rubbery hydrogel phase behind it. Just ahead of the front, the presence of solvent plasticizes the polymer and causes it to undergo a glass-to-rubber transition [145]. Now, the following possibilities arise: (i) If the glass transition temperature Tg of polymer is well below the experimental temperature, the polymer will be in the rubbery state and polymer chains will have a high mobility that allows easier penetration of the solvent into the loaded hydrogel and subsequent release of the drug molecules into the release medium [146]. This clearly results in Fickian diffusion (Case I) which

is characterized by a solvent (or drug) diffusion rate Rdiff slower than the polymer chain relaxation rate Rrelax (Rdiff  Rrelax ). (ii) If the experimental temperature is below Tg , the polymer chains of hydrogels are not sufficiently mobile to permit immediate penetration of the solvent into the polymer core. The latter situation gives rise to a non-Fickian diffusion process which includes Case II and anomalous diffusion, respectively, depending on the relative rates of diffusion and chain relaxation (for Case II, Rdiff  Rrelax , and for anomalous diffusion, Rdiff ∼ Rrelax ). The entire mechanism is modelled in Fig. 18. Water soluble drug transport in crosslinked polymeric material was investigated [147] to determine the effect of polymer morphology, composition and solute properties on release behavior. Two crosslinked polymer systems, P(HEMA-co-MMA) and PVA were used to study water transport and solute release experiments. Drug release rates, drug diffusion coefficients and the swelling interface number were used to characterize solute transport. Release experiments were conducted using eight different solutes. Poly(diethylaminoethyl methacrylate-co-hydroxy ethyl methacrylate), poly(diethylamino ethyl acrylate-cohydroxy methacrylate) and poly(methacrylamine propyl ammonium chloride-co-hydroxyethyl methacrylate) were synthesized by free radical polymerization [148]. The anomalous transport behavior in these polymers was analyzed and the swelling front velocity was determined. The transport was found to be anomalous at acid pH values where the polymer networks were ionized. Release of

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Fig. 18. The mechanisms of Case II and anomalous diffusion.

insulin, myoglobin and albumin was studied and found to be strongly dependent on the mesh size of the polymer network. Polyionic complexes of CS and PAA were prepared [149] for a wide range of copolymer compositions and release of amoxicillin trihydrate and amoxicillin sodium from these complexes was studied. The diffusion of amoxicillin trihydrate was controlled only by the swelling/eroding ratio of the polyionic complexes. The swelling of amoxicillin sodium loaded hydrogels was greater than that of amoxicillin trihydrate formulations. It was concluded that the water uptake was mainly governed by the degree of ionization. Bajpai et al. [150] prepared a semi-IPN of PVA and poly(acrylamide-co-styrene) (PAMS) and assessed its potential for controlled release of tetracycline. They investigated the influence of various experimental parameters, such as percent loading, composition of the IPN, thickness of the loaded device, pH and nature of the release medium, on the release profile of the drug. In another study, these authors [151] also studied modulation of in vitro release of crystal violet from a binary polymer hydrogel system of different chemical compositions at varying pH. Bajpai and Bhanu [152] prepared Sem-IPNs of PEG, PVA and polyacrylamide as a support for enzyme immobiliza-

tion and performed kinetic studies by immobilization of ␣-amylase. The effect of IPN composition on the extent of immobilization was investigated and percentage of relative activity of the immobilized enzyme was evaluated as a function of chemical architecture of the IPNs, pH and temperature, taking starch as a substrate. The authors also evaluated the potential of these hydrogels for insulin release [153]. The release data were analyzed by Fick’s power law and the influence of various factors on the plausible mechanism of insulin release was examined. Two polymeric hydrogels containing PVP-crosslinked PAM and PVA-crosslinked PAM were loaded with sulphamethoxazole to study the swelling and drug release dynamics at fixed pH and room temperature [154]. The effect of various factors such as the composition of the hydrogel, crosslink density and drug loading capacity were studied on the swelling and drug-release pattern of the hydrogels. 3.5. Modulated-release systems In these systems, the drug release is controlled by external stimuli such as temperature, pH, ionic strength, an electric field, electromagnetic radiation or UV light, etc. Hydrogels that respond to these external stimuli can be used as controlled-release devices. The following discus-

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sion focuses on various drug delivery systems responsive to external signals. 3.6. Temperature-sensitive release systems Temperature is the most widely utilized triggering signal for a variety of triggered or pulsatile drug delivery systems. The use of temperature as a signal has been justified by the fact that the actual body temperature often deviates from the physiological value (37 ◦ C) in the presence of pathogens or pyrogens. This deviation can be a useful stimulus to activate release of therapeutic agents from various temperature-responsive drug delivery systems for diseases accompanied by fever. Drug-delivery systems responsive to temperature utilize various polymer properties, including the thermally reversible transition of polymer molecules, swelling change of networks, glass transition and crystalline melting [155]. Temperature-sensitive hydrogels are probably the most commonly studied class of environment-sensitive polymer systems in drug delivery research. These hydrogels are able to swell or deswell as a result of change in the temperature of the surrounding fluid. For convenience, temperature-sensitive hydrogels are classified into negatively thermosensitive, positively thermosensitive and thermally reversible gels [156]. 3.6.1. Negative thermosensitive hydrogels Negative temperature-sensitive hydrogels have a lower critical solution temperature (LCST), which may be defined as the critical temperature below which the polymer swells in the solution while above it the polymer contracts. Below the LCST, the enthalpy term, related to the hydrogen bonding between the polymer and the water molecules, is responsible for the polymer swelling. When the temperature is raised above the LCST, the entropy term (hydrophobic interactions) dominates, leading to polymer contraction. The efficiency of the hydrogen bonding process has negative temperature dependence; and above the LCST the hydrogen bonds between the monomer side groups and water molecules will be increasingly disrupted with increasing temperature [157]. The backbones of the polymer, the long chains of C–C bonds to which the side chains are attached are hydrophobic and tend to reduce their surface area exposed to the highly polar water molecules. They can do so by forming aggregates as shown in Fig. 19. Normally, when hydrogen bonds between the side groups and the water are present, the aggregation of the backbone is prevented because the hydrogen bond interactions with the water molecules are stronger than the backbone interactions. When the hydrogen bonds are broken by increasing thermal agitation, aggregation takes place, resulting in shrinkage of the thermosensitive hydrogel with increasing temperature. A well-known polymer with LCST at 32 ◦ C is poly(Nisopropyl acrylamide) (NIPAAm) which has been extensively employed as a negative thermosensitive hydrogel. These hydrogels show an on–off drug release [158] with “ON” at a low temperature and “OFF” at high temperature

Fig. 19. Backbones of the temperature-sensitive hydrogel in the (a) swollen and (b) aggregated condition.

allowing the desired drug release. Fig. 20 shows a schematic illustration of the on–off release. It is a common strategy to modulate the LCST of a polymer by incorporating hydrophilic or hydrophobic moieties in the polymer structure. For example, when NIPAAm is copolymerized with hydrophilic monomers such as acrylamide (AAm), the LCST increases up to about 45 ◦ C when 18% of AAm is incorporated in the polymer, whereas LCST decreases to about 10 ◦ C when 40% of hydrophobic N-tert-butyl acrylamide (N-tBAAm) is added to the polymer [159]. Some other polymers belonging to the NIPAAm family are poly(N,N -diethyl acrylamide) (LCST 26–35 ◦ C) [160], poly(dimethylamino ethylmethacrylate) (LCST ≈ 50 ◦ C) [161] and poly(N-CL)1-hydroxymethyl) propylmethacrylamide (LCST ≈ 30 ◦ C) [162]. LCST systems are relevant for controlled release of drugs, and of proteins in particular [163]. Thermosensitive polymers may be fixed on liposome membranes; and in that case liposomes exhibit control of their content release [164]. 3.6.2. Positive thermosensitive hydrogels A positive temperature-sensitive hydrogel has an upper critical solution temperature (UCST). The hydrogel contracts upon cooling below the UCST. Polymer networks of poly(acrylic acid) (PAA) and polyacrylamide (PAAm) or poly(acrylamide-co-butyl methacrylate) exhibit positive temperature dependence of swelling [165].

Fig. 20. Schematic illustration of on–off release from a squeezing hydrogel device for drug delivery.

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Most commonly used thermoreversible gels are prepared from poly(ethylene oxide)-b-poly(propylene oxide)b-poly(ethylene oxide) (Pluromics, Tetronics, Poloxamer). The polymer solution is a free-flowing liquid at ambient temperature but gels at body temperature. Such a system would be easy to administer into a desired body cavity. In some cases, if lowering the amount of thermogelling polymer is necessary, it may be blended with a pH-sensitive, reversibly gelling polymer. Recently, new series of biodegradable triblock copolymers were designed. The polymers, consisting of poly(ethylene glycol)–poly(d,l-lactic acid-co-glycolic acid)–poly(ethylene glycol) (PEG–PLGA–PEG) [166] or PLGA–PEG–PLGA [167], were investigated for sustained injectable drug delivery systems. Some natural polymers like xyloglucan may also form thermoreversible gels [168]. Cappello et al. [169] developed novel protein polymers “ProLastins”, which undergo an irreversible sol–gel transition. When injected as a solution into the body, the material forms a firm and stable gel within minutes. It remains at the site of injection providing absorption times from less than 1 week to many months. Temperature-sensitive hydrogels can also be placed inside a rigid capsule containing holes or apertures. The on–off release is achieved by the reversible volume change of the temperature-sensitive hydrogel [170]. Such a device is called a squeezing hydrogel device because the drug release is affected by the dimensions of the hydrogel. In this type of system, the drug release rate was found to be proportional to the rate of squeezing of the drug-loaded polymer. A US patent [171] describes a medical device containing a thermosensitive cellulose gel structure, which can deliver a bioactive-solute compound to a target location in the body. The gel structure deswells at a certain temperature and expels the biologically active solute upon an increase in gel temperature. The shrinking cellulose gel releases the solute in a sustained, convective release pulse such that substantially all of the loaded solute is released in a relatively short time under the influence of increased body temperature. The polymer network is characterized by positive molecular interactions between the different components of the system. These interactions may be physical, such as chain entanglements, or chemical, such as ionic interaction, hydrogen bonding, van der Waals attraction and covalent bonding. Another US patent [172] describes a class of semiinterpenetrating polymeric networks that includes linear polymer molecules functionalized with a bioactive moiety. The linear polymer is physically entangled with the bioactive moiety. The polymer network, which can be used as a matrix in tissue engineering, is fluid at room temperature, but becomes solid or semisolid at elevated temperature, e.g. at mammalian body temperatures. Finally, US patent [173] describes a transbody-surface drug delivery device for the administration of bioactive molecules to an individual at a therapeutically effective rate. The system comprises a thermo-responsive gel with the capability of reversibly swelling and deswelling to control the release of embedded bioactive molecules.

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3.7. pH sensitive release systems A polyelectrolyte is a macromolecule that dissociates to give polymeric ions on dissolving in water or other ionizing solvent. Because of the repulsion between charges on the polymer chain, the chain expands when it is ionized in a suitable solvent. However, if the solvent prevents ionization of the polyelectrolyte, the dissolved chain remains in a compact, folded state. If the polyelectrolyte’ chains are hydrophobic when unionized in a poor solvent, they collapse into globules and precipitate from solution. The interplay between hydrophobic surface energy and electrostatic repulsion between charges dictates the behavior of the polyelectrolyte [174]. Since the degree of ionization of a weak polyelectrolyte is controlled by pH, and the ionic composition of the aqueous medium, “smart” polymers dramatically change conformation in response to minute changes in the pH of the aqueous environment. All pH sensitive polymers contain pendant acidic or basic groups that either accept or donate protons in response to the environmental pH [175]. Swelling of a hydrogel increases as the external pH increases in the case of weakly acidic (anionic) groups, but decreases if the polymer contains weakly basic (cationic) groups. When the ionic strength of the solution is increased, the hydrogel can exchange ions with the solution. By doing so, the hydrogel maintains charge neutrality and the concentration of free counter ions inside the hydrogel increases. An osmotic pressure difference between the hydrogel and the solution arises and causes the gel to swell. When the ionic strength is increased to high levels (1–10 M), the hydrogel will shrink. This is due to the decreasing osmotic pressure difference between the gel and the solution. The solution now has an osmotic pressure in the range of the osmotic pressure inside the gel, as shown in Fig. 21. Most anionic pH-sensitive polymers are based on polyacrylic acid (PAA) or its derivatives. In addition to PAA, polymethacrylic acid (PMMA), poly(ethylene imine) and poly(l-lysine) have also been explored for use in drug delivery [176]. Microparticles of poly(methacrylic acid-g-ethylene glycol) P(MAA-g-EG) loaded with insulin exhibited unique pH-responsive characteristics whereby interpolymer complexes formed in acidic media and dissociated in neutral/basic environments [177]. Consequently, insulin release from the gel was significantly retarded in acidic media while rapid release occurred under neutral/basic conditions. Copolymer networks of poly(methacrylic acid) grafted with poly(ethylene glycol), showing reversible pH-dependent swelling behavior due to the formation of interpolymer complexes between protonated pendant acid groups and the etheric groups on the graft chains, have been developed [178]. Gels containing equimolar amounts of MAA/EG exhibited less swelling at low pH. pH-responsive polymers based on acrylic acid derivatives were introduced by Palasis [179]. The polymers were derivatized to contain moieties cationically charged at pH below their pKa value. Thus, they attract negatively charged therapeutic agents. At pH values above their pKa , the polymers become predominantly uncharged and substantially release the therapeutic agents. Bae and Park [180]

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Fig. 21. The swelling of a pH-sensitive hydrogel as a function of the ionic strength.

described pH-sensitive polymers containing suflonamide groups, which show changes in swellability and solubility depending on pH. The pH-sensitive polymer may be linear, grafted, a copolymer or a hydrogel. pH-dependent hydrogels may be grouped into two main classes: (a) cationic hydrogels and (b) anionic hydrogels. 3.7.1. Cationic hydrogels Cationic hydrogels swell and release a drug in the low-pH environment of the stomach. Risbud et al. [181] developed a pH-sensitive CS/PVP based controlled drug release system using air-dried and freeze-dried amoxicillin. Porous freeze-dried hydrogel exhibited superior pH-dependent swelling properties over non-porous airdried hydrogels. Freeze dried membranes released around 73% of the amoxicillin (33% by air dried) in 3 h at pH 1.0 and thus, had the better drug release properties. Spherical crosslinked beads using chitosan, glycine and glutaraldehyde were prepared by Gupta and Ravi kumar [182]. The swelling behavior of the beads was monitored as a function of time in solutions of different pH. The release experiments were performed using thiamine hydrochloride as the model drug. The chitosan beads showed a pH-dependent swelling behavior which makes them appropriate for delivery of drugs in an acidic environment. 3.7.2. Anionic hydrogels Hydrogels of PAA or PMA can be used to develop formulations that release drugs in a neutral pH environment [183,184]. Hydrogels of polyanions (e.g. PAA) crosslinked with azoaromatic crosslinkers were developed for colonspecific drug delivery. Swelling of such hydrogels in the stomach is minimal and thus the drug release is also minimal. The swelling increases as the hydrogel passes down the intestinal tract due to an increase in pH leading to ionization of carboxylic groups. But only in the colon can the azoaromatic crosslinks of the hydrogels be degraded by azoreductase produced by the microbial flora of the colon

[185,186] as shown in Fig. 22. The degradation kinetics and degradation pattern can be controlled by the crosslinking density. The kinetics of swelling of hydrogels can be controlled by changing the polymer composition [187], which can be changed as the pH of the environment changes. Some pendant groups, such as N-alkanoyl (e.g. propionyl, hexanoyl and lauroyl) and o-acylhydroxylamine moieties, can be hydrolyzed as the pH changes from acidic to neutral; and the rate of side-chain hydrolysis is dependent on the length of the alkyl moiety. In a novel study by Kim and Peppas [188], the mesh sizes of the hydrogels were very small (18–35 Å) in the collapsed states at pH 2.2 and very large (70–111 Å) in the swollen states at pH 7.0. In addition, as the content of MAA in the feed monomers was increased, the mesh size decreased at pH 2.2 but increased at pH 7.0. When the crosslink-

Fig. 22. Schematic illustration of oral colon-specific drug delivery using biodegradable and pH sensitive hydrogels.

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Fig. 23. Schematic illustration of copolymerization of star polymer with the ionizable MAA.

ing ratio of the copolymer increased, the swelling ratio decreased at both pH 2.2 and pH 7.0. These hydrogels are useful for the development of oral protein delivery carriers. In another study [189] it was found that an increase in the degree of ionization contributed to the electrostatic repulsion between adjacent ionized groups, leading to chain expansion, which, in turn affected macromolecular chain relaxation. However, for both P(MAA-co-MAA) and P(MAAg-Eg) hydrogels, the swelling mechanism exhibited little dependence on the copolymer composition of each hydrogel at the same pH. The conventional calcium-alginate and newly formulated calcium-pectinate or calcium-alginate-pectinate systems have been evaluated for their release properties in simulated gastric and intestinal environments [190]. It was shown that the drug entrapment capacity of a weak acidic drug, such as diclofenac sodium, was significantly enhanced by a pH drop of the crosslinking solution. Dissolution studies over a pH range similar to that of the human gastrointestinal tract demonstrated that the rate of drug release from different pellets in simulated small-intestine environments ranged from rapid to slow (i.e. 100% drug release ranged from 4 h to 10 h) but always in a controlled manner, depending on the type of formulation employed. Temperature- and pH-sensitive IPN hydrogels of PMMA and PNIPAAM were prepared by a sequential UV polymerization [191]. These hydrogels exhibited a swelling transition at 31–32 ◦ C, the LCST of the PNIPAAM network, and a pH of approximately 5.5, the pKa of PMAA. Both pH and temperature have a strong influence on the permeability. A hypothetical mechanism has been proposed to explain the phenomenon of a model drug having the highest permeability at the physiological state of 37 ◦ C and pH 7.4—which is desirable in membrane and drug delivery system. Responsive and recognitive polymeric networks based on PEG star polymers were prepared to respond to pH

changes in their environment and to distinguish between d-glucose and similar sugars [192]. The large number of functional groups provided by the star polymer was essential for interactions between templates and the network during polymerization; network sensitivity to changes in pH was obtained by copolymerization of star polymer with the ionizable MAA (Fig. 23). Such systems have potential for oral drug delivery and actuators. Hydrogels based on n-alkyl methacrylate esters (n-AMA), AA and AAm crosslinked with 4,4 di(methacryloylamino) azobenzene have also been prepared [193]. The factors that exert greatest influence on the swelling behavior of the gels include the degree of crosslinking, the lengths of the n-AMA side chains and pH. By adjusting these factors, the degree of swelling of the hydrogel in the small intestine can be controlled and consequently the drug may be prevented from being released before reaching the colon. Anionic interpenetrating HA–PHEA gels were prepared by copolymerization of GMA–HA and HEA [194]. Release of chloropromazine HCl as a model cationic drug from HA–PHEA gels was suppressed significantly in water. The release was found to increase with increasing ionic strength and decreasing pH of solutions. pH-responsive insulin hydrogels were obtained by copolymerization of MA-insulin with AA in aqueous solution using APS/TMEDA as an initiator system [195]. At low AA content, the swelling of hydrogel at pH 7.4 was predetermined by increased crosslinking density, the swelling increased with increasing AA owing to an enhanced ionic osmotic contribution and low polymerization efficiency. The results show that the change in swelling in response to pH changes increases with increasing AA content. When hydrogels of polyacrylic acid are crosslinked with MDIC and ODIC, the amounts of drug released from hydrogels at pH 5 and pH 7.4 are different [196]. This is explained by the fact that a hydrophobic complex between PAA and

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Fig. 24. A potential system for drug-delivery of peptide and protein drugs to different regions of intestinal tract.

PEG units is formed at pH 5. The existence of compact hydrophobic domains in the polymeric network hampers the drug release. However, at pH 7.4 a large number of carboxylic groups of PAA are ionized. This leads to destruction of the complex and swelling of the network, resulting in full release of the drug. At pH 2, the complex is as stable as at pH 5 but the amount of drug released is higher. The most probable reason for this is that in artificial stomach juice, with a pH 2, the hydrolysis of the PAA network takes place at a greater rate and to a larger degree. pH/temperature-sensitive polymers P(NIPAAm-coBMA-co-AA) were used to prepare insulin releasing beads via loading in aqueous solution [197]. At acidic pH and body temperature, the beads were insoluble, and thus no drug was released in the stomach. At pH 7.4 and body temperature, the low-molecular weight hydrophilic polymeric beads displayed a hump-like profile and dissolved within 2 h (bead dissolution controlled release mechanism), while the high-molecular weight hydrophilic polymeric beads swelled and released insulin slowly over a period of 8 h. Thus, the unique properties of the pH/temperaturesensitive polymeric beads make this a potential system for drug delivery of peptide and protein drugs to various regions of the intestinal tract (Fig. 24). 3.8. Magnetic-sensitive release systems Magnetism has a profound influence on living organisms. The hemoglobin in our blood, an iron-containing

protein, is magnetic. Magnetotactic bacteria are perhaps the first living organisms to orient themselves with the earth’s magnetic field [198]. These bacteria are known to contain aligned chains of magnetic particles of various shapes. There is now substantial evidence that all living organisms, including animals and humans, contain magnetic particles that act as magnetic receptors [199]. It is an established fact that the magnetism and magnetic materials have a strong role to play in health care and biological applications [200–203]. The combination of fine particles and magnetism in the field of biology and biomaterial has been found useful in sophisticated biomedical applications such as cell separation [204–206] gene and drug delivery, and magnetic intracellular hyperthermia treatment of cancer [207,208]. Fig. 25(a) demonstrates a procedure for preparing iron oxide (␥-Fe2 O3 ) nanoparticles encapsulated with arginine (Arg) peptide (RRRRR CK-FITC) conjugated with poly(d,llactide-co-glycolide) (PLGA). It should be noted that the Arg peptide is well known as a cell penetration peptide leading to intracellular transport across the cellular membrane into the cytoplasm and nucleus by a seemingly energyindependent mechanism. Fig. 25(b) shows the sequence of Arg peptide. The fluorescein conjugated Arg-peptide was prepared by Peptron (Peptron Inc., South Korea, established in 1997 and focused on custom peptide synthesis, www.peptron.com). The FITC was conjugated to observe the intracellular translocation of magnetic nanoparticles into the cell [209].

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Fig. 25. (a) The schematic diagram for preparing the ␥-Fe2 O3 -PLGA-Arg-FITC nanoparticles. (b) The sequence of arginine-peptide.

Encapsulation of magnetic particles with preformed natural or synthetic polymers is the simple and classical method to prepare magnetic polymeric particles. The development of emulsion polymerizations, such as the conventional emulsion polymerization process, miniemulsion and microemulsion polymerization, and soap-free emulsion polymerization, have led to new synthesis methods [210–213]. The typical method based on emulsion polymerization is to suspend magnetic particles in the dispersed phase and then polymerize the monomer in the presence of the magnetic particles to form magnetic polymer particles. The miniemulsion polymerization is very suitable for making magnetic polymeric particles. In miniemulsion polymerization, the monomer droplets with magnetic nanoparticles act as “nanoreactors”. Xu et al. [214] used inverse miniemulsion polymerization to synthesize magnetic polyacrylamide particles containing nanosize magnetic iron oxide. The particle size of magnetic polymeric particles is about 100 nm and the particles are superparamagnetic. A US patent [215] describes magnetosomes comprising magnetic monocrystals having a maximum diameter of 45 nm surrounded by a membrane and their use in triggered release in the body. These magnetosomes with cationic charges increase the probability that antibodies and therapeutic agents can be correctly bound to them. Another US patent [216] discusses nanosized particles termed as “nanoclinics” or “nanoparticles” or “nanobubbles”. These nanoparticles are prepared from iron oxide particles by a multistep synthetic process. A tracking agent

such as the “two-photon fluorophore” C625 is attached to the surface of the particles to track the nanoparticles using two-photon laser scanning microscopy, a powerful technique for probing the three-dimensional structure of a cell providing inherent optical sectioning capability without any significant interference from autofluorescence. Sodium silicate is added to preform the silica shell. A carbon spacer (e.g. C625) is then attached to the silica surface. The spacer reduces the steric hindrance for the binding of the targeting agent to its target molecules. Another US patent [217] focuses on a treatment method that involves the administration of a magnetic material composition containing single-domain magnetic particles. The application of an alternating magnetic field to inductively heat the magnetic material composition causes the triggered release of therapeutic agents at target tumor or cancerous cells. Pankhurst and co-workers [218] studied the physical principles underlying some current biomedical applications of magnetic nanoparticles, i.e. the relevant physics of magnetic materials and their responses to applied magnetic fields. They studied the way these properties are controlled with reference to (a) magnetic separation of labeled cells and other biological entities, (b) therapeutic drug, gene and radionuclide delivery, (c) radio frequency methods for catabolism of tumors via hyperthermia, and (d) enhancement agents for magnetic resonance imaging applications. Tertaj et al. [219] presented various synthetic routes for the preparation of magnetic nanoparticles useful for biomedical applications.

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Robin [220] studied cancer treatment by electromagnetically activated “nanoheaters.” Heating cancer cells in a clinical setup can be achieved by various technologies such as focused ultrasound, radio frequency, thermal radiation, lasers and magnetic nanoparticles. The size of the magnetic nanoparticles lies in the range 5–90 nm, which is about one-thousandth the size of the smallest biological cell but also about a thousand times larger than typical molecules. To turn these particles into heaters, they are subjected to an oscillating electromagnetic field, where the field direction changes periodically. Application of the magnetic field generates a directional force on each magnetic particle. As the magnetic field oscillates at high frequency, the average force on the particle is zero. The energy of the oscillation is converted into heat, raising the temperature of the nanoparticles and their biological material. 3.9. Electric-sensitive release systems Electrically responsive delivery systems are prepared from polyelectrolytes (polymers that contain a relatively high concentration of ionizable groups along the backbone chain) and are thus pH-responsive as well. Under the influence of an electric field, electroresponsive hydrogels generally deswell or bend, depending on the shape and orientation of the gel. The gel bends when it is parallel to the electrodes, whereas deswelling occurs when the hydrogel lies perpendicular to the electrodes. Ramanathan and Block [221] examined and characterized the use of chitosan gels as matrices for electrically modulated drug delivery. In electrification studies, release time profiles for neutral (hydrocortisone), anionic (benzoic acid) and cationic (lidocaine hydrochloride) drug molecules from hydrated chitosan gels were monitored in response to milliampere currents as a function of time. Likewise, chondroitin-4sulfate hydrogels were examined by Lensen et al. [222] as potential matrices for electrocontrolled delivery of peptides and proteins. Kim et al. [223] synthesized an interpenetrating polymer network (IPN) hydrogel composed of poly(vinyl alcohol) and chitosan which exhibited electrosensitive behavior. They investigated the response of the IPN hydrogel in electric fields. A swollen PVA chitosan IPN was placed between a pair of electrodes and bending behavior in response to the applied electric field was noted. The bending angle and the bending speed of the PVA/chitosan IPN increased with increasing applied voltage and concentration of NaCl aqueous solution. Synthetic as well as naturally occurring polymers either alone or in combination, have also been used. Examples of naturally occurring polymers include hyaluronic acid, chondroitin sulfate, agarose, xanthan gum and calcium alginate. The synthetic polymers are mostly methacrylate and acrylate derivatives such as partially hydrolyzed polyacrylamide, polydimethylaminopropyl acrylamide, etc. Electrically induced anisotropic gel deswelling was first explained by Tanaka et al. [224], who suggested that the electric field produces a force on both the mobile counter ions and the immobile charged groups of the gel’s polymeric network. For example, in partially hydrolyzed polyacrylamide gels fixed to one of the electrodes, the

Fig. 26. Representation stress at cathode and anode.

mobile H+ ions migrate toward the cathode while the negatively charged immobile acrylate groups in the polymer networks are attracted toward the anode. The anionic polymeric gel network is pulled toward the anode. The pull creates a uniaxial stress along the gel axis, the stress being greatest at the anode and smallest at the cathode (Fig. 26). This stress gradient is thought to contribute to the anisotropic gel deformation. When the gel is not fixed to either electrode, the attractive forces between the immobile negative charges of the polymer network and the anode can result in translational movement of the gels towards the anode. In a recent study by Bajpai and coworkers [225], polyaniline (PANI) was impregnated into a macromolecular matrix of poly(vinyl alcohol)-g-poly(acrylic acid) and electrical conductivity and electroactive behavior of the resulting nanocomposite was studied. 4. Challenges and prospects Synthetic strategies to fabricate different types of drug delivery systems, as discussed in the preceding sections, present a scenario of drug delivery science. The situation, however, is not simple. There have to be considered a number of factors that create hurdles for ultimate approval of the drug-loaded system in polymer medicine. These factors may be, for example, biocompatibility of the device, cytotoxicity, in vivo studies, FDA approval, efficiency, inconvenience caused to patients, cost effectiveness, etc. Only after these factors are examined, can a drug delivery system be therapeutically acceptable. The delivery of a drug at a predetermined rate over a specified time to a selected target organ has been the ideal requisite in drug delivery technology and pharmacokinetics. Moreover, the need for carriers that exhibit oscillatory behavior of the releasing bioactive agent has also emerged as a significant problem of drug design and formulation in recent years. The traditional methods of drug administration in conventional forms, such as pills and subcutaneous or intravenous injection, are still the predominant routes for drug administration. But pills and injections offer limited control over the rate of drug release into the body; usually they are associated with an immediate release of the drug. Consequently, to achieve therapeutic levels that

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extend over time, the initial concentration of the drug in the body must be high, causing peaks that gradually diminish over time to an ineffective level. In this mode of delivery, the duration of the therapeutic effect depends on the frequency of dose administration and the half-life of the drug. This peak-and-valley delivery is known to cause toxicity in certain cases, most frequently with chemotherapy for cancer. Thus, the design of a drug delivery system with optimum performance in specific circumstances poses challenges. In an overview of the whole scenario, the field of drug delivery systems has to confront the following challenges: • • • • • •

Improved efficacy. Targeted delivery and reduced side effects. Optimum performance. Interfacing and pacing with modern methodologies. Guarantees of safe environment. Ease of fabrication and application in reality.

These benefits may be realized by adopting approaches that basically involve judicious combination of highly specific monomers and polymers of both synthetic and natural origin. The use of smart materials in drug delivery technologies has not only to focus on the possible medical benefits but also must consider economic aspects of the developed materials and/or technology. Furthermore, huge effort on synthetic polymer chemistry must be undertaken to design tailor-made macromolecular systems that will offer novelty in their operation and performance. Above all, the systems developed must be acceptable to the patient community who are the end-users of any successful research and technology. A logical consideration of the possibilities about bright prospects for controlled drug delivery gives rise to positive signals and, therefore, more effort deserves to be put into its growth and expansion. Since smart materials have specific mode of operability and are prone to typical experimental conditions, there is large scope for synthetic polymer chemistry to design multiresponsive delivery systems. Despite the tremendous research input that has been applied to achieve high performance technologies, a number of aspects still remain to be worked on: • Designing of drug-delivery systems with multistimuliresponsive potential. • More precise synthetic routes for making responsive materials with greater responsive sensitivity. • Assurance of economic viability so as to popularize devices on a large commercial and population scale. • Design of more localized drug delivery systems. • Oral delivery of insulin using body-friendly natural polymers with enhanced absorption in blood. Thus, it may be concluded that although much advancement has been demonstrated by untiring efforts of researchers worldwide, still there exist numerous challenges that have to be addressed. Moreover, the field of controlled drug delivery offers a wide scope and future prospects to build-up a technology with high performance, economically viable and potentially efficient.

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5. Conclusions Polymer hydrogels exhibiting a desired response to external signals such as pH, temperature, electric field, magnetic field, etc. form the foundation of a novel class of materials that find unique applications in biomedical and pharmaceutical sciences. Among various types of polymer hydrogels, polymer blends of natural and/or synthetic polymers possess many advantages such as easy fabrication of devices, manipulation of device properties, etc. Another class of responsive polymers of vital utility is interpolymer complexes formed as a result of non-covalent association of polymers. Some very effective pH-sensitive interpolymer complexes are PMAA/PEG and proton acceptor polymers such as poly(vinyl esters), polyacrylamide, etc. Interpenetrating polymer networks (IPNs) have also been recognized as popular responsive polymers having enhanced physical properties as compared with conventional blends of their components. Block copolymers have recently been studied extensively. In particular, amphiphilic block copolymers, Zwitterionic polymers, di- and tri-block copolymers, biodegradable aliphatic polyesters, polysilane-b-PMAA, etc. have been evaluated for biomedical application. Among the various synthetic routes adopted for designing specific block copolymers are GTP, ATRP, RAFT, etc. Depending on the end-application and desired functioning, various geometrical and chemical architectures like polymer brushes and nanoparticles have been designed to achieve high performance. Precisely speaking, polymer brushes refer to an assembly of polymer chains that are tethered by one end to a surface or interface. These novel polymeric materials find application in stimuli-responsive surfaces, chemical gates, cell-growth confinement, etc. Controlled drug delivery systems offering unique advantages over conventional therapy fall into various categories such as diffusion-controlled, chemically controlled, solvent activated and modulated release systems. In diffusion-controlled systems, drugs diffuse through polymer either via leakage or degradation of the material. These systems are further classified into reservoir and matrix systems. In the former, drug diffusion through the membrane is the rate limiting step and controls the overall drug release rate, whereas in the latter case, the system exhibits first-order release behavior with decreasing release rate. In chemically controlled systems, the release of bioactive agent occurs by (i) biodegradation of drug containing polymer system, (ii) hydrolytic or enzymatic cleavage of unstable bend, and (iii) diffusion of a drug. Other widely used drug delivery systems are swelling-controlled systems. The extent of swelling determines the amount of released drug. The mechanism of drug transport depends upon the chemical architecture of the device which controls the relative rates of drug diffusion and polymer chain relaxation, the ultimate determinant of the release mechanism. Drug release systems governed by external stimuli are termed as ‘modulated release systems’. These systems demonstrate variable release rates with changing

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experimental conditions: temperature, pH, ionic strength, electromagnetic radiation, etc. Temperature-sensitive gels fall into negative and positive thermosensitive systems which exhibit a lower critical solution temperature and an upper critical solution temperature, respectively. Whereas the former gels contract above the LCST, the latter do the same below the UCST. These gels may be placed inside a rigid capsule containing holes and used as on–off release system. pH-sensitive drug delivery systems are basically polyelectrolytes of either charge and operate by widening of the mesh sizes of their network resulting from repulsive forces developed due to ionization or protonation of the constituent polymer chains. Whereas cationic polyelectrolytes like chitosan work well in the low pH environment of the stomach, anionic hydrogels such as polyacrylic acid, polymethacrylic acid, etc. work efficiently in the alkaline environment of the colon. Polymers responsive to external applied magnetic and electric fields have emerged as a novel strategy to diagnose tumor related physiological disorders. Encapsulation of nanosized iron oxide particles into a natural or synthetic polymer has been recognized as a common synthetic route to design iron oxide containing polymer nanocomposite type materials. The most commonly employed polymers are starch, polylactides, liposomes, etc. which contain iron oxide particles up to dimensions of 45 nm—often called ‘magnetosomes’. These magnetosomes have been used for triggered release in the body. The literature describes various US patents for magnetosomes as “nanoclinics” or “nanobubbles” as tracking agents in laser scanning microscopy. Likewise, electric responsive release systems show modulating release kinetics with a changing external field. Such electroactive polymers normally contain ionizable carboxylic groups along the macromolecular chains. The most commonly used synthetic polymers in this category are polyacrylic acid and polymethacrylic acid while natural polymers may be hyaluronic acid, agarose, chitosan, xanthan gum, calcium alginate, etc. Thus, polymer hydrogels of different chemical architecture with novel physico-chemical properties have shown potential to find applications as promising drug carrying vehicles in various drug delivery technologies. Although their synthesis and in vitro study seems to be simple, from the view point of in vivo applications the polymer systems need to be judged with extreme care before they can be accepted ultimately for commercial applications. References

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