Drug delivery, as the name suggests, is the method or process of administering a pharmaceutical compound (drug) to achieve a therapeutic effect in humans or animals. Key factors are to deliver the drug to the right area, at the right time and at the right concentration. However there are many obstacles to achieving successful drug delivery. To name a few, these obstacles include the drugs': low solubility, environmental or enzymatic degradation, fast clearance rates from the body, non-specific toxicity and inability to cross the biological barriers [32,33]. In order to overcome these obstacles, drug delivery carriers are being used, most of which are based on polymers [32-34].

Nevertheless, even when using a carrier to deliver the drug it is difficult to have the desirable controlled release rate. Specifically, the delivery of a drug is often problematic because the concentration of the released drug at the target area is either too low or too high and is not delivered for the desirable amount of time, as shown with the gray dotted lines in the graph (Figure 2). Thus, the desirable-controlled drug release (shown with the blue line in the graph), that is often characteristic of a zero-order kinetic rate, is not easily achievable. Consequently “smart” polymeric carriers are used to deliver drugs. These carriers allow delivery of the drug at the right time and concentration by only releasing the drug in response to an external stimulus. For example the polymer chains of a carrier may expand as a result of the temperature increasing, thus enabling the drug to diffuse out and be released from the carrier [35]. Several studies aimed at controlled drug delivery using thermoresponsive polymer based carriers are discussed later in this review, categorized by their architecture/structural characteristics.

Gene therapy aims at the treatment of many genetic diseases as it is a technique for correcting defective genes that are responsible for these genetic diseases. Specifically, the delivery of the appropriate, therapeutic gene (DNA) into the cells that will replace, repair or regulate the defective gene that causes the disease is a vital step for gene therapy. DNA, however, is a negatively charged, hydrophilic molecule; thus its delivery into the nucleus of the cell which requires it to pass through the also negatively charged and hydrophobic cell membrane is not feasible. Consequently, gene delivery carriers (also called vectors or vehicles) have been developed [36,37]. Nature's way to carry genes is viruses and these were the first carriers used for gene delivery [36,37]. However viruses have many disadvantages, the most severe of which is the immune response that they can cause and this is why non-viral carriers have been developed [36-41]. Many of these are polymer-based because polymers are cheaper and safer than viruses and also easier to tailor compared to other gene delivery carriers like liposomes [36-38,42].

When using a polymeric carrier, the main steps of gene delivery (Figure 3), also called transfection, involve the: (1) DNA and polymer complexation; (2) addition of DNA/polymer complex (sometimes also called polyplex) onto cells for a period of time commonly called the transfection time; (3) removal of complex from the cells; (4) incubation time that is when the cells are left to incubate for a time period until results are observed. Complexation is usually performed at room temperature whilst the incubation and transfection periods are at 37 °C (body temperature that the cells need to be in order to survive). Interestingly, thermoresponsive polymers have been used to enhance the transfection efficiency by changing the temperature either during the complexation and/or during the incubation or transfection period. These studies are discussed below.

In gene delivery studies where thermoresponsive polymers were used the temperature at which one or two of the aforementioned steps were performed at was changed. In particular, in studies where polyethyleneimmine (PEI) with grafted PNIPAAm [43], chitosan grafted with PNIPAAm [44], linear and branched NIPAAm, DMAEMA and PEI polymers [2] and PEG polymers with grafted PEI chains [45] were used, the complexation and transfection temperature were changed to enhance the transfection efficiency. In other studies only the incubation or complexation temperature were varied using random terpolymers of P(NIPAAm-co-DMAEMA-co-BuMA) [18,46] or PNIPAAm copolymers [47], while both complexation and incubation temperature were varied using a polyarginine polymer conjugated with PNIPAAm [48].

In an interesting study by Zhou et al. using [poly(N,N-dimethylamino propyl acrylamide)-b-PNIPAAm]-star polymers, they used the polymers thermoresponsive ability in a different manner from the above studies [49]. The complexation was undertaken at room temperature which was below the polymer LCST and then the complex was deposited on a surface above the LCST [49]. On that surface, the cells were allowed to incubate at 37 °C [49]. The result of this was increased transfection to cells cultured on the surface compared to cells grown on a surface where DNA with no polymer was deposited [49].

Another group of crosslinked hydrogels are the interpenetrating networks (IPN). Interpenetrating networks, according to IUPAC, consist of two covalently linked polymer networks which are bound together by physical entanglement as opposed to covalent bonds. Specifically, this requires the polymerization of both networks simultaneously and results in two intermixed networks that can only be separated by breaking bonds. These materials are of interest due to their ability to introduce new properties when the networks interact [95,96] or two different properties when acting independently.

In contrast to the studies discussed above, where the drug release was slowed by the temperature increase, Shin et al. demonstrated a hydrogel with a positive thermoresponse, i.e., drug release increases with increasing temperature [95]. Specifically, an interpenetrating network of polyacrylic acid (PAA) and polyacrylamide (PAAm) forms hydrogels that swell above their upper critical solution temperature, UCST, due to hydrogen bonding between the two different networks being disrupted at higher temperatures allowing the networks to swell [95]. This preliminary work showed the possibility of increasing drug release with increased temperature [95]. Recent work on the same IPN with grafted β-cyclodextrin showed a faster thermoresponse and lower UCST (35 °C) and a lowered effect of salt on the swelling [96]. Incorporation of a model drug, ibuprofen, showed a positive drug release with a controlled rate above and below the UCST [96].

Semi interpenetrating networks (SIPN) differ from IPNs in the sense that they are not entwined networks. SIPNs consist of a crosslinked network with linear or branched polymer chains penetrating them. An example of this is the SIPN synthesized by Mundargi et al. formed from crosslinked gellan gum microspheres with interpenetrating PNIPAAm chains [95]. Drug loading into the microspheres allows for pulsatile drug release. Specifically release below the LCST is greater due to swelling of the spheres and greater water transport through the spheres [97].

Kim et al. produced photodegradable injectable seminterpenetrating networks of PNIPAAm-co-PAAc with linear PAAc-co-peptide chains and showed the degradation to be dependent on peptide crosslinks and PAAc-co-peptide concentrations [98]. They also showed the degradation products to be relatively non cytotoxic demonstrating a promising tissue engineering scaffold [98].

Combining hydrophilic and hydrophobic monomers into block copolymers allows the formation of ordered structures in solution, the most common of these being the micelle. Micelles are useful for encapsulating hydrophobic drugs and delivering them into an aqueous environment.

In 1998 Chung et al. showed the effect of attaching hydrophilic or hydrophobic moieties to the terminal end of a PNIPAAm [99]. The LCST and the rate of the phase transition dramatically change depending on the nature of the end moieties [99]. More specifically, hydrophilic monomers stabilized the chain and raised the LCST and the inverse was seen when hydrophobising the chain [99]. This effect however was less pronounced when the hydrophobic modification was undertaken using a long hydrophobic chain [99]. This was found to be due to the internalization of the hydrophobic ends into micelles in solution allowing the structures to exhibit the phase transition of PNIPAAm alone [99].

Several groups have focused on using PNIPAAm as the thermoresponsive block in the formation of thermoresponsive micelles [100-104]. Akimoto et al. produced micelles of P(NIPAAm-co-DMAAm)-b-PLA, where PLA was poly(lactic acid), and showed that these micelles were able to internalize into cells above their LCST, specifically due to the increased interaction between the hydrated NIPAAm outer sphere and the cells [105]. Interestingly, they also showed that fictionalization of the PNIPAAm chain ends with thiol groups can increase this interaction and internalization, the results show promising methods for intracellular delivery of therapeutic molecules [100].

Diblock copolymers of PNIPAAm and poly(methyl methacrylate) (PMMA) or 10-undecenoic acid were synthesized by Wei et al. [103,104]. The copolymers were found to form micelles in water capable of releasing the encapsulated drug in relation to temperature [103,104]. At temperatures below the LCST, drug release stops after the initial burst release [103,104]. This is due to the hydrated PNIPAAm chains limiting drug release [103,104]. Heating above the LCST causes a big structural change in the micelle with the PNIPAAm collapsing and becoming hydrophobic, allowing release of the contained drug with an on/off drug switching [103,104].

Degradable copolymers of poly(NIPAAm-co-HPMAm-lactate)-b-PEG [106] and PEG-b-(HEMAm-lactate) [107] were shown by Hennink et al. to form micelles above a critical micelle temperature dependant on the polymer LCST [106,107]. Cleavage of the lactate moiety in vitro or in vivo causes an increase in the polymer LCST and so a destabilization of the micelles, releasing any encapsulated drug [106,107].

Preparation of thermoresponsive micelles is not limited to PNIPAAm. Figure 6 above shows a block copolymer synthesized by Huang et al., utilizing an orthoester containing acrylamide block which becomes hydrophobic above its LCST, and a hydrophilic PEG block [108]. This copolymer formed micelles above its LCST but also showed pH responsive degradation giving another option for drug release [108]. Polymers with similar properties were produced by Hruby et al. based on poly oxazolines [109] and Wang et al. based on PEO [110].

The thermoresponsive properties of terpolymers of PEGMA-b-(DMAEMA-co-MMA) were shown to have an interesting pH and temperature sensitivity [111]. Polymers transitioned from molecularly dissolved at low temperature and low pH, to micelles with either a PEGMA core upon increasing temperature or a DMAEMA core upon increasing pH [111].

San Miguel et al. synthesized ABA triblock copolymers of PDMAEMA (block A) and PVCL (block B) and demonstrated the formation of thermoresponsive micelles with sustained drug release in vitro [17].

Wei et al. synthesized a thermoresponsive star block copolymer based on L-Lactide and NIPAAm [112]. These star polymers were found to self assemble into large micelle structures in water which showed a fast on/off drug switching with temperature [112].

The formation of micelles in solution is a reversible and dynamic process; this can be overcome by crosslinking polymer chains once micellization has occurred. Selective crosslinking of specific blocks within the micelle can lead to interesting structures with varied properties. A review of non thermoresponsive shell crosslinked micelles was published by Armes et al., it focused on the crosslinking of the middle block of ABC triblock copolymers [113]. This modification allowed for synthesis without the need for low concentrations as no intermicellar crosslinking took place [113].

Wei et al. produced a number of shell crosslinked hybrid micelles using inorganic based crosslinks and PNIPAAm thermoresponsive groups [114,115]. They found drug release to be sustained for around 100 hours followed by a slow leveling off of the release rate [114,115]. The shell crosslinked layer slowed the release of drug by entrapment but allowed an increase in the release with increased temperature [114,115].

Cohn et al. utilized the formation of micelles above the LCST of PEG-b-PPG-b-PEG triblock copolymers, where PPG was the hydrophobic poly(propylene glycol) block, also called poly(propylene oxide) (PPO), to form functional nanoshells, Figure 7 [116], by incorporating a methacrylate group onto each end of the block copolymers and crosslinking them while in a micellar configuration a collapsible and expandable shell structure was formed with an extremely large void spaces capable of encapsulating numerous moieties [116].

Polymersomes, another class of self assembled structures, have also been used for drug and DNA delivery. They are based on self assembled amphiphilic block copolymers with a hydrophilic interior and exterior (Figure 8). They provide an advantage of the ability to encapsulate hydrophilic molecules in their interior protected from the outside environment; compared to micelles which preferentially encapsulate hydrophobic molecules. Reviews on this subject that include thermoresponsive polymersomes have been published previously by Meng et al. [117], Du and O'Reilly [118] and Kim et al. [119].

A diblock copolymer, synthesized by Qin et al., comprising of PEG-b-PNIPAAm with 7–20% PEG content spontaneously formed polymersomes in water at temperatures of 37 °C and above. These polymersomes would allow for encapsulation of both hydrophilic molecules within the interior, and hydrophobic molecules in the membrane promoting several possible therapeutic applications [120].

Polymersomes based on copolymers of PNIPAAm and poly(N-(3-aminopropyl)methacrylamide hydrochloride) (PAMPA) [121], PNIPAAm and PEG [122], P(LA-star-b-NIPAAm) [123] and hyperbranched poly[3-ethyl-3-(hydroxymethyl)oxetane]-g-PEG [124] where also studied, where NIPAAm or PEG where the thermoresponsive blocks.

McCormick et al. and Qin et al. utilized the LCST of a thermoresponsive block to direct the formation of polymersomes [121,122]. McCormick used the LCST of PNIPAAm to induce formation of polymersomes and crosslinked the shell by ionic crosslinking using the polyelectrolyte PAMPA [121]. Non-crosslinked polymersomes by Qin et al. were synthesized using PEG and PNIPAAm [122]. Above the LCST of PNIPAAM the polymersomes formed spontaneously, encapsulating drugs loaded into the solution beforehand [122]. These polymersomes showed no drug release until further cooling to destabilize them [122].

Expanding on the research by Qin et al., several other studies have utilized the thermoresponsive ability of polymers to aid the release from polymersomes. Star polymers synthesized by Wei et al. of a hydrophobic PLA block and thermoresponsive PNIPAAm blocks spontaneously formed polymersomes in solution [123]. Heating above the LCST hydrophobized the PNIPAAm blocks releasing any encapsulated drugs [123]. Star copolymers having a hydrophobic hyperbranched poly[3-ethyl-3-(hydroxymethyl)oxetane] (HBPO) core and many hydrophilic PEG arms were synthesized by Zhou et al. [124]. These polymers form polymersomes below their LCST, however, above the LCST the polymersomes destabilize undergoing a “vesicle-to-membrane” transformation and releasing any encapsulate [124].

In a recent interesting study, Agut et al. investigated both the pH and temperature affect on PDMAEMA containing polymersomes [125]. At low pH, the DMAEMA units were protonated and polymersomes formed due to electrostatic interactions [125]. Upon increasing the pH the mechanism changed with polymersomes only forming with increasing temperatures above the LCST of the DMAEMA [125].

Copolymer films of PNIPAAm and poly(N-butylacrylamide) were shown by Wilson et al. and Doorty et al. to give a sustained release of drugs from the film over a considerable time period [3,126]. They showed the released amounts of drug loaded at room temperature to be inversely proportional to the hydrophobic monomer content once heated to 37 °C [3,126].

Dinarvand et al. investigated the possibility of using a copolymer of PNIPAAm with PAAm as a stimuli responsive membrane for the control of permeation of molecules for numerous applications like drug delivery [127]. They found that an increasing temperature above the LCST reduces the transport through the membrane by collapse of the PNIPAAm structure [127].

An interesting use of thermoresponsive polymer films was shown by Zakharchenko et al. where a bilayer of PVCL on top of PNIPAAm with encapsulated magnetic nanoparticles was prepared [128]. At temperatures greater than the LCST the films were flat and allowed for adsorption of nanoparticles, cells or drugs onto the surface, Figure 9, upon cooling the films rolled up entrapping the absorbed particles which could then be release by heating again [128]. This is a novel approach for the encapsulation and release of nanoparticles and cells with the addition of the magnetic particles allowing manipulation of the films by an external field [128].

Sukhishvil and Zhu showed a very interesting approach to film preparation using the assembly of diblock copolymer micelles [129]. A block copolymer consisting of PNIPAAm and poly(N-vinylpyrrolidone) (PVPON) was prepared [129]. These polymers formed micelles above the LCST of PNIPAAm with PNIPAAm cores [129]. Assembly of these micelles onto a substrate was achieved layer by layer with incorporation of poly(methacrylic acid) (PMAA) between the micelles [129]. Upon cooling, hydrogen bonding between the PMAA and PVPON in the micelle shells stopped the disassembly of the micelles [129]. These films showed good incorporation of hydrophobic drugs with a slower release above the LCST due to collapse of drug containing PNIPAAm cores [129].

Schmaljohann produced a series of graft copolymers of PNIPAAm, PDMAAm and PEG which showed suitable swelling/deswelling and surface changes for cell attachment and detachment in response to temperature [130]. Membranes of triblock copolymers of PNIPAAm-b-PVCL-b-PNIPAAm produced by Li et al. also displayed cell attachment and detachment in response to temperatures above and below the LCST, respectively [131].

Abraham et al. produced a P(NIPAAm-co-MMA) polymers and grafted this with phosphorylate HEMA [132]. Cells were found to adhere well to the graft copolymer and spread across the surface due to the phosphate moieties [132]. Upon cooling below the LCST, whole cell sheets were released from the surface [132].

Previously, the conjugation of insulin to PNIPAAm was studied by Chen et al. and, more recently, Hatakeyama demonstrated that grafting of the polymer to a surface allows cell growth at insulin levels 10 times lower than free insulin and allows for easy detachment when polymer is cooled below its LCST [133,134].

Ito et al. prepare plates of polystyrene grafted with PNIPAAm copolymers using a mask to create a micropatterned surface [135,136]. Cell growth and selected cell detachment was shown to be achievable with this approach [135,136]. Tsuda et al. also utilized patterning for cell culture by combing a patterned surface of PNIPAAm and PNIPAAm-graft-PBuMA with two different LCSTs. This substrate was able to selectively adhere cells at different temperatures allowing formation of cell sheet monolayers comprising two cell types in an ordered pattern [137]. A review on micro-patterning using stimuli responsive polymers by Ito gives a much more extensive overview of the subject [138].

The encapsulation of biological substrates into a PNIPAAm layer for controlled release during cell growth and also showed the possible release of whole cell sheets from the PNIPAAm grafted plates [139]. A review on cell sheet engineering by Kikuchi and Okano also showed the release of cell sheets from PNIPAAm films and the dependence on film thickness [140-142]. Films 20–30 nm show the best cell adhesion and detachment [140-142]. However films with thicknesses greater than 30 nm have minimal cell adhesion [140-142]. This is due to the hydration of the outer PNIPAAm layer being too hydrophilic and hence support less protein and cell adhesion [140-142]. This problem was addressed by Moran et al. who produced films of thicknesses greater than 30nm and coated them with cell adhesion promoters such as collagen of laminin [143]. Cells successfully adhered to these films and the thermal detachment of cells was not affected [143]. A further solution to this was demonstrated by Shimizu et al. who treated the PNIPAAm surface with an oxygen plasma which roughened the surface and removed its thermoresponsive ability making it hydrophobic at all temperature. Release of the cell layer at low temperatures takes the hydrophobic surface layer of PNIPAAm with it resulting in cell sheets with an increased mechanical strength [1].

Cheng et al. showed the production of plasma polymerized PNIPAAm films onto microheater arrays produced using photolithography [144]. This method allows for localized heating and specific area detachment of cells with many possible applications [144].

An interesting 3D cell culture method was envisaged by Poon et al. They synthesized chitosan-graft-PEG-graft-methacrylate copolymers which gelled at 37 °C but were also UV curable [145]. Cells were encapsulated layer by layer by heating each layer above the LCST and the final construct was cured with UV to enable cooling below the LCST without hydrogel breakdown [145].

Li et al. synthesized nanoparticles of thermoresponsive polymers by fast heating of a solution of the thermoresponsive polymer P(PEGMA-co-Boc-Cyst-MMAm) [146]. The particles produced were dependant on the heating rate during formation [147]. These particles possessed sensitivity to a reducing environment, such as the intracellular cytoplasm, by reduction of the disulfide bonds in the polymer chain resulting in breakdown of the nanoparticles [147]. This would allow for intracellular delivery of numerous therapeutic molecules [146].

Much work has been done on the modification of inorganic particles with thermoresponsive polymers [148-154]. Zhang et al. coated insoluble nanoparticles with PNIPAAm rendering them stable in aqueous solutions with temperature dependant solution properties and suggested uses in drug delivery and biological sensing [148]. Several groups have shown the formation of a novel drug delivery device with magnetic targeting capability by coating a magnetic nanoparticle, conjugated to active drug, with a thermoresponsive polymer. Upon heating above the LCST, the coating shrinks and releases the active drug [149,151,152,155,156]. Purushotham also showed the possibility of using a magnetic field to heat the composite and hence release entrapped drugs form the thermoresponsive coating [150]. Work by Papaphilippou et al. has involved the preparation of hybrid micelles utilizing a thermoresponsive PEGMA containing polymer and a magnetite nanoparticle core [154]. These micelles showed a thermoresponsive ability with LCST of 60 °C and a magnetic ability with future prospects aiming controlled drug delivery [154].

Pitch et al. demonstrated the stabilisation of magnetite nanoparticles by a thermoresponsive polymer [157]. In particular, microgels containing thermoresponsive PVCL and a hydrophobic monomer, acetoxy ethylmethacrylate, were prepared and loaded with the nanoparticles. These microgels showed thermoresponsive swelling/deswelling whilst stabilizing the suspension of nanoparticles [157].

A comprehensive review of combining magnetic particles and thermoresponsive polymers was written recently by Yuan et al. [158].

Das et al. showed another approach to polymer nanoparticle composites [159]. By incorporating gold nanorods into NIPAAm microgels, photo-induced heating of the gold nanorods provided enough heat to induce the phase transition of the microgel providing a possible photoinduced drug release [159]. With similar work on grafting NIPAAm onto gold nanoparticles being done by Pérez-Luna et al. [160].

Nanoparticles of crosslinked P(NIPAAm-co-AAm) were prepared by Fundueanue et al. by dispersion of the polymer in mineral oil and crosslinking with gluteraldehyde [161,162]. When prepared below the LCST, these particles were found to have a uniform pore size and sharp thermoresponse with a large increase in pore size when brought to body temperature showing a promising possibility for drug encapsulation and delivery [161,162].

Jun Wang et al. have produced a series of thermoresponsive micellar nanoparticles from PVCL and polyphosphoester and examined their cell change in size with temperature [163,164]. They found that as the temperature was increased above the LCST the micelles became more hydrophobic and formed aggregates in a reversible way [163,164]. The micelles were found to be both biodegradable and non cytotoxic showing a promising drug delivery vector [163,164].

Another class of nanoparticles are microgels. Microgels are microscopic sized crosslinked hydrogel particles with similar properties to crosslinked hydrogels. The advantage of a microgel is that, although it is crosslinked outside of the body, it is still easy to insert it into the body via injection. Vihola et al. synthesized PVCL and PVCL-graft-PEG microgels were formed by heating the polymer above its LCST and using salicylic acid as a crosslinker [7]. The salicylic acid formed hydrogen bonds between the polymer chains forming a physical hydrogel [7]. By adding a solution of polymer and drug to a solution containing the crosslinker at temperatures greater the LCST, hydrogel particles were formed which showed sustained release [7]. Interestingly, the PEG graft copolymers showed a slower drug release due to an increase in hydrogen bonding and hence increase packing from the PEG chains [7].