Injectable biodegradable copolymer hydrogels, which exhibit a sol–gel phase transition in response to external stimuli, such as temperature changes or both pH and temperature (pH/temperature) alterations, have found a number of uses in biomedical and pharmaceutical applications, such as drug delivery, cell growth, and tissue engineering. These hydrogels can be used in simple pharmaceutical formulations that can be prepared by mixing the hydrogel with drugs, proteins, or cells. Such formulations are administered in a straightforward manner, through site-specific control of release behavior, and the hydrogels are compatible with biological systems. This review will provide a summary of recent progress in biodegradable temperature-sensitive polymers including polyesters, polyphosphazenes, polypeptides, and chitosan, and pH/temperature-sensitive polymers such as sulfamethazine-, poly(β-amino ester)-, poly(amino urethane)-, and poly(amidoamine)-based polymers. The advantages of pH/temperature-sensitive polymers over simple temperature-sensitive polymers are also discussed. A perspective on the future of injectable biodegradable hydrogels is offered.
Hydrogels are three-dimensional hydrophilic polymeric networks that can absorb and retain a considerable amount of water with maintenance of shape.1, 2 Injectable biodegradable hydrogels have been widely used in biomedical applications, such as drug/cell delivery and tissue engineering, because of their highly hydrophilic characteristics. Such hydrogels are of particular interest because drugs, proteins, and cells can be easily incorporated into polymer solutions prior to administration. Importantly, no surgical procedures are required for the insertion of gels into the body; the gels are administered by simple injection.3–5
Injectable hydrogels can be formed in situ by either chemical or physical crosslinking methods. Chemically crosslinked hydrogels, prepared through photopolymerization,6 disulfide bond formation,7 or reaction between thiols and acrylate or sulfones,7 undergo significant volume changes during the phase transition. In contrast to chemical hydrogels, physically crosslinked hydrogels, formed by the self-assembly of polymers in response to environmental stimuli (for example, temperature, pH, or both) display sol–gel transitions without marked volume changes. The sol–gel transition systems are relatively low-viscosity aqueous solutions (sol state) prior to injection, but rapidly convert into a gel under physiological conditions post-injection.1–3 Physical hydrogels have several advantages over chemical hydrogels, because they do not require photo irradiation, use of organic solvents or crosslinking agents, and do not release heat during polymerization at the gelation site, which may denature incorporated proteins and damage embedded cells and surrounding tissues. Thus, the physical systems have recently attracted increased attention. For use in drug/cell delivery and tissue engineering, hydrogels should be low-viscosity solutions (free-flowing) prior to subcutaneous injection, and should rapidly gel in the human body, where ultimate degradation of the hydrogels is desired.
This article provides insights into recent advances in synthesis and biomedical applications of injectable, biodegradable, polymeric hydrogels that exhibit sol–gel transitions in response to temperature and pH/temperature changes. The discussion covers poly(ethylene glycol) (PEG)/polyester block copolymers, polyphosphazenes, polypeptides, chitosan, polymers based on sulfamethazine, poly(β-amino ester), poly(amino urethane), poly(amidoamine), and others.
Hydrogels that are sensitive to temperature are useful for both in vitro and in vivo applications, because temperature control is generally easy. Temperature-sensitive hydrogels undergo a sol–gel phase transition when the temperature is increased from room temperature to physiological temperatures.
Poly(ethylene glycol) (PEG)/Polyester
Copolymers of hydrophilic biocompatible PEG with biodegradable biocompatible aliphatic polyesters, for example polylactide (PLA), polyglycolide (PGA), poly(ε-caprolactone) (PCL), or poly[(R)-3-hydroxybutyrate] (PHB), have received increasing attention as promising biomaterials.
The first reported biodegradable thermosensitive hydrogel was PEG-poly(L-lactide)-PEG (PEG-PLLA-PEG) (Scheme 1).9 PEG-PLLA-PEG was synthesized by the ring-opening polymerization (ROP) of L-lactide using monomethoxy PEG (MPEG) of molecular weight (MW) 5 000 as macroinitiator, to form MPEG-PLLA diblock copolymers. Triblock copolymers (PEG-PLLA-PEG), with PLLA blocks in the 2 000–5 000 Da molecular weight range were next produced by coupling the resulting diblock copolymers using hexamethylene diisocyanate (HMDI). The concentrated copolymers (10–30 wt.-%) dissolved in water demonstrated a gel-to-sol transition with increases in temperature. The gel-to-sol transition was precisely controlled by the biodegradable PLLA block length when a PEG block was fixed at both ends. The gel-to-sol properties of the PEG/polyester diblock and triblock copolymers in water depended on the hydrophobic/hydrophilic balance, block length, hydrophobicity, and stereoregularity of the hydrophobic block.10, 11
PEG-(D,L-lactide)-PEG (PEG-PLA-PEG) triblock copolymers (MW of MPEG = 2 000–5 000) were recently produced by coupling MPEG-PLA diblock copolymers using adipoyl chloride.12 Rheological measurements showed a gel-to-sol transition in concentrated polymer solutions, with an increase in temperature. Gelation at lower temperatures was attributed to hydrogen bonding between PEG blocks, and the hydrogen bonds were broken at elevated temperatures, leading to the gel-to-sol transition. A series of biodegradable star-shaped PLLA-PEG block copolymers were obtained by the pairing of two components: star PLLA and monocarboxy-MPEG, using dicyclohexylcarbodiimide (DCC).13, 14 The gel-to-sol transitions, induced by increases in temperature, were observed at concentrations beyond the critical gelation concentration (CGC). With the same PEG block length, use of longer PLLA blocks led to a decrease in the CGC, thus widening the gelation window. Gelation was attributed to micellar packing and the breaking of micellar packing structures caused by the partial hydration of the PEG block, which resulted in the gel-to-sol transition at higher temperatures.
However, the abovementioned gels underwent gel-to-sol transitions, reducing their suitability for encapsulation of some drugs or proteins. In addition, injection at temperatures that are elevated with respect to body temperature is uncomfortable for patients. PEG-poly(D,L-lactide-co-glycolide)-PEG (PEG-PLGA-PEG) triblock copolymers with short PEG blocks (MW ≤ 750) were soluble in water at low temperatures and converted into a gel at elevated temperatures (Scheme 2).15 The triblock copolymers were shown to be biocompatible with blood.16–18 The gelation window spanned the physiological temperature range (Figure 1). 13C NMR and dynamic light scattering (DLS) studies revealed that the gelation of PEG-PLGA-PEG was enhanced by micellar growth and close packing of micelles (Figure 2). The upper gel-to-sol transition at higher temperatures was driven by the breakage of the micellar structure caused by the partial dehydration of PEG and PLGA blocks. The sol–gel transition could be adjusted by variation in PLGA and PEG block lengths, PLGA composition, and the use of particular additives. In other studies,19, 20 the gelation mechanism of PEG-PLGA-PEG (550-2810-550) was investigated by rheology, DLS, differential scanning calorimetry (DSC), and small-angle neutron scattering (SANS). The results indicated that the macroscopic liquid–liquid phase separation induced gelation of the triblock copolymer. A transparent gel was formed in situ after injection of a 33 wt.-% PEG-PLGA-PEG (550-2810-550) aqueous solution into rats, and gel integrity persisted for one month.21
Interestingly, thermosensitive hydrogels based on PLGA-PEG-PLGA (BAB-type) (Scheme 3) showed a sol-to-gel transition similar to that of PEG-PLGA-PEG (ABA-type) copolymer hydrogels.22–24 However, the synthetic procedure for BAB-type hydrogels was simpler than that for the ABA-type, using HMDI as a coupling agent. The gelation mechanism for BAB-type hydrogels was different from that of ABA-type hydrogels because of the presence of two PLGA end blocks. The former formed micelles with intermicellar bridges, whereas the latter formed regular micelles with a PLGA block core and a PEG block shell, in aqueous solution. At temperatures below the critical gelation temperature (CGT), some bridging micelles of PLGA-PEG-PLGA copolymers formed, but they were not stable because of the low hydrophobicity of PLGA. With increasing temperatures to the CGT, a bridged micelle network was formed because of an increase in the hydrophobicity of the PLGA segment, leading to gelation (Figure 3). In vitro and in vivo degradation of the PLGA-PEG-PLGA (1500-1000-1500) copolymer (ReGel) was studied.23 A 23 wt.-% ReGel was found to degrade completely in vitro after 6–8 weeks at 37 °C, whereas no hydrogels were observed at the end of the fourth week after subcutaneous injection of a 23 wt.-% ReGel solution into rats.
To overcome the molecular weight constraints of the PLGA-PEG-PLGA and PEG-PLGA-PEG triblock copolymers, and to control the time of persistence of gels, graft copolymers of PEG-g-PLGA and PLGA-g-PEG were investigated (Scheme 4).25–27 The former copolymer was synthesized by the ROP of D,L-lactide (LA) and glycolide (GA) using a hydroxy-pendant PEG as a macroinitiator. The PEG-g-PLGA copolymer in water showed a sol–gel transition at concentrations above 16 wt.-%. Gel integrity persisted for one week under physiological conditions. The latter copolymer was prepared by the one-step ROP of LA, GA, and epoxy-terminated PEG.25, 26 Aqueous solutions of PLGA-g-PEG copolymers also exhibited a sol-to-gel transition upon heating. After injection of the copolymer solution (29 wt.-%) into rats, the gel persisted for more than two months, a significant improvement over the one week persistence time observed for the PEG-g-PLGA copolymer hydrogel. The gelation mechanism of the PLGA-g-PEG copolymer was examined by 13C NMR spectroscopy, SANS, rheology, and infrared (IR) spectroscopy.28 Partial dehydration of PEG caused micellar aggregation, leading to gelation, and significant dehydration of PEG led to macroscopic separation at the gel-to-sol transition temperature. The sol-to-gel transition temperature was tuned from 15 to 45 °C by tailoring the polymer composition, the number of PEG grafts present, and the PEG molecular weight. By mixing these copolymers at different composition ratios, the sol-to-gel transition temperature could be adjusted, and the gel duration could be varied from one week to three months. Both copolymers yielded rather soft gels with storage moduli (G′ values) of 100 Pa.26
Poly(ε-caprolactone) (PCL) is a hydrophobic crystalline polymer that is both biodegradable and biocompatible. PCL copolymers have a powdery morphology, making them easier to handle than are PLGA and PLLA copolymers, which have a sticky paste morphology. Aqueous solutions of MPEG-PCL diblock copolymers (MPEG MW ≥ 2 000) underwent a gel-to-sol transition on temperature increase. The gelation window was strongly influenced by PEG and PCL block length. In vivo gelation of MPEG-PCL (MW 2 000–2 300) was also studied. When a 23 wt.-% copolymer solution at 42 °C was injected into a rat, the gel formed immediately and persisted for one month with only a small amount of inflammation at the injection site.29 MPEG-PCL diblock copolymers with a lower molecular weight MPEG (MW = 750) were subsequently reported.30 Interestingly, synthesized diblock copolymers with PCL block lengths of 1 400–3 000 were soluble in water and underwent a sol-to-gel-to-sol transition as a result of temperature changes. The sol–gel phase diagram depended on PCL block length. Recently, PEG-PCL-PEG and PCL-PEG-PCL (Scheme 5) triblock copolymers have been introduced, and aqueous solutions thereof showed a clear sol-to-gel-to-turbid sol transition with an increase in temperature.31, 32 DLS and 13C NMR studies indicated that the clear sol-to-gel transition arose from micellar aggregation, whereas the gel-to-turbid sol transition was driven by the breakage of the core–shell structure. Because of differences in structural topology, PCL-PEG-PCL had a higher G′ (10 000 Pa) (Figure 4) than did PEG-PCL-PEG (100 Pa) (Figure 5). However, because of crystallization of the PCL block, a clear aqueous solution of PCL-PEG-PCL (20 wt.-%) turned turbid within 1 h at 20 °C, which may be problematic for injectability. The crystallization problem was, however, solved by using poly(caprolactone-co-trimethylene carbonate)-PEG-poly(caprolactone-co-trimethylene carbonate) (PCTC-PEG-PCTC).33 A PCTC-PEG-PCTC triblock copolymer in water exhibited a sol-to-gel-to-syneresis transition upon heating. However, a very soft gel with a G′ of 1 Pa was obtained. This gel was stable with respect to hydrolysis in phosphate buffers at 37 °C for 50 d, but degraded markedly in rats. In addition, the crystallizability of PCL decreased as the molecular weight of PCL rose. Therefore, a multiblock copolymer of (PCL-PEG-PCL)n was prepared by coupling PCL-PEG-PCL triblock copolymers (1 000-1 000-1 000), using terephthaloyl chloride. A 20 wt.-% aqueous solution of the multiblock copolymer underwent a sol-to-gel-to-sol transition that did not turn turbid at room temperature, indicating that the material may be convenient for practical applications, such as drug formulation and injection. The multiblock showed a lower G′ (100 Pa) than did the triblock (10 000 Pa). The gelation mechanism was attributed to crystallization of the multiblock copolymer.34
The multiblock topology affected the gelation behavior of other copolymer hydrogels. A series of PEG/PLLA alternating multiblock copolymers was synthesized by coupling PEG (MW = 600) to PLLA (MW = 1 100–1 500) using succinic anhydride.35 The multiblock PEG/PLLA aqueous solution underwent a sol-to-gel-to-sol transition with increasing temperature. The transition temperature and gel modulus could be controlled by varying the PLLA block length, PEG molecular weight, and PEG/PLLA ratio. The gelation mechanism was considered to be micelle aggregation. Stereochemistry also affected gelation.36 The PEG/poly(D,L-lactide) (PDLLA) and PEG/PLLA multiblock copolymers with identical block lengths and total molecular weight were prepared. Relative to amorphous PEG/PDLLA, the stereoregular PEG/PLLA multiblock copolymer had a reduced CGC, a lower sol-to-gel transition temperature, a broader gel area, and a larger maximal gel modulus. 13C NMR, X-ray diffraction (XRD), and UV-visible spectroscopy studies indicated that the different gelation properties could be attributed to the slower dynamic molecular motion of the methyl groups in PLLA. The isotactic arrangement in PLLA induced strong aggregation of the PEG/PLLA multiblock. In contrast with the sol-to-gel transition behavior of the PCL-PEG-PCL (1 000-1 000-1 000) multiblock copolymer,34 the PCL-PEG-PCL multiblock copolymer, with a higher molecular weight of both PEG and PCL, exhibited a gel-to-sol transition upon heating.37 In addition, multiblock copolymers of PEG-sebacate (PEG-SA) were synthesized by simple condensation polymerization.38 A soft gel formed when a 25 wt.-% aqueous solution of the PEG-SA multiblock was heated to 37 °C, and gel integrity persisted for more than three weeks in phosphate buffer, pH 7.4, at 37 °C.
In addition, it was found that a stereocomplex of the enantiomeric triblock copolymers (10 wt.-%), PLLA-PEG-PLLA (1 300-4 600-1 300), and poly(D-lactide)-PEG-poly(D-lactide) (PDLA-PEG-PDLA) (1 100-4 600-1 100), in water could induce temperature-dependent gelation, although the individual enantiomeric copolymers did not show thermal gelation. Stereocomplex formation during gelation was confirmed by wide-angle X-ray scattering (WAXS).39 A mixture of PEG-PLLA-PEG (2 000-2 000-2 000) and PEG-PDLA-PEG (2 000-2 000-2 000) (35 wt.-%) exhibited a gel-to-sol transition with an increase in temperature. WAXS results showed that gelation was not attributable to complexation of PLLA and PDLA, but rather to interdigitation of the helical PEG chains induced by the complementary arrangement of PDLA and PLLA helices.40 Enantiomeric PEG12500-(PLA)2 and PEG21800-(PLA)8 aqueous solutions showed gel-to-sol transitions with increasing temperatures. Although neither enantiomeric polymer solution formed a gel, a stereocomplex with a PLA block could form a gel at the same concentration. Rheology studies revealed that PEG-(PLA)8 showed a higher gel modulus than PEG-(PLA)2, because of the higher crosslinking density of the former. The PLA block length, PEG content, polymer topology, and polymer concentration influenced the gel–sol transition, gel modulus, and kinetics of gelation.41, 42 Subsequently, the PEG-PLLA and PEG-PDLA multiblock copolymers were developed by coupling the corresponding triblock copolymers using diisocyanatobutane.43 The multiblock copolymers showed a much lower CGC, faster gelation, and a higher gel storage modulus, compared with the parent triblock copolymers, because of an increase in the crosslinking density of the multiblock copolymers.
The end groups of the thermosensitive copolymers also influenced gelation.44, 45 A series of PLGA-PEG-PLGA triblock copolymers with several different end caps (hydroxy, acetyl, propionyl, and butanoyl groups) was synthesized and characterized. Triblock copolymers containing acetate and propionate groups exhibited a sol-to-gel transition as a function of temperature, whereas the copolymer containing the butyrate group precipitated in water. An increase in the hydrophobicity of the copolymer lowered the transition temperature, and the end cap caused a significant change in the position of the gelation window. Cholesterol end-capped star PEG-PLLA copolymers (above 3 wt.-%) in water also exhibited thermal gelation, but PEG-PLLA itself did not.46 Gelation was induced by the strong hydrophobic association of the cholesterol group.
Poly[(R)-3-hydroxybutyrate] (PHB) is a natural biodegradable polyester produced by bacteria. The crystallinity and hydrophobicity were higher than those of other synthetic polymers, such as PLA and PCL. A thermosensitive and amphiphilic poly(ether ester urethane) multiblock copolymer consisting of PHB, PEG, and poly(propylene glycol) (PPG) (PEG/PPG/PHB) was synthesized.47, 48 This copolymer, in aqueous solutions and at very low concentrations (2–5 wt.-%). showed a sol-to-gel transition as a function of temperature change. The gelation of the copolymer solution was associated with micellar packing.
Biodegradable polyphosphazenes, consisting of a hydrophilic PEG block and hydrophobic amino acids or a peptide block, such as L-isoleucine ethyl ester (IleOEt), D,L-leucine ethyl ester (LeuOEt), L-valine ethyl ester (ValOEt), or di-, tri-, and oligo-peptides in the side groups, were synthesized (Scheme 6).49–51 Aqueous solutions (10 wt.-%) of polyphosphazenes with MPEG350 and IleOEt exhibited a sol-to-gel transition as a function of temperature.49 The maximal viscosity was 30 Pa · s at 37 °C. The gelation properties were adjusted by varying the composition of the substituents, MPEG molecular weight, and concentration. In a subsequent study, polyphosphazenes with oligopeptides (tri- or tetra-peptides) and MPEG 350 as side groups were also found to exhibit a phase transition.50 The gelation properties of the thermogelling polyphosphazenes depended on the structure of the oligopeptide and the hydrophobic side groups. Such gels also showed higher gel strength compared with former gels.49 Intermolecular association of hydrophobic oligopeptides was responsible for thermally induced gelation of the copolymer. The sol–gel transition and gel strength of polyphosphazenes were also modulated by the blending of hard and soft polymers. By mixing the two polymers at Tmax (the temperature at which the viscosity is maximal) values of 31 and 42 °C, with blend ratios of 2:2, the Tmax of the blended polymers was observed to be 35–41 °C.52 The degradation rate could be controlled by the content of incorporated depsipeptides. Polyphosphazenes, incorporated with the depsipeptides, degraded faster than those in gels without depsipeptides, because the hydrolysis of depsipeptides produced carboxylic acid, which triggered degradation of hydrophobic amino acid groups.51
Polypeptides are important biomaterials offering favorable characteristics, such as biocompatibility and biodegradability. Polypeptides can interconvert among a variety of conformations, such as α-helix, β-sheet, and random coil, and building blocks with hydrophobic, hydrophilic, ionic, and non-ionic characteristics can be synthesized. An artificial triblock protein with short leucine-zipper end blocks flanking a water-soluble polyelectrolyte domain underwent reversible gelation in response to changes in pH and temperature.53 Gelation of the triblock protein was driven by formation of coiled-coil aggregates of the terminal leucine-zipper domains, and the gel changed to a viscous solution when coiled-coil aggregates dissociated with increasing pH and temperature. The thermally induced hybrid hydrogels were prepared by combination of water-soluble synthetic polymers and engineered proteins.54 The proteins were used as crosslinkers for synthetic polymers. Changes in protein conformation as a result of temperature changes triggered the formation of hybrid hydrogels. Diblock copolypeptide amphiphiles, consisting of charged and hydrophobic blocks, were synthesized.55 Aqueous solutions could form thermally stable gels (up to 90 °C), but the gels rapidly broke down under an applied stress. Gelation was believed to be promoted by association of the hydrophobic domains, which was triggered by the ordered packing of α-helical and β-strand segments, whereas rapid recovery after stress was attributable to the nature of the physical gelation process and the low molecular weight of the copolypeptides. A 10 wt.-% β-lactoglobulin aqueous solution showed a sol-to-gel transition at 85 °C.56 β-lactoglobulin is a component of milk whey, with two disulfide bonds and one free sulfhydryl group. Gelation was attributed to the formation of hydrophobically linked aggregates, followed by formation of disulfide-bonded aggregates. A synthetic polypeptide of poly(ferrocenylsilane)-poly(γ-benzyl-L-glutamate) (PFS-PBLG) was soluble in hot toluene, but formed a transparent gel at room temperature.57 The gelation process arose from formation of an α-helix and random-coil structure in the diblock copolymer in toluene.
VPGVG (V = valine, P = proline, and G = glycine) is a prominent amino acid sequence in elastin. A hexahistidine metal-binding motif was incorporated into an elastin-like polypeptide.58 A 6 wt.-% aqueous solution of the polymer was a clear at 4 °C, but changed to a gel at 25 °C. This material may be useful in heavy metal removal applications. Copolymers that contained a collagen peptide and an elastin peptide formed a thermally induced gel in water.59 Polypeptides with 82–86 mol-% VPGVG composition showed a sol-to-gel transition when the temperature was increased. The collagen acted as a hydrate unit and the elastin peptide acted as a thermosensitive crosslinking point. A de novo-designed peptide showed a thermosensitive gelation transition.60 A 2 wt.-% aqueous solution of the MAX3 peptide was a rigid gel at 75 °C with G′ of 1 100 Pa. A transition from a random coil to a β-hairpin produced a hydrogel network as the temperature increased. A sol-to-gel transition was also evident in aqueous solutions (above 3 wt.-%) of amphiphilic poly(N-substituted α/β-asparagines) (Scheme 7).61 The phase diagram was strongly influenced by hydrophilic blocks (amino alcohols) and polymer concentration. Recently, poly(alanine)-poloxamer-poly(alanine) (PA-PLX-PA) was synthesized as a thermosensitive hydrogel.62 Aqueous solutions of PA-PLX-PA underwent a sol-to-gel transition as the temperature increased. The sol-to-gel transition temperature was influenced by the molecular weight of each block and by the composition of PA. Based on FT-IR, DLS, 13C NMR, circular dichroism (CD), transmission electron microscopy (TEM), and fluorescence spectroscopy studies, the PA transition from random coil to β-sheet and the decrease in molecular motion of PLX resulted in gelation. The hydrogels were stable in phosphate buffer but degraded quickly in the presence of enzymes. A new thermogelling poly(N-vinyl pyrrolidone)-PA (PVP-PA) has been reported.63 Aqueous solutions of the polymers showed sol-to-gel transitions. Gel formation was attributed to hydrophobic association and formation of a β-sheet structure.
Chitosan, a polysaccharide derived from the partial deacetylation of chitin from crustacean shells, has been widely used as a biomaterial because of biodegradability, biocompatibility, non-toxicity, and bioadhesive properties (Scheme 8). Chitosan was approved by the US Food and Drug Administration and has been used in drug delivery, tissue engineering, and cosmetics.64, 65 An injectable thermogelling hydrogel was prepared by the combination of chitosan and β-glycerol phosphate (C/GP).66 Chitosan was dissolved in hydrochloric acid, and a GP solution was then slowly added to obtain a clear solution. At pH 7.15, the C/GP aqueous solution remained in a clear liquid state, but gelled rapidly in the vicinity of 37 °C when heated. The gelation temperature increased as the degree of deacetylation decreased, but was not affected by the molecular weight of the chitosan. Gelation was driven by hydrophobic association of the neutral chitosan molecules, promoted by the influence of GP on water at elevated temperatures. When subcutaneously injected into rats, the C/GP solution rapidly gelled. PEG-g-chitosan aqueous solutions exhibited thermal gelation.67 A 3 wt.-% solution of chitosan grafted onto 55 wt.-% PEG showed an increase in viscosity at 37 °C, from < 1 Pa · s to > 6 Pa · s within 1 250 s, whereas the chitosan solution alone did not show any change in viscosity, even after 3 500 s. A network gel resulted from the association of chitosan molecules and reduction in the mobility of the PEG segments at high temperatures. Pluronic was grafted onto chitosan (C-g-P), and the resulting C-g-P in aqueous solution displayed a sol-to-gel transition as the temperature increased.68 The gelation temperature was controlled by chitosan content, and gelation did not occur when the chitosan content was >17 wt.-%. More recently, hydrophobic N-palmitoyl moieties were grafted onto chitosan (NPCS) to produce a pH-triggered hydrogel within the pH range of 6.5–7.0.69 The G′ of the NPCS aqueous solution at pH 6.5 was about 100 Pa (Figure 6) and was strongly dependent on the shear rate, indicating that the NPCS aqueous solution might be able to squeeze through the needle during injection. The G′ of the hydrogel at pH 7.0 was higher than that at pH 6.5. This material was non-toxic in vitro, but a large degree of inflammation was observed at the interface between the tissue and the hydrogel after two weeks of implantation. No chronic inflammation was observed after six weeks of implantation. A balance between charge repulsion and hydrophobic interactions in NPCS aqueous solutions in the pH range 6.5–7.0 was involved in the gelation process.
Other Thermosensitive Block Copolymers
Poly(trimethylene carbonate) (PTMC) is biodegradable, biocompatible, and has soft mechanical properties. A PEG-PTMC diblock copolymer was synthesized by the ROP of trimethylene carbonate (TMC) onto MPEG, using stannous octoate as a catalyst.70 The diblock copolymer solution (≥25 wt.-%) underwent a sol-to-gel transition on an increase in temperature. The phase diagram was mapped by varying the concentration, molecular weight, and composition of the diblock copolymer. On the basis of 13C NMR, DLS, and TEM studies, micellar aggregation, through dehydration of PEG, was attributed to the gelation process. The hydrogel was stable in vitro for up to 90 days, but a 15% weight loss occurred over 20 days in vivo. The degradation in vivo, which is different from the degradation in vitro, may be due to the influence of body fluid. The degradation of PTMC in vivo produced alcohol and carbon dioxide, which did not reduce the pH at the interface between the hydrogel and the tissue.71 ABA-type block copolymers consisting of poly(propylene fumarate) (PPF) and MPEG were synthesized by a simple transesterification method.72 The triblock copolymer with MPEG molecular weights of 570 and 800 in aqueous solution exhibited a thermosensitive gelation process in the concentration range of 5–25 wt.-%. The sol–gel transition was influenced by salt concentration and MPEG molecular weight. Highly unsaturated double bonds in PPF could form in situ crosslinks. Poly(propylene phosphate) (PPP) has been used in biomaterial applications, such as drug delivery, tissue engineering, and gene delivery, because of its biodegradability and biocompatibility.73 PPP aqueous solutions did not exhibit a thermally induced gelation transition, but underwent a sol-to-gel transition in the presence of calcium ions. Polyacetal grafted with MPEG was prepared, to make a thermogelling hydrogel.74 When a 15% graft ratio of MPEG and a 5% graft ratio of poly(orthoester) were introduced onto the polyacetal backbone, the resulting polymer in aqueous solutions (25 wt.-%) underwent a sol-to-gel transition at 34 °C. PEG-poly(ethyl-2-cyanoacrylate) (PEG-PEC) was synthesized by addition polymerization.75 A PEG-PEC aqueous solution (750-450) showed a unique closed-loop phase transition in the concentration range of 4–15 wt.-%. In contrast with the systems described above, such as PCL and PLGA, the sol-to-gel transition temperature of PEG-PEC increased when the polymer concentration increased. Closed-loop gelation resulted from a balance between aggregation and stabilization of micelles as a function of gelation temperature and concentration.
The thermosensitive block copolymer hydrogels have potential applications as biomaterials. However, they suffer from some limitations that restrict the range of applications in which they may be utilized. First, when a thermosensitive polymer solution is injected into the body using a syringe, the increase in temperature to the physiological temperature (37 °C) during injection causes gelation inside the needle, creating a blockage. This makes it difficult to inject thermosensitive polymer solutions into the body. Second, a lack of functional groups limits applications of these materials with respect to delivery of ionic peptides/proteins. Third, it takes a long time to dissolve the thermosensitive polymers in water, thus the polymer should be stored prior to use. The biodegradable polymer could be degraded during storage and circulation for commercial use. Therefore, the reconstitution problem of the polymer solution is of concern. Fourth, the degradation of polyester generates acidic products sometimes that change the local pH. The resulting low pH damages incorporated proteins or cells. Thus, it is important to maintain neutral pH during the degradation.
pH/temperature-sensitive copolymer hydrogels were prepared by combining a pH-sensitive moiety with a temperature-sensitive block to solve the abovementioned drawbacks. Acidic sulfamethazine oligomers (OSMs) were coupled with thermosensitive poly(ε-CL-co-LA)-PEG-poly(ε-CL-co-LA) triblock copolymers to produce pH/temperature-sensitive hydrogels (OSM-PCLA-PEG-PCLA-OSM) (Scheme 9).76 These copolymer hydrogels were synthesized in two steps: First, a carboxylic acid-terminated OSM was obtained by conventional radical polymerization in the presence of a chain transfer agent (3-mercaptopropionic acid), and a PCLA-PEG-PCLA triblock copolymer was produced by the ROP of CL and LA using PEG as a macroinitiator. Second, the carboxylic group in OSM was coupled to the hydroxy groups at both ends of PCLA-PEG-PCLA using 4-(dimethylamino) pyridine (DMAP) as a catalyst. The parent PCLA-PEG-PCLA triblock copolymer aqueous solutions (15 wt.-%) showed a sol-to-gel transition in response to changes in temperature but not in pH (Figure 7a). In contrast, the 15 wt.-% OSM-PCLA-PEG-PCLA-OSM solution exhibited a sol-to-gel transition as a function of both pH and temperature (Figure 7b). The gel window became wider with increasing pH and/or PCLA/PEG ratio. The sol–gel transition could be controlled by the polymer concentration, hydrophobic/hydrophilic balance, PEG block length, and OSM molecular weight.77 An association of bridged micelles was suggested as a mechanism for gelation of the pentablock copolymer. A schematic gelation mechanism is illustrated in Figure 8. At pH 8.0 and in the temperature range of 10–70 °C, the polymer solution existed as a sol state because the OSM was ionized. At pH 7.4 and 15 °C, the OSM deionized and became more hydrophobic, but the pentablock copolymer still exhibited a sol state because of weak interactions with the hydrophilic PCLA block at low temperatures. In contrast, at pH 7.4 and 37 °C, the PCLA blocks became hydrophobic, inducing a strong hydrophobic interaction between PCLA-OSM blocks and leading to a micellar interconnecting gelation process. At pH 8.0, the polymer solution did not form a gel in the temperature range of 10–70 °C, suggesting that the solution was easily injectable using a syringe. A strong gel formed quickly when the polymer solution was injected into a pH 7.4 phosphate buffered saline (PBS) solution, whereas dispersion of the polymer in a pH 8.0 PBS solution was observed. The pentablock copolymer hydrogel maintained its integrity for more than two weeks in pH 7.4 PBS at 37 °C and showed a slower degradation than did the parent PCLA-PEG-PCLA polymer.78 After one month, the molecular weight of the pentablock copolymer decreased from 6 550 to 4 830. The pH drop as a result of degradation of the parent PCLA-PEG-PCLA triblock copolymer was significant, from pH 7.4 to 2.2, whereas the pH drop resulting from OSM-PCLA-PEG-PCLA-OSM degradation was from only pH 7.4 to 5.5 after one month. The buffering effect of OSM moieties minimized the effects of acidic degradation products. A hydrogel formed rapidly after injection of the pentablock copolymer solution (20 wt.-%, pH 8.0) into rats. Good cytotoxicity against HeLa cells was observed in vitro for concentrations up to 10 mg · mL−1 of the pentablock copolymer. A histology study revealed that acute inflammation was found during the first two weeks, but decreased notably after six weeks. Well-defined OSM-PCLA-PEG-PCLA-OSM pentablock copolymers were synthesized by atom transfer radial polymerization (ATRP).79 The Br-PCLA-PEG-PCLA-Br was synthesized by conjugating 2-bromoisobutyryl bromide with PCLA-PEG-PCLA. The pentablock copolymer was then prepared by polymerization of sulfamethazine methacrylate monomers using Br-PCLA-PEG-PCLA-Br as an ATRP macroinitiator. The molecular weight distribution of the polymer was narrow relative to that obtained from conventional radical polymerization. Subsequently, OSM-PCGA-PEG-PCGA-OSM copolymers were synthesized.80 The OSM-PCGA-PEG-PCGA-OSM hydrogel degraded at a faster rate than did the OSM-PCLA-PEG-PCLA-OSM copolymer hydrogel.
Recently, copolymer hydrogels based on a basic poly(β-amino ester) (PAE) were prepared.81, 82 PAE is known to be a pH-sensitive, non-cytotoxic, biodegradable polymer, and the positive charge of PAE facilitated an electrostatic linkage with plasmid DNA (pDNA) at pH 7.2.83 MPEG-PCL-PAE block copolymers were synthesized by the Michael addition polymerization of piperazine, hexan-1,6-diol diacrylate (HDA), and MPEG acrylate.81 Aqueous solutions of the resulting copolymers exhibited a gel-to-sol transition at pH values above 6.0, when the temperature was increased. The phase diagram could be tailored by varying the MPEG molecular weight and PCL block length. Gelation was related to packing of the micelles upon heating. Subsequently, a PAE-PCL-PEG-PCL-PAE pentablock copolymer was prepared by Michael addition polymerization of 4,4-trimethylene dipiperidine (TMDP), PCL-PEG-PCL diacrylate, and butane-1,4-diol diacrylate (BDA) (Scheme 10).82 The parent PCL-PEG-PCL aqueous solution (20 wt.-%) showed a sol–gel transition as a function of temperature but not pH. In contrast, the PAE-PCL-PEG-PCL-PAE aqueous solution at pH values above 6.0 underwent a sol-to-gel transition in response to both temperature and pH changes (Figure 9). When the pentablock copolymer was mixed with insulin, the sol-to-gel transition temperature was lowered because of an ionic complex formed between the polymer and insulin. The PAE-PCL-PEG-PCL-PAE copolymer hydrogel degraded in two steps: first, fast degradation of PAE was noted, followed by slow degradation of the PCL-PEG-PCL triblock copolymer. The PAE in the hydrogel degraded completely within 12 d, whereas 18 d were required for degradation when the hydrogel was mixed with insulin (a complex gel) in a pH 7.4 PBS solution. The sol–gel transition of the polymer/insulin solution shifted relative to the transition of the polymer solution alone.82 Thus, it was important to control this shift near physiological conditions for practical applications.84 The gel window could be tailored by varying the PEG molecular weight, PAE block length, PCL/PEG ratio, and concentration. In addition, the degradation of pentablock copolymers could be controlled by substituting PCLA for the PCL block.85 The PAE of the PAE-PCLA-PEG-PCLA-PAE hydrogel degraded within 10 days compared to the 12 days for the PAE-PCL-PEG-PCL-PAE hydrogel, because of faster degradation of the PCLA block compared to the PCL block.
Subsequently, a series of pH/temperature-sensitive multiblock copolymers based on poly(amino urethane) (PAU) were reported.86 The multiblock copolymers were synthesized by polyaddition of HO-PCL-PEG-PCL-OH, bis-1,4-(hydroxyethyl)piperazine (HEP), and 1,6-diisocynato hexamethylene (HDI) (PCL-PEG-PCL-PAU)n (Scheme 11). At pH values below 7.0, the polymer solution (20 wt.-%) existed as a sol state over a temperature range of 0–80 °C. In contrast, at pH 7.0 and above, the polymer solution exhibited a sol-to-gel-to-aggregation transition upon heating. The sol–gel phase diagram could be controlled by varying the hydrophobic/hydrophilic balance and block length. A hydrogel formed quickly when the polymer solution (20 wt.-%) was subcutaneously injected into rats. The (PCL-PEG-PCL-PAU)n copolymers containing the hydrophobic PCL blocks showed incomplete solubility in water, even at a relatively low pH. A double hydrophilic polymer that did not contain a hydrophobic block dissolved easily in water at low pH. Therefore, a series of pH/temperature-sensitive multiblock copolymers based on the double hydrophilic polymer were synthesized by polyaddition of HO-PEG-OH, HEP, and HDI (PEG-PAU)m.87 A 20 wt.-% aqueous solution of (PEG-PAU)m showed a sol-to-gel-to-sol transition in the pH range of 6.8–7.4 with increasing temperature. Because of complete dissolution in water at low pH, the double hydrophilic multiblock copolymers may be easily mixed with drugs or proteins, suggesting that these copolymers are promising candidates for biomaterial applications.
Unfortunately, it is difficult to control the molecular weight and composition of PAU-related copolymers, because of the presence of bifunctional monomers and the high reactivity of HDI, resulting in multiblock copolymers. A poly(amidoamine)-PEG-poly(amidoamine) (PAA-PEG-PAA) triblock copolymer was synthesized by Michael addition polymerization of PEG, TMDP, and 1,10-decylene diacrylamide (DDA) (Scheme 12).88 The composition of this copolymer was easily controlled, resulting in a triblock copolymer. The PAA block formed a hydrophilic block at relatively low pH (such as pH 3.0), but became hydrophobic at higher pH (such as pH 7.4). The decrease in pKa of the PAA-PEG-PAA polymer with increasing temperature indicated that the PAA blocks became more hydrophobic at higher temperatures. PAA-PEG-PAA aqueous solutions underwent a sol-to-gel-to-condensed gel transition with a very high viscosity of 43.6 kPa · s at 37 °C and pH 7.4. At the condensed gel temperature, water was squeezed from the gel matrix. The gelation process was attributed to the self-assembly, hydrogen bonding, and hydrophobic interactions of PAA blocks. The sol–gel window was tailored by varying the polymer composition and concentration.89 The PAA-PEG-PAA copolymer showed bioadhesive properties, suggesting potential for applications involving mucosal surfaces. The mucoadhesive properties were attributed to hydrogen bond interactions between amide groups and the mucin, and the interactions of positively charged PAA blocks with the negatively charged sialic acid of mucin. In contrast to PAA-PEG-PAA copolymers, which displayed a sol-to-gel transition with increasing temperature, concentrated PAE-PEG-PAE triblock copolymers (30 wt.-%) in water showed a gel-to-sol transition at higher temperatures and at pH > 6.4.90 PAE-PEG-PAE was synthesized by Michael addition polymerization of PEG diacrylate, TMDP, and HDA. The resulting polymer could be easily dissolved in water at a relatively low pH, because of its doubly hydrophilic character. The sol–gel transition point was adjusted by changing the composition. This polymer also showed potential for bioadhesive applications.
Biomedical Applications of Injectable Biodegradable Hydrogels
Biodegradable polymeric hydrogels that are responsive to temperature or pH/temperature changes have been widely used for biomedical applications, such as drug/protein delivery and tissue engineering. Aqueous polymer solutions may be loaded with drugs, proteins, or cells at a specific temperature before injection into particular sites in the body. Once formed, the hydrogels act as drug delivery matrices or cell growth depots. Table 1 summarizes the biomedical applications of the biodegradable injectable hydrogels.
Table 1. Various biodegradable injectable hydrogel systems and their applications.
Thermogelling PEG-PLLA-PEG hydrogels were investigated for the sustained release of fluorescein isothiocyanate (FITC)-labeled dextran, which was incorporated into polymer solutions at 45 °C, and the drug-loaded polymer solutions were changed to gels by cooling to body temperature. The release rate of dextran was influenced by polymer concentration. With a 35 wt.-% polymer gel, dextran was released at a constant rate over the course of 12 days, without a burst release, in contrast with the burst effect observed for 23 wt.-% polymer gels.9
To study the influence of hydrophobicity and hydrophilicity on sustained release using a PEG-PLGA-PEG triblock copolymer hydrogel, spironolactone and ketoprofen were employed as drug models.91 The hydrophilic ketoprofen was released over two weeks with a first-order release profile, whereas hydrophobic spironolactone required two months for release and yielded an S-shaped release trace. A diffusion mechanism was proposed for release of the hydrophilic drug, whereas diffusion at the first stage, followed by degradation at later stages, was the mechanism envisaged to describe the release of the hydrophobic drug. pDNA was released from a PEG-PLGA-PEG matrix at an approximately constant rate (zero order) over two weeks, and the drug half life was five days.92 When a hydrogel mixed with luciferase pDNA was applied to skin wounds of CD-1 mice, the expression of luciferase reached a maximum at 24 h and then dropped to 94% at 72 h. The supercoiled structure of the released pDNA was preserved, although a small quantity of linear pDNA was observed. A plasmid TGF-β1-loaded PEG-PLGA-PEG hydrogel was used to accelerate wound healing in diabetic mice.93 Reepithelialization was complete at day 9 using hydrogels containing TGF-β1, at day 11 employing a commercial wound dressing (Humatrix) that contained TGF-β1, and at day 14 in the absence of treatment. Fibroblast proliferation and collagen organization were significantly enhanced by hydrogel treatment relative to treatment with TGF-β1-loaded Humatrix. These results suggest that PEG-PLGA-PEG hydrogels could be used as a gene delivery system for treatment of skin disorders and wound healing. A hydrogel mixed with FITC was found to prolong drug exposure in the bladder of rats.94 The FITC-loaded hydrogel showed sustained release for up to 24 h after instillation, whereas all administered free FITC was observed in the urine of rats after 8 h.
The in vitro release of human insulin from PLGA-PEG-PLGA (ReGel) systems proceeded at a constant rate over two weeks, without an initial burst effect.95 The presence of 0.2% (w/v) zinc could enhance the insulin release rate, and almost 90% of the insulin was released over the course of two weeks. An in vivo study revealed that a steady amount of insulin was secreted from the ReGel/0.2 wt.-% Zn-insulin depot within a period of two weeks. A related study showed that the incretin hormone glucagon-like peptide-1 (GLP-1) was released from ReGel in vitro and in vivo.96 The in vitro study of GLP-1/ReGel showed a release profile over five days, whereas the release profile of zinc-complexed GLP-1/ReGel (ZnGLP-1/ReGel) exhibited a zero-order release profile over two weeks, without an initial burst. After a single injection of ZnGLP-1/ReGel into Zucker diabetic fatty (ZDF) rats, plasma GLP-1 was maintained at levels that were significantly higher than the control group over the course of two weeks. In addition, plasma insulin levels increased and the blood glucose levels fell.
Testosterone was released over the course of three months from PLGA-PEG-PLGA systems.97 The drug release rates were affected by drug concentration, solvent composition, and composition of the copolymer. The release of indomethacin and 5-fluorouracil from PLGA-PEG-PLGA systems was also reported.98 Hydrophilic 5-fluorouracil and hydrophobic indomethacin were secreted from the hydrogel over five days and one month, respectively, and the release rate depended on copolymer composition. Ganciclovir-loaded PLGA microspheres were dispersed in a PLGA-PEG-PLGA gel.99 The microspheres/hydrogel showed a slower release rate compared to that from PLGA microspheres alone. The release of levonorgestrel, interleukin-2, and ceftazidime from a PLGA-PEG-PLGA hydrogel was reported.100–102 ReGel was used to release a hydrophobic drug, paclitaxel (ReGel/paclitaxel: OncoGel),23 over the course of 50 d. The in vitro release study showed diffusion-controlled release within the first two weeks, followed by combined diffusion/polymer degradation.
A PLGA-g-PEG/PEG-g-PLGA hydrogel was used for the sustained release of insulin.103 After a single injection of insulin-loaded hydrogel, blood glucose levels could be adjusted from 5 to 16 days in diabetic rats, depending on polymer composition. Also, a chondrocyte-loaded PLGA-g-PEG gel was used to repair an articular cartilage defect. Poly(N-isopropyl acrylamide)-co-acrylic acid/hydroxyapatite collagen sponges containing chondrocytes were used as a control. The cartilage defect was completely repaired using the PLGA-g-PEG hydrogel, in contrast with persistence of control defects. The superior efficacy of cartilage defect repair was attributed to the biodegradability of the PLGA-g-PEG.
The in vivo osteogenic differentiation of rat bone marrow stromal cells (rBMSC) was studied.104 Histological analysis demonstrated that in situ formed MPEG-PCL hydrogels containing rBMSC and dexamethasone were biocompatible and enhanced bone formation. The in vitro release of FITC-labeled bovine serum albumin (FITC-BSA) from the MPEG-PCL hydrogel showed a sustained release profile for more than 20 days.105 Although in vivo release of FITC-BSA was sustained for 30 days, an initial burst was observed. In addition, MPEG-PCL wafers could be prepared as an implantable material using a direct compression method.106 BSA was released from the wafers over the course of 30 days with an initial burst release.
L929 cells encapsulated in cholesterol end-capped star PEG-PLLA hydrogels (10 and 20 wt.-%) were viable and proliferated in three dimensions within the hydrogels,46 indicating that the polymer could be used as an injectable cellular scaffold. The PEG/PPG/PHB hydrogel is a promising candidate for the controlled release of protein.48 The in vitro release of BSA was controllable over 70 d. Diffusion control was proposed as the release mechanism during the first stage of release, and the mechanism at the later stages was proposed to be erosion control.
Polyphosphazenes were used for the sustained release of FITC-dextran and human serum albumin (HSA) over the course of two weeks.107 The release of FITC-dextran was dependent on polymer concentration. The solubility of doxorubicin (DOX) was enhanced in polyphosphazene hydrogels.108 DOX was released at a controlled rate from the polyphosphazene hydrogel over 20 d, and the release rate was affected by gel strength. The antitumor activity of DOX in the mouse lymphoblast cell line P388D1 was maintained over the course of one month. The release of FITC-albumin from polyphosphazene hydrogels was controlled using chitosan,109 and was sustained over two months without an observed initial burst in the presence of chitosan, in contrast with the observed release over one month without chitosan. The extension of release time was attributable to formation of an ionic complex between chitosan and FITC-albumin. The polyphosphazene hydrogel has also been used to entrap pancreatic islets.110 In comparison with both rat islets entrapped in other hydrogels, and free islets, rat islets in the polyphosphazene hydrogel retained higher cell viability and insulin production over a 28-day culture period. In subsequent work, polyphosphazene hydrogels were used to encapsulate hepatocytes as spheroids or single cells.111 Over a 28-day culture period, the spheroid hepatocytes maintained a higher viability and produced albumin and urea, whereas single hepatocytes reduced the level of albumin secretion from the hydrogel. These results suggest that the hydrogels could be used as a three-dimensional cell culture system. More recently, paclitaxel within covalently conjugated polyphosphazene gels showed sustained release over one month at pH 7.4, and three days at pH 6.8.112 The in vitro evaluation of antitumor activities against several cancer cell lines indicated that paclitaxel was released from the hydrogel without inhibiting tumor growth. However, the paclitaxel-conjugated polyphosphazene hydrogel showed a higher level of antitumor activity over one month compared with the control, after a single intratumoral injection of the hydrogel into HSC-45M2 human gastric cancer cell-containing nude mice.
C/GP was used as a hydrogel for the sustained release of camptothecin.113 The release profile showed zero-order release kinetics within the first four weeks. In contrast with the blank C/CP, tumor growth was significantly delayed after injection of C/GP-containing camptothecin. Insulin was mixed with G/GP and released over 350 h. At a fixed concentration of C (2.5 v/w), the higher GP content resulted in a faster release of insulin, due to the higher mobility of the protein in the gel. The diffusion controlled process was attributed to the release mechanism.114 A chitosan-g-pluronic (CP) system was used for injectable cell delivery to generate cartilage.115 When bovine chondrocytes were encapsulated in a CP hydrogel, cell viability and synthesized glycosaminoglycan content increased after 28 days of cell culture.
OSM-PCLA-PEG-PCLA-OSM hydrogels with good biocompatibility were used for the controlled release of paclitaxel.116 An in vitro release (pH 7.4, 37 °C) study showed sustained release at a constant rate over 20 d. An in vivo study was carried out on C57BL/6 male mice. Good anti-tumor activity was observed over the course of two weeks after subcutaneous injection of the paclitaxel-loaded copolymer hydrogel into tumor-bearing mice. After two weeks of treatment, the tumor volume of saline-treated mice was about 17 cm3, whereas that of paclitaxel/hydrogel-treated mice smaller than 7 cm3. The anti-tumor effect depended on the concentration of paclitaxel. At the paclitaxel concentration of 1 mg · mL−1, the body weight was constantly maintained for two weeks while the body weight decreased during the first six days and slightly increased. The release of paclitaxel from OSM-based hydrogels was controlled using PLGA-PEG-PLGA instead of PCLA-PEG-PCLA.77 In comparison with the release profile of paclitaxel from the OSM-PCLA-PEG-PCLA-OSM hydrogel, the release of paclitaxel from the OSM-PCGA-PEG-PCGA-OSM hydrogel was faster, because of more rapid degradation of the PCGA block. A OSM-PCLA-PEG-PCLA-OSM solution containing human mesenchymal stem cells (hMSCs) and recombinant human bone morphogenetic protein-2 (rhBMP-2) was injected into the backs of mice.117 After seven weeks, mineralized tissue with high levels of alkaline phosphate activity had formed, indicating that this material could be used as an injectable scaffold for bone tissue engineering.
pH/temperature-sensitive PAE-PCL-PEG-PCL-PAE hydrogels were employed for the controlled release of insulin.82 The polymer exhibited as a solution at pH < 6, thus, it was easy to mix the polymer and insulin at pH 3.0–4.0 at 2 °C. An in vitro study carried out at pH 7.4 and 37 °C showed a constant release rate profile for up to 20 days. After a single injection of the PAE-PCL-PEG-PCL-PAE/insulin solution into SD rats, the serum insulin concentration was sustained for one month and plasma insulin levels were maintained at a constant level over 15 days, without an initial burst (Figure 10). In contrast, a 1-day duration of plasma insulin level and a marked initial burst were observed when the PCL-PEG-PCL/insulin mixture was employed. The physiological effects of pentablock copolymer/insulin injection were investigated in diabetic rats.118 After a single injection into a diabetic rat, steady blood glucose levels were maintained for more than one week, without a decrease in body weight. The blood glucose levels and body weight showed a dose dependence of insulin loaded into the polymer. The sustained release of insulin was related to the ionic interactions between partial positive charges in PAE blocks and negative charges in insulin at physiological pH. The sustained release of insulin was attributed to degradation of the PAE blocks, suggesting that a degradation-controlled process was the major mechanism of insulin release.
The (PCL-PEG-PCL-PAU)n multiblock hydrogel was used for sustained release of paclitaxel over the course of one month, and chlorambucil was released from the (PEG-PAU)m multiblock hydrogel for 10 days under physiological conditions. For these polymers, drug encapsulation was easily performed at pH 5.0–6.0 and 2 °C because of their solubility character.86, 87 PAE-PEG-PAE and PAA-PEG-PAA hydrogels may have potential for drug delivery at mucosal surfaces.89, 90 Lidocaine and flurbiprofen were released from PAE-PEG-PAE and PAA-PEG-PAA hydrogels over the course of one day, indicating that these hydrogels could be used as drug carriers.
Conclusions and Perspectives
Significant progress has been achieved in the development of injectable biodegradable polymeric hydrogels, and each hydrogel system has special intrinsic properties (including gel strength, pH after degradation, and degradation rate) that may be appropriate for a particular application. In this review, the characteristics, sol–gel mechanisms, and biomedical applications (such as drug/cell delivery and tissue engineering) of temperature- and pH/temperature-sensitive hydrogels are summarized. Some challenges remain for improving the applicability of hydrogels.
The first challenge in the development of practical injectable applications of thermosensitive block copolymer hydrogels is administration. The risk of the syringe clogging upon injection can be addressed by modulating the gelation temperature and lowering the polymer concentration. Slowly degrading polymers can be used for long-term applications in drug/cell delivery and tissue engineering, thus, the lifetime of the gel prior to degradation must be considered. For controlled drug release, an initial burst release is a limiting factor, especially for low-molecular-weight drugs, hydrophilic drugs, and proteins. The requirement of storage in solution for commercial use may cause degradation, thus, the reconstitution problem should be considered. The products generated from the degradation of some polymers, such as PLLA and PLGA, may lower the pH of the surrounding environment and should, therefore, be considered. Such degradation products may cause inflammation at the injection site or damage incorporated proteins/cells.
pH/temperature-sensitive block copolymer hydrogels show several advantages over thermosensitive block copolymer hydrogels, such as the absence of clogging during injection, which allows facile injection into deep sites in the body, avoidance of the local low pH environment caused by degradation, which protects proteins/cells from damage, and the ease of handling and storage. In particular, cationic block copolymer hydrogels can form ionic linkages with the anionic proteins or pDNA at physiological conditions, which result in sustained release without an initial burst. The sustained release of the low-molecular-weight and hydrophilic drugs may be obtained by ionically bonding with the polymers. The pentablock copolymers possess complicated structures that may make FDA approval difficult, thus, well-defined pH/temperature-sensitive copolymers should be considered. In addition, the degradation rate of such well-defined copolymers should be optimal for the sustained release of proteins over one month.
For both kinds of the above mentioned block copolymer hydrogels, appropriate time for clearance of the hydrogels from the body is quite important. The degradation depends on polymer composition, crystallinity of the polymer, and topological structure. This understanding may help to design hydrogels with fine tuning of the degradation rate, which results in the corresponding release rate. The changes of the gel phase and gel strength after mixing with drugs, cells, or proteins should be a challenge. The integration of drugs may be considered for potential applications. Biodistribution, elimination routes, and the effect of degradation products on organs, are also of concern. For tissue engineering and cell growth, cell adhesion to the gels, molecular-level characteristics of the tissue/polymer interface, effects of the degradation products on cells, and the effect of hydrogels on histogenesis, should be further investigated.
The authors thank Dr GuangJinIm for his comments on this paper. This research was financially supported by the MEST and NRF (20090093631)
Minh Khanh Nguyen graduated from the Department of Chemical Engineering from HoChiMinh City University of Technology (VietNam) in 2003 and is currently pursuing his Ph.D. under the guidance of Professor Doo Sung Lee at Sungkyunkwan University (Korea). His main research is focused on the development of functionalized and biodegradable injectable polymeric hydrogels for controlled drug and protein delivery.
Doo Sung Lee studied chemical engineering at the Seoul National University. He completed his M.Sc. and Ph.D. at Korea Advanced Institute of Science and Technology in 1984 working in the field of interpenetrating polymer networks. He joined Sungkyunkwan University as Assistant Professor in the Department of Textile Engineering. He was one of the founders of the Department of Polymer Science and Engineering and was appointed Professor of this department in 1993. He served as a Dean of the College of Engineering at Sungkyunkwan University in 2005–2007. He started to study biomaterials in 1995 when he was in the Department of Pharmaceutics and Pharmaceutical Chemistry at the University of Utah as a visiting Professor. He served as an Editor-in-chief of PolymerScienceandTechnology (2001) and on the Editorial board of MacromolecularResearch for four years (2002–2005). He was elected as a member of Korean Academy of Engineering in 2007. He is currently a vice president and an editorial board member of BiomaterialsResearch of the Korean Society of Biomaterals (2006–present). He has authored and co-authored about 300 papers, including about 120 in peer-reviewed journals, and seven book chapters, and filed 26 patents. His main research interest is functionalized and biodegradable injectable hydrogels and micelles for the controlled drug and protein delivery and molecular imaging.