Negatively temperature-responsive hydrogels that undergo a sol-gel transition on warming above a lower critical solution temperature (LCST) have been extensively investigated for drug delivery applications, owing to their fast temperature-induced gelation without requiring organic solvents or chemical reaction. LCST can be due to a variety of mechanisms and can be determined by turbidity, rheology, differential scanning calorimetry, or spectroscopy.
Polymers of N-substituted acrylamides exhibit an LCST transition over a narrow temperature range in aqueous solution.32 Among these, poly(N-isopropylacrylamide) (poly(NIPAAm)) is by far the most commonly investigated, as it has a convenient LCST of about 30 °C, which can be controlled by the content and composition of additional comonomers.17, 18 Typically, poly(NIPAAm) gels tend to undergo a high degree of syneresis (i.e., shrinking and separation from solvent) on gelation, which can result in high burst release, particularly of hydrophilic drugs. Another disadvantage of poly(NIPAAm) homopolymer is its lack of degradability, and therefore, additional co-monomers must be included to make the gel degradable.
For a copolymer containing NIPAAm to form an injectable hydrogel, it must initially have an LCST below body temperature. Before injection below the LCST, this polymer is soluble in an aqueous solution. Then, on injection (body temperature > LCST), this polymer precipitates and forms a hydrogel if the polymer concentration is sufficient. The release rate from physical (i.e., not cross-linked) gels is inversely proportional to the gel's viscosity in the absence of intermolecular interactions between the polymer and the drug.23, 33 For a formulation to be degradable based on hydrolysis34–38 or enzymatic degradation39 of side chains, the LCST must increase to above body temperature on hydrolysis or enzymatic degradation and the polymer must not be sufficiently chemically cross-linked to remain as a solid below the LCST so that the polymer redissolves. In these types of materials, the LCST starts below the body temperature to make the material injectable, but becomes resorbed after the LCST increases to above body temperature. The use of dynamic LCST polymers based on NIPAAm was first demonstrated by Neradovic et al.34 using hydrolyzable poly(NIPAAm-co-hydroxyethylmethacrylate (HEMA)-monolactide), although these materials did not have a final LCST above body temperature. Similar materials were subsequently developed by Lee and Vernon37 using poly(NIPAAm-co-AAc-co-HEMA-lactide) and Ma et al.36 using poly(NIPAAm-co-lactide methacrylate-co-HEMA), which exhibited LCST increase to above body temperature on hydrolysis. In the former case, the degradation was rapid (2–8 days) due to increased hydration of the polymer chains owing to the hydrophilicity of acrylic acid (AAc), whereas Ma et al. reported much longer degradation times of about 200 days due to grafting of the lactide side group onto methacrylic acid, leading to a much more hydrophobic material. However, these materials also underwent rapid deswelling over the course of 2–6 h and had an equilibrium water content under 50%,36 likely rendering them unsuitable for delivery of hydrophilic drugs that will partition into the water phase as it leaves the hydrogel system during syneresis.
Alternatively, copolymers of NIPAAm containing hydrolyzable dimethyl-γ-butyrolactone acrylate (DBLA) were shown by Cui et al.38 to exhibit degradability without the loss of low-molecular-weight by-products. Copolymers containing NIPAAm, DBLA, and AAc were subsequently shown to have favorable biocompatibility in vivo.40 Li et al.41 used similar copolymers of NIPAAm, DBLA, AAc, and HEMA-poly(trimethylene carbonate) (HEMA-PTMC) to deliver the protein superoxide dismutase. The release profile was highly dependent on the protein loading. Gels physically mixed with collagen at high-protein loading (4 mg/mL) exhibited nearly constant protein release over a period of 3 weeks; however, all gels with low-protein loading (2 mg/mL) showed fast release over the first day and almost no release thereafter.
Another method allowing for gel degradation is cross-linking NIPAAm-based polymer chains with degradable cross-linkers such as poly(lactic acid),42 dextran,42, 43 or peptide sequences44, 45—however, in this approach, unless the LCST of the individual polymer chains increases to above body temperature, the poly(NIPAAm) portions will remain permanently insoluble at the injection site. In order to be injectable, cross-linked materials must have a very low cross-link density, as they will otherwise form solid, swollen hydrogels below the transition temperature.
As a poly(NIPAAm)-based gel is heated above its LCST, a homogeneous polymer-rich phase (i.e., the gel) separates from a fraction of the original solvent and then tends toward an equilibrium.46 Usually, solutions are formulated at <30 wt % and deswell such that the final polymer concentration in the gel is about 50%,47–50 so a significant fraction of water is lost during gelation. Rapid deswelling and syneresis (i.e., loss of the initially entrapped aqueous liquid) has been shown to be associated with fast drug release on heating above the LCST of NIPAAm-based physical gels.51–53 Accordingly, limiting syneresis on the LCST transition may provide slower release of hydrophilic drugs, which partition into the aqueous phase when the gel phase separates. Graft copolymer architecture is useful for controlling swelling with relatively minor effects on the material LCST.29, 54–57 We have demonstrated this concept using graft copolymers of poly(NIPAAm-co-Jeffamine® M-1000 acrylamide), of which some formulations exhibit almost no change in volume when heated above the LCST.53 Homopolymer gels reduced over 90% of the entrapped model drug ovalbumin within 3 h, while 16% release was observed over 6 days from gels containing Jeffamine grafts.
A promising and yet rarely used method for controlling drug release from these materials may be the copolymerization of a drug derivative as a hydrolyzable side group. Such a strategy was reported by Shah et al.58 using N-hydroxysuccinimide as a model drug in the polymer-drug conjugate poly(NIPAAm-co-N-acryloxysuccinimide). These materials have the advantage of releasing drug in concert with the rate of degradation of the bond linking the drug to the polymer backbone, and degradation leaves an acid group on the polymer backbone, raising the LCST and allowing for dissolution of the polymer. However, this strategy is only compatible with certain drugs (those containing hydroxyl groups).
pH-Responsive Poly(N-isopropylacrylamide)-Based Hydrogels
Dual-responsive polymers sensitive to both temperature and pH have been developed by copolymerizing NIPAAm with either AAc55 or pH-sensitive substituted acrylates such as 2-(dimethylamino)ethyl methacrylate (DMAEMA).59 In terms of drug delivery, it is convenient to have a pH-induced transition at 37 °C that occurs over the neutral range of pH around 5.0–7.4.60 Using propylacrylic acid as a second monomer having pKa near 6.0, Garbern et al.61 reported prolonged release of active vascular endothelial growth factor (VEGF) up to 4 weeks at pH 5.0 and for between 1 and 4 days at pH 7.4, despite the polymer being soluble. Such prolonged release at low pH may have been due to the high content of propylacrylic acid (17 mol %) affecting the overall properties of the polymer in its protonated (insoluble) state. The somewhat prolonged release at neutral pH despite the polymer being soluble was attributed to electrostatic affinity between the VEGF (isoelectric point pI = 8.5) and the negatively charged polymer.
Block Copolymer Hydrogels
Solutions of block copolymers with alternating hydrophilicity (usually ABA triblock or alternating multiblock copolymers) have LCST in aqueous solution, which is dependent on block length, composition, and polymer concentration. Typically, relatively high polymer concentrations (>10 wt %) are required for these materials to exhibit LCST behavior, although exceptions do exist.62 These materials often have a central hydrophobic block such as poly(propylene oxide) (PPO) or poly(lactide-co-glycolide) (PLGA) and hydrophilic blocks, which are almost always comprised of poly(ethylene oxide) (PEO, also called PEG for poly(ethylene glycol)). The LCST of these materials is thought to be due to increased hydrophobicity of the hydrophobic segment on heating, leading to micellar aggregation.63, 64
Much of the early work on this family of polymers was based on PEO-PPO-PEO block copolymers such as the Pluronics made by BASF,63–68 but these are nondegradable, mechanically weak, and tend to be highly permeable to drugs.68 The first degradable block copolymer hydrogels with LCST behavior for drug delivery were reported by Jeong et al.69 who used poly(L-lactic acid) (PLLA) as the hydrophobic central block. Unlike poly(NIPAAm)-based systems, these polymers are backbone degradable and become soluble after a single ester is hydrolyzed. These gels were shown to release FITC-dextran (20 kDa) over at least 12 days with very low initial burst release. Increasing the polymer concentration from 25 to 35 wt % led to a more constant-release profile in vitro.69 Similar materials of PEG-PLGA-PEG (500-2810-500) were shown to have convenient LCST below body temperature above about 16 wt %.70 These materials released the hydrophilic low-molecular-weight drug ketoprofen mostly over the first 3–5 days in vitro, with very low release thereafter. The release rate was not tunable over a wide time frame, with similar release observed from 20, 25, and 33% gels. Using the same copolymers, the hydrophobic drug spironolactone was released in an S-shaped release profile, with the first part being due to diffusion over about 4 days, followed by almost no release for the next 10 days, and then accelerating release thereafter for the next 35 days due to degradation of the gels. Much work over the last decade has used this block copolymer design with various degradable hydrophobic groups and various architectures, such as PEG-PTMC diblocks,71 poly(caprolactone) (PCL) groups,72, 73 modified Pluronics,74–76 enzyme-degradable poly(amino acid)s,77 and even hydrophobic segments of poly(NIPAAm),78, 79 which give rise to a sharp gelation temperature based on the LCST of the NIPAAm chains.
For drug release, block copolymer hydrogels are rather flexible in terms of the types of drugs that can be loaded. A formulation of PLGA-PEG-PLGA developed under the name ReGel® showed nearly linear in vitro release of paclitaxel over 50 days with about 20% initial burst release over the first day at 23 wt % concentration.5 Similar gels showed first-order release of various proteins over 1–2 weeks. The difference in release among these can be explained by the fact that hydrophobic drugs are relatively immobile due to poor solubility and hence exhibit partition-controlled release, whereas hydrophilic drugs (either small molecules or proteins) tend to release more quickly and in a first-order profile based on diffusion. Partition-controlled release refers to the phenomena that hydrophobic drugs preferentially dissolve (partition) in the hydrophobic polymer phase and only slowly diffuse into the water phase and thus become released due to their low water solubility. The release of the protein GLP-1 was slowed from about 2–15 days and made nearly linear by complexing the drug with zinc, thereby reducing its solubility while preserving its activity.80 Injection of the ReGel® into diabetic rats led to reduced blood glucose levels for 15 days postinjection. A similar system using cationic poly(β-amino ester) (PAE) as the outside blocks of a PAE-PCL-PEG-PCL-PAE pentablock copolymer resulted in electrostatic retention of insulin and a linear release profile for up to 20 days in vitro, with sustained insulin concentrations provided in vivo for up to 18 days.81 Still, this class of materials tends to release hydrophilic drugs very quickly. For example, Gong et al.82 reported PEG-PCL-PEG gels that released the hydrophilic model drug VB12 completely over about 24 h, with high burst and incomplete release of both bovine serum albumin (BSA) and the hydrophobic model drug honokiol after 14 days (about 20% cumulative release), although it is worth noting that the release was not measured throughout the 50-day degradation of the gels. Tang and Singh83 synthesized 11 variants of mPEG-PLGA-mPEG and reported release of lysozyme over 20 days at the longest from very high-concentration (40 wt %) gels.
pH-Responsive Block Copolymer Hydrogels
Block copolymer hydrogels can be made to respond to both temperature and pH. Determan et al.84 modified Pluronic F127 with end blocks of poly(2-diethylaminoethyl-methyl methacrylate), having pKa near 7.5. They showed linear release of lysozyme over about 4 days at pH 7.0 and slightly slower release, projected to last about 6 days, at pH 8.0. Pluronic® P104-based gels with acid-sensitive acetal linkages were reported by Garripelli et al.85 They reported release of FITC-dextran (40 kDa) over 2 days at pH 5.0, 9 days at pH 6.5, and about 30 days with an S-shaped release profile at pH 7.4. Such pH-dependent release might be useful in terms of releasing drugs more quickly in response to local acidic conditions.
Hybrid Temperature-Responsive Materials
Natural and natural/synthetic hybrid temperature-responsive materials that thicken or gel on warming to body temperature have been reported, usually with the goal of prolonging release using high-molecular-weight natural polymers or for combining the advantages of high biocompatibility and degradability of natural materials with the control over composition and properties of synthetic materials.
Aqueous solutions of chitosan in the presence of glycerophosphate (GP) salts can undergo gelation on heating at neutral pH.86 Gelation time is affected by salt concentration, temperature, degree of deacetylation, and slightly by drug loading.87 These materials were shown to release paclitaxel in vitro at a rate, which was dependent on the drug loading, with 40% cumulative release in 30 days when loaded at 64 mg/mL.88 When evaluated in vivo in a subcutaneous tumor model, gels showed some antitumor efficacy both in the presence and absence of drug, indicating an antitumor effect of chitosan alone.88 Gels with drug showed tumor volumes at 17 days that was similar to that of daily intravenous injections for the first 4 days. However, it is unclear whether the higher efficiency (similar outcome using a single administration) of the gels is related to the material or simply the intratumoral administration of the drug in the gel group. A notable disadvantage of these natural materials is that the release rate of the drug is not easily controlled by adjusting the properties of the device. Although the authors note that the drug itself could be modified to alter the release (in this case, to increase the release rate for a high drug loading), this approach is undesirable in terms of cost and regulatory issues. Cross-link density of chitosan-GP interpenetrating polymer networks (IPNs) with hydrophilic poly(vinyl alcohol) was used to control the release of lysozyme,89 though it is worth noting that covalently cross-linked gels are likely to not be injectable or to break into many pieces on injection through a needle, and the poly(vinyl alcohol) used was of relatively high molecular weight (66 kDa), which could lead to renal toxicity.
Another common design for hybrid temperature-responsive hydrogels is to graft temperature-responsive synthetic polymers, such as poly(NIPAAm), onto a natural, hydrophilic polymer such as hyaluronic acid,90 gelatin,91 or chitosan.92–94 The poly(NIPAAm) chains in these materials tend to be of low molecular weight, and so the resulting gels tend to be very weak and viscous. Further, the drug release is typically fast. A variety of low-molecular-weight drugs were shown to release from NIPAAm-grafted gelatin within 6 h,91 and proteins released from NIPAAm-grafted hyaluronic acid within 12 h.90 Similarly, fast release over 2 days or less was observed for NIPAAm-grafted chitosan for various model drugs92–94—to prolong the release of hydrophobic drugs to 14 days, the authors embedded PLGA microparticles within the gel.94 In addition to the short duration of drug release so far demonstrated from this class of materials, another disadvantage is that the thermosensitive polymer chains in these designs will remain insoluble after degradation of the natural component.
Another group of linear polymers having LCST-like behavior in aqueous solution are the poly(organophosphazene)s, which can be made degradable through the incorporation of amines, amino acids, or alkoxy groups. These gels thicken over a wide range of temperatures and remain transparent rather than phase-separating like most other LCST materials, and so the thermal transition is typically characterized by viscometry. Drug release can be controlled either by incorporation into the polymer95, 96 or simply by diffusion or degradation, over a period of up to about 35 days.7 Kang et al.97 used hydrophobic side groups of L-isoleucine ethyl ester and hydrophilic side groups of α-amino-ω-methoxy-PEG (molecular weight (MW) 550) and reported duration of release for two model protein drugs to be between 3 and 15 days. Similar gels released the hydrophobic low-molecular-weight drug doxorubicin steadily over about 30 days.98
Elastin-like polypeptides (ELPs) contain pentapeptide repeat units Val-Pro-Gly-X-Gly, where X is any natural amino acid except proline.8 These materials have an LCST that tends to occur over a narrow temperature range. Because they are genetically encoded, they can be made to be monodisperse, and are enzymatically degradable in vivo. Although these materials have also been investigated as thermally triggered polymer-drug conjugates,99–102 hydrogels have also been made from these materials for injectable drug delivering matrices, primarily complexed with silk to comprise so-called silk-elastin like polypeptides or SELPs. These materials show injectability due to reduced crystallinity of the silk component by mixing with ELPs.103 However, the release from these matrices is rapid, occurring almost completely in a matter of h, even for large protein drugs.104