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Keywords:

  • drug delivery;
  • in situ forming;
  • injectable hydrogels;
  • temperature-responsive;
  • tissue engineering

Abstract

  1. Top of page
  2. Abstract
  3. INTRODUCTION
  4. INJECTABLE HYDROGELS FOR DRUG DELIVERY
  5. INJECTABLE HYDROGELS FOR TE AND REGENERATIVE MEDICINE
  6. INJECTABLE HYDROGELS FOR SPACE-FILLING APPLICATIONS
  7. CONCLUSIONS
  8. REFERENCES AND NOTES
  9. Biographical Information
  10. Biographical Information
  11. Biographical Information
  12. Biographical Information

Hydrogels are promising for a variety of medical applications due to their high water content and mechanical similarity to natural tissues. When made injectable, hydrogels can reduce the invasiveness of application, which in turn reduces surgical and recovery costs. Key schemes used to make hydrogels injectable include in situ formation due to physical and/or chemical cross-linking. Advances in polymer science have provided new injectable hydrogels for applications in drug delivery and tissue engineering. A number of these injectable hydrogel systems have reached the clinic and impact the health care of many patients. However, a significant remaining challenge is translating the ever-growing family of injectable hydrogels developed in laboratories around the world to the clinic. © 2012 Wiley Periodicals, Inc. J Polym Sci Part B: Polym Phys, 2012


INTRODUCTION

  1. Top of page
  2. Abstract
  3. INTRODUCTION
  4. INJECTABLE HYDROGELS FOR DRUG DELIVERY
  5. INJECTABLE HYDROGELS FOR TE AND REGENERATIVE MEDICINE
  6. INJECTABLE HYDROGELS FOR SPACE-FILLING APPLICATIONS
  7. CONCLUSIONS
  8. REFERENCES AND NOTES
  9. Biographical Information
  10. Biographical Information
  11. Biographical Information
  12. Biographical Information

Biomaterials encompass a large and diverse array of materials that range from metallic orthopedic implants to polymeric constructs aimed at replacing, restoring, or regenerating lost tissue structure and/or function. An ever-growing class of biomaterials is polymeric hydrogels, classically defined as three-dimensional (3D), water-swollen polymer networks formed as a result of physical or chemical cross-linking.1 Their thermodynamically driven interactions with aqueous media cause equilibrium swelling/shrinking behavior that result in important properties in drug delivery and tissue engineering (TE) applications.2 Because of their high water content and mechanical resemblance to natural tissues, hydrogels show promising biocompatibility and potential for medical/biological applications.1 Injectable hydrogel formulations are especially attractive due to their minimally invasive delivery procedure, providing reduced healing time, reduced scarring, decreased risk of infection, and ease of delivery compared with surgically implanted materials.3 Injectable hydrogels are especially useful for applications where the final form and shape are either not important or are defined by the void or space into which they are injected.

There are a variety of schemes used to make hydrogel systems injectable. Table 1 summarizes these schemes and provides a number of advantages and disadvantages for examples. The most common schemes include materials that form in situ on injection due to physical and/or chemical cross-linking. Physical cross-linking occurs in some injectable hydrogels in response to an environmental condition such as temperature, pH, or ionic strength. Figure 1 shows an example of this mechanism with the gelation of a N-isopropylacrylamide copolymer in response to a change in temperature. Chemical cross-linking between soluble precursor materials in situ can be achieved through a variety of chemical processes including enzymatic, photoirradiation, and self-reactive reactions. Figure 2 shows concept of chemical cross-linking for injectable hydrogels. Hydrogels can also be rendered injectable by preforming the gels into microparticles or nanoparticles. However, particulate systems (e.g., micelles, liposomes, polymer-drug conjugates, microparticles, and nanoparticles) that are injectable by virtue of their small size constitute a vast field of research17–21 beyond the scope of this review. To this end, the objectives of this review are to provide the reader with an overview of injectable hydrogel systems, describe examples of their use in drug delivery, TE/regenerative medicine, and space-filling applications, and use these examples to expound on the mechanisms used for in situ gelation.

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Figure 1. Physical gelation via stimuli-responsive polymer is one mechanism used to create injectable hydrogels. This example shows a temperature-responsive N-isopropylacrylamide copolymer undergoing a phase transition on change of temperature.

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Table 1. Examples, Advantages, and Disadvantages of the Different Cross-Linking Mechanisms Used to make Hydrogels Injectable
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INJECTABLE HYDROGELS FOR DRUG DELIVERY

  1. Top of page
  2. Abstract
  3. INTRODUCTION
  4. INJECTABLE HYDROGELS FOR DRUG DELIVERY
  5. INJECTABLE HYDROGELS FOR TE AND REGENERATIVE MEDICINE
  6. INJECTABLE HYDROGELS FOR SPACE-FILLING APPLICATIONS
  7. CONCLUSIONS
  8. REFERENCES AND NOTES
  9. Biographical Information
  10. Biographical Information
  11. Biographical Information
  12. Biographical Information

Injectable hydrogels are often engineered to function as carriers that provide controlled drug release.22–26 These materials are particularly useful for local drug delivery, providing greater drug concentration at the desired site of action while minimizing the systemic drug concentration and the associated site effects.27 Injectable hydrogels can also be easily delivered into sites that are difficult or infeasible to access in a surgery. These formulations are typically designed to provide prolonged activity from a single administration, reducing risks associated with patient compliance. Because the size and shape of an injectable hydrogel is determined by its local environment,28 good interfacial contact is usually achieved between the hydrogel and the nearby tissue,29 allowing for drug to elute directly into the local tissue and, in some cases, facilitating healing. Further, injectable and degradable materials allow for easy delivery of additional doses to the same site. Figure 3 shows concept of injectable, degradable polymers for drug delivery.

Some drawbacks exist for using injectable hydrogels for drug delivery. Most hydrogels are too weak to be load-bearing, and some (physical gels, in particular) are prone to plastic deformation in response to stress due to their viscoelastic character. Insoluble drugs can become heterogeneously distributed within the gels, leading to sample-to-sample variability of drug release rates. The greatest drawback is typically that drug release from hydrogels suffers from either a rapid initial burst release of drug followed by sustained release, or rapid release of drug altogether.23 Burst release is a primary concern for injectable formulations because slow gelation in situ can lead to loss of drug before the gel fully forms. Changes in gel volume after gelation can also cause rapid release.

Although the ideal properties of the hydrogel depend heavily on the application for which it is intended, there is a basic set of properties that are advantageous for drug delivery.30, 31 Desirable hydrogel properties for drug delivery include:

  • i
    Low/medium viscosity of the solution before injection.
  • ii
    Fast gelation that minimizes the initial burst release of drug.
  • iii
    High water content and good biocompatibility.
  • iv
    Compatibility with a wide variety of drugs (low molecular weight, proteins, nucleic acids).
  • v
    Efficient drug loading.
  • vi
    Maintenance of volume and interfacial contact with tissue after gelation.
  • vii
    Control of drug release rate over a range of time frames.
  • viii
    Long (>3 months) shelf life.
  • ix
    Easy to purify and sterilize.
  • x
    Degradability into low molecular weight, soluble byproducts either by hydrolysis or enzymatic degradation.
  • xi
    Suitable drug release profile (i.e., cumulative release vs. time) for the application.

In this section, injectable drug delivering matrices reported in the literature will be presented with respect to these advantages according to the mechanism for gelation (temperature-responsive, in situ cross-linking, ionic interaction, and self-assembly) that allows for injectability.

Temperature-Responsive Hydrogels

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.

N-Isopropylacrylamide-Based Hydrogels

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.

Poly(organophosphazene)s

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

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

In Situ Cross-Linking Hydrogels

Hydrogels can be formed from liquid or soluble precursors by in situ chemical reaction by a number of mechanisms including thermal polymerization, photopolymerization, or in situ polymer–polymer cross-linking reactions. For any application, it is desirable for such reactions to take place under biologically compatible conditions and usually with minimal side reactions to nearby tissue. In situ cross-linking is ideal for incorporating a wide variety of biocompatible hydrophilic biomaterials (such as natural soluble polymers or PEG) into an injectable hydrogel. The kinetics of gelation is an important consideration for drug delivery, particularly for systems that do not gel beginning on the outside, forming an initial “skin” and then proceed to gel toward the center of the material. The material must have low enough viscosity to be injectable and yet gel quickly enough after injection to avoid burst release. Efficient cross-linking is also important for minimizing the toxicity associated with reactive chemical species or leachable small molecules (e.g., monomers, cross-linkers). Discussion of these systems will be divided into natural or hybrid materials that are comprised of at least one natural component, and materials made from fully synthetic precursors.

Natural or Hybrid In Situ Cross-Linking Hydrogels

Natural biomaterials are often high molecular weight, linear polymers that can subsequently be cross-linked together through a controlled fraction of functional groups on their sides, most commonly hydroxyls or amines. Alginate oxidized to contain terminal aldehyde groups can cross-link through proteins such as gelatin, which is accelerated in the presence of borax.105 The gelation time of these materials was 20–50 s, and even when the materials were allowed to gel for 10 min before release was measured in vitro, a 30% or greater burst release of primaquine was observed within 6 h for all gels tested. The remainder of the drug was released slowly over at least 5 days. Similar chemistry has been used with an aldehyde-functionalized alginate derivative cross-linking through adipic acid dihydrazide, cross-linking on the formation of two hydrazone linkages.10 A disadvantage of aldehyde chemistry is its lack of specificity—aldehydes are prone to react with amines, which are present in proteins, as is the case for the fixatives glutaraldehyde and formaldehyde. Conversely, a chemistry that has not been evaluated for drug delivery but is capable of forming cross-linked networks with high specificity is high-affinity noncovalent binding, such as that between avidin and biotin.106

A common cross-linking chemistry in both synthetic and natural materials is the Michael addition reaction between thiols and acrylates. Cai et al.11 developed thiol-modified hyaluronic acid and chondroitin sulfate materials for in situ cross-linking with PEG diacrylate (PEGDA) and release of basic fibroblast growth factor (bFGF). Gelation time was dependent on pH, with gels forming more rapidly at higher pH. Gels formed within 1 min at pH 8.5. Release of FITC-HSA was prolonged over about 5 days in vitro. Using covalently bound heparin to complex with the bFGF, its release was prolonged over about 28 days, and this formulation showed improved neovascularization relative to gels without heparin or bFGF alone. Hiemstra et al.107 used thiol-modified dextran with either four-arm PEG-tetraacrylate or dextran functionalized with vinyl sulfone to form gels in under 1 min. At 20 wt %, these gels became rather strong, with storage modulus near 100 kPa. The gel properties of these materials are particularly sensitive to the ratio of reactive groups, with the strongest gels obtained using a thiol:acrylate ratio near 1:1.

Enzyme-mediated in situ cross-linking can be done using tyramine-functionalized materials. Jin et al.108 functionalized hydroxyl groups on dextran with pendant phenol moieties, which are coupled to each other in the presence of hydrogen peroxide and horseradish peroxidase. Similar systems based on hyaluronic acid showed fast burst release in the first 5 h up to a fraction of the total protein, which was related to the cross-link density of the gels.109 This period was followed by degradation-based release of the entrapped model proteins lysozyme and α-amylase only in the presence of hyaluronidase.

Cross-linking can also be initiated using light as an external energy source. For example, chitosan functionalized with azide and lactose groups was cross-linked by UV light with a gelation time under 30 s and used for release of FGF-2 in a wound dressing.110–112 The reaction proceeds by release of nitrogen gas from azide groups, converting them to reactive nitrene groups which then cross-link either with amines or other nitrenes to yield a stable azo linkage. These materials provided some release for 1 day after gelation and then nearly zero release thereafter, although protein release was measured to be incomplete.111 It is unclear what fraction of the FGF-2 remained active following the UV exposure, but improved wound closure with time was observed, indicating that some active protein was released.112 These materials released the alkaline model drug toluidine blue completely within 1 day in vitro, whereas almost no acidic model drug (trypan blue) was released for 5 days. This difference was attributed to the alkaline character of the chitosan itself, demonstrating the potential for control over release rate depending on the affinity between the drug and the device.

Synthetic In Situ Cross-Linking Hydrogels

In situ polymerization of monomers with macromolecular cross-linkers can be used for controlled drug delivery applications provided that gelation is fast and the monomers have low toxicity. For example, West and Hubbell113 developed gels based on PEGDA and N-vinyl pyrrolidone (NVP), which cross-link by radical polymerization into solid gels within 20 s of exposure to UV light. The cross-linker in these designs can be engineered to control the degradability of the gels due to either hydrolysis113 or the action of specific enzymes.114 These gels were shown to release a variety of proteins (6–66 kDa) in vitro at a steady rate over a period of up to 5 days depending on molecular weight. The UV photopolymerization was also shown to not substantially affect the activity of entrapped tissue plasminogen activator (tPA). In situ photopolymerized gels using poly(ε-caprolactone fumarate) diacrylate as the macromer required over 3 min to form gels when exposed to blue light.12 Thin cylindrical shaped gels (8 mm diameter, 2 mm thickness) optimized for maximum cross-linking showed diffusion-based release of the low molecular weight, hydrophobic drug tamoxifen citrate over a period of about 5 days in vitro. Although the cross-linked gels are relatively nontoxic to cells, the macromers themselves can be relatively cytotoxic. Studies on acrylate-functionalized macromers115 and initiators116 show concentration-dependent toxicity of these components of in situ polymerizing hydrogels.

Some systems rely on cross-linking reactions between individual low-molecular-weight precursors to form gels rather than a polymerization reaction. This approach may be advantageous in that it does not require an external energy source or give rise to radicals. The most common among these in situ cross-linking systems are systems of thiols and acrylates, of which the average functionality of the precursors must be greater than 2 for network formation (i.e., gelation) to occur.117, 118 Elbert et al.119 mixed PEG-dithiol and multiarm PEG acrylates at 40 wt % to yield gels which released albumin over 8–12 days, with slower release and longer degradation time observed in gels with greater average functionality. Vernon et al.120 used PEGDA and the tetra-thiol precursor pentaerythritol tetrakis(3-mercaptopropionate) (QT) at a very high concentration (75 wt %) for partition-controlled release of progesterone. Gels with high drug loading showed prolonged and steady release over at least 50 days.

Drug release data from studies using in situ cross-linking gels must be evaluated carefully because the data is likely to underestimate the amount of burst release that would occur from the gels if used in vivo. Because the standard method in the literature involves curing the liquid precursors completely before exposing the gels to the release medium, most measurements of drug release behavior provided by these materials represent a “best case” in terms of minimizing burst release. Although studies on temperature-responsive hydrogels may show similar bias by pregelling a drug-loaded injectable solution before exposing the gel to the release medium, burst release from temperature-responsive hydrogels tends to occur more slowly on syneresis after gelation, and is therefore still likely to be reflected more accurately in vitro.

Hydrogels Forming by Ionic Interactions or Self-Assembly

Systems that undergo gelation in the presence of ions in the solvent have also been evaluated.121 A change in the concentration of specific ions in solution can trigger solution to gel (sol-gel) transition in some materials due to ion exchange that causes a change in the interactions between polymer molecules and the solvent. For example, alginate, a naturally derived polymer, undergoes gelation in the presence of Ca2+ and other divalent ions.1 This material was shown to release active VEGF for over 2 weeks when first allowed to gel for 30 min122, 123 and show improved motor function in a rat model of Huntington's disease relative to a control without VEGF.124 Although alginates can entrap a wide variety of drugs, they have a long in vivo degradation time.125 Alginate gels also have a time-sensitive gelling process which poses a challenge for drug delivery—too fast, and the material is not injectable; too slow, and the drug is released rapidly before gelation can occur. Alginate gels undergo faster gelation without a preliminary mixing step when used in ophthalmic applications because the higher calcium concentration in the eye allows for alginate gelation without an initial mixing step. In these applications, release of low-molecular-weight drugs was prolonged over less than 12 h.126, 127

Electrostatic interactions between oppositely charged materials yield hydrogels following a premixing step. Van Tomme et al.128 reported formation of mostly elastic gels on mixing of oppositely charged microspheres comprised of dextran cross-linked either with negatively charged methacrylic acid or positively charged DMAEMA. These gels showed degradation-based release of rhodamine-B-dextran (70 kDa), which can be controlled by adjusting the cross-link density, microsphere weight fraction, and charge balance.129 Negatively charged chondroitin 6-sulfate and positively charged type A gelatin were used to form a gel by complex coacervation, which was used to control the gelation and drug release from a methylcellulose (MC) hydrogel.130 However, this system involves a salt (ammonium sulfate) that is unlikely to remain in the gel for several days after injection in vivo, limiting its use to short-term formulations.

Hydrogels can form due to noncovalent interactions between similar molecules131–134 or complexation of two mixed precursors.135, 136 Such gels are often referred to as self-assembling or nanoassembling gels. Self-assembling hydrogels are shear thinning that allows for injectability after subjecting the material to stress and then gels on subsequent recovery of its mechanical properties in situ. Injectable self-assembling peptide nanofibers comprised of short peptides were shown to release PDGF-BB in vivo in a myocardial infarct model for up to 14 days.131 These peptides contain an alternating polar/nonpolar pattern of amino acids which allows them to form nanofibers or other nanostructures which form gels on entanglement of those structures.137 This design is advantageous because of the low molecular weight of the gelator, the low concentration required for gelation, and the enzyme-degradable nature of the materials. Using a similar approach, amphiphilic prodrugs based on acetominophen conjugated to fatty acids were shown to deliver drugs without initial burst release.133 Alternatively, a non-drug gelator such as ascorbyl palmitate can be used to entrap drugs.134 Ascorbyl palmitate forms a gel, and drugs can be physically entrapped within the gel. Drug release was enzyme-mediated, with almost zero release of physically entrapped drugs in the absence of enzyme.133, 134 Drug is released in response to up-regulated enzymes, including matrix metalloproteinase (MMP)-2, MMP-9, and esterases such as lipase, as the ester of ascorbyl palmitate is degraded. Ascorbyl palmitate is an alkyl-modified ester of vitamin C. Such low release in the “off” state over as long as 7 days is remarkable for a hydrogel material. This design is particularly promising because it uses materials that are natural and generally regarded as safe, offers flexibility for the incorporation of multiple types of drugs (hydrophilic drugs in the gelator itself and hydrophobic drugs entrapped physically), and allows for very high drug loading since a significant fraction of the device itself is composed of drug. However, because the degradation and release are dependent on enzyme action, the rate of release is likely to be governed mostly by the amount of enzyme activity in the injection site. As injection of any foreign material is likely to induce some acute inflammatory response, proteases involved in this process might trigger faster release of drug. The release of the hydrophilic compound in such a case could be controlled through changing the gelator concentration. Another concern with these materials is their long gelation time, which may produce burst release in vivo even if not observed after the 15–45 min gelation time. Over this amount of time, it is reasonable to expect that some release might occur as the gel recovers its properties.

Synthetic polymers can also be used in injectable self-assembling hydrogels. One synthetic approach using segments of PEG-poly(hydroxybutyrate)-PEG block copolymer hydrogels with α-cyclodextrin has been used to cause gelation, compared with the block copolymer alone which is soluble.138, 139 Cyclodextrins form a necklace-like molecular structure with linear polymers such as PEG,140, 141 which contribute to the crystallinity of the polymer. The gels are shear-thinning and thixotropic, reducing in viscosity by about five-fold after 10 min of agitation and then recovering over a period of 12 h. The gels released entrapped FITC-dextran (20 kDa) over 25 days at a nearly constant rate after gelation. However, the same concern as above applies—the gels were allowed to thicken for 12 h before the start of the release study. Also the viscous and weak (∼100 Pa*s viscosity at low shear rate)139 rheological properties of these gels may lead to irreversible deformation under nearly any load.

INJECTABLE HYDROGELS FOR TE AND REGENERATIVE MEDICINE

  1. Top of page
  2. Abstract
  3. INTRODUCTION
  4. INJECTABLE HYDROGELS FOR DRUG DELIVERY
  5. INJECTABLE HYDROGELS FOR TE AND REGENERATIVE MEDICINE
  6. INJECTABLE HYDROGELS FOR SPACE-FILLING APPLICATIONS
  7. CONCLUSIONS
  8. REFERENCES AND NOTES
  9. Biographical Information
  10. Biographical Information
  11. Biographical Information
  12. Biographical Information

TE and regenerative medicine are highly interdisciplinary fields that combine the principles of engineering and biological sciences with the goal of developing systems capable of replacing, restoring, and/or regenerating damaged or diseased tissues. TE products have only recently entered the market, with the first TE construct receiving FDA-approval in 1996 (Integra™, Integra Lifescience). Since then, the field of TE and regenerative medicine has been rapidly evolving and has incredible potential for advancement through research and growth in the marketplace.

The emergence of TE and regenerative medicine allows for the engineering of multifunctional biomaterials that acts as a scaffold and/or matrix capable of regulating and guiding cellular behavior through molecular and structural interactions. Two basic modes of delivery are used for implanting TE constructs: (1) implanting preformed scaffolds in the area of concern and (2) using an injectable approach (i.e., in situ gelling or flowable hydrogels). Figure 4 highlights the concept of an injectable TE construct. Injectable TE constructs have been shown to be efficient cell carriers,142, 143 enable minimally invasive delivery techniques, and provide superior interfacial contact between the hydrogel and the nearby tissue.144, 145 In contrast, preformed, noninjectable systems have shown limited therapeutic efficacy due to the often inadequate capacity to conform to host defect size and shape and the inability to seed cells deep into the scaffold.146, 147 One important drawback for some injectable polymeric scaffolds is that they can be mechanically weak, limiting their potential in load-bearing applications, such as orthopedic TE.147 However, several means of addressing this issue have been explored—for example, hydrogels reinforced with rigid microspheres, as discussed in more detail in later sections.2, 3

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Figure 2. Chemical gelation via covalent cross-links is another mechanism used to create injectable hydrogels. Multifunctional injectable “liquid” precursors react on mixing to transition from liquid to solid. Injection time depends on kinetics of the chemical cross-linking. This example shows a chemically cross-linked PPODA/pentaerythritol tetrakis(3-mercaptopropionate) gel.

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Figure 3. In a typical drug delivery application, a liquid solution of the drug would initially be injected. On injection, the hydrogel forms (either in response to temperature change or to chemical cross-linking) to entrap the loaded drug and control the release. After depletion of the drug, the hydrogel would typically be resorbed through degradation processes.

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Figure 4. Injectable neural tissue engineering system—Schematic depicting the key elements of an injectable neural tissue system, including cellular components, adhesive molecules, soluble growth factors, and proteolytic degradation sites.

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Rationally designed injectable TE constructs should encompass key design parameters (Fig. 5) to enhance the efficacy in any physiological system.26, 148, 149 An injectable hydrogel system should:

  • i
    evoke a minimal immune/inflammatory response;
  • ii
    match 3D structure and/or mechanical properties of host tissue;
  • iii
    support exogenous cell transplantation and/or favorable host cell infiltration and interaction,;
  • iv
    provide cellular cues through incorporation of bioactive molecules (i.e., extracellular matrix [ECM] proteins, growth factors, and/or chemokines); and
  • v
    degrade at a rate that complements the rate of tissue remodeling.
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Figure 5. Key considerations for rationally designed injectable tissue engineered systems.

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The goal of these design parameters is to produce a fine-tuned microenvironment for cells to interact and behave in a predictable and reproducible manner (e.g., migration, attachment, proliferation, and, in the case of stem cells, differentiation). Recent advances in microfabrication/nanofabrication techniques and the development of innovative synthetic chemistries have provided investigators with a number of tools that allow for exceptional control over engineering TE construct microenvironments. Details regarding interesting developments in scaffold fabrication are described in a relevant review of current techniques and future trends by Hutmacher et al.148 TE strategies also aim to simulate the native cellular microenvironment by including ECM proteins and other molecules that provide signaling to maintain cell survival, differentiation, and migration.150, 151 Synthetic materials typically lack these cues and thus the ability to direct cell fate. Therefore, they often require modification with bioactive molecules via chemical immobilization or passive mixing to influence favorable cell behavior from transplanted and/or surrounding host cells. These biomolecules can range from full-length proteins to sections of known active protein domains (i.e., recombinant proteins) to short active amino acid sequences (i.e., peptides).152–154 This section will highlight injectable hydrogel systems developed for TE and regenerative medicine applications as reported in the recent literature. In keeping with the theme from the previous drug delivery section, the discussion is organized based on the mechanism responsible for in situ gelation.

Physical Gels

Temperature-Responsive Systems

Temperature-responsive and other types of physically cross-linked hydrogels are attractive materials primarily due to the ease of tailoring the physical properties for specific applications. These systems exhibit relatively a biologically compatible in situ setting process and cells, bioactive molecules or even complex 3D structures (microbeads/tubes) can simply be mixed in solution at low temperatures (in materials that exhibit an LCST) and injected in the defect area to induce gelation.

Chitosan-based (i.e., chitosan–GP salt–hydroxyethyl cellulose; CH-GP-HEC) polymers are one class of biodegradable, temperature-responsive polymers. GP salt in this system acts as a mild base that shields electrostatic repulsion between chitosan monomers. Gelation of CH-GP is thought to be a result of the hydrophobic effect (between [BOND]CH3 groups), hydrogen bonding (between [BOND]OH and [BOND]NH groups), and ionic cross-linking.155 Temperature sensitivity is likely due to the strong hydrophobic effects and hydrogen bonding, because the intensity is correlated to temperature. The exact mechanism of gelation is yet to be elucidated, but certain embodiments of the system formulated in high concentration solutions of GP were observed as a solution at room temperature and physiological pH, and undergoes gelation at physiological temperatures (∼37 °C).155, 156 However, gels formulated in high concentrations of GP resulted in poor cell growth, possibly due to excessive osmotic stress. Moreover, lower concentrations of GP increased the LCST of the system drastically.155 As a result, Yan et al. investigated the effects of a second cross-linking agent, HEC. The biocompatibility, structural properties, and gelation temperature are found to be correlated to the HEC content of the material. Their studies also show CH-GP-HEC to be highly biocompatible and capable of promoting cell proliferation with murine MSCs.156

MC is another example of a temperature-responsive hydrogel. Gelation mechanism is primarily due to hydrophobic interactions between methoxy-group containing molecules that produce a reversible 3D physical cross-linked network.9 The resulting physical and gelation properties of MC depend on the molecular weight and degree of methoxy substitution of the MC polymer and the ionic strength of the solvent. Stabenfeldt et al.157 demonstrated that the temperature-responsive injectable physical hydrogel MC functionalized with laminin (ECM protein) significantly improved neural stem cell (NSC) survival, differentiation, and migration compared with native MC in vitro. The same bioactive MC hydrogel protected NSCs from apoptosis in an in vitro model of neural injury.158

Agarose is polymer system that exhibits temperature-sensitive gelation properties due to hysteresis mechanisms. Agarose melts at high temperatures (∼85 °C), then as the temperature is decreased to below the setting temperature (ranging from 17 to 35 °C depending on chemical modifications), a gel is formed due to the formation and subsequent aggregation of double helical structures.159 Jain et al.160 used hydroxyethylated agarose hydrogels to deliver brain-derived neurotrophic factor (BDNF) and active Rho GTPases to stimulate neuronal axons regeneration after spinal cord injuries (SCI). Axonal infiltration through CS-rich regions and the implant material was observed in animal models treated with Rho GTPases and BDNF, relative to controls. However, behavioral studies looking at changes in motor control are missing. One important drawback to this system is the low temperature (17 °C) required for the hydroxyethylated agarose gelation—the investigators required a special cooling system to induce in situ gelation of the material already injected in the animal models.161

Articular cartilage tissue is composed of specialized cells in addition to ECM composed mainly of collagen fibers, glycosaminoglycans (GAGs), and proteoglycans (PGs). Natural repair of cartilage tissue is a slow process where the limiting factor is the availability and transport of new chondrocytes to the area of injury.162 To address this limitation, development of materials for cartilage TE is largely focused on efficient means of delivering chondrocytes to the site of injury. A number of hybrid injectable hydrogel constructs have been proposed where data suggests that the incorporation of cartilage-specific GAGs, PGs (e.g., HA, heparin sulfate, tyramine, and chondroitin sulfate) and peptide sequences can support chondrocyte growth and promote formation of new ECM matrix both in vitro and in vivo.163–166 Materials designed for cell encapsulation, such as chondrocytes, should minimize syneresis. A hydrated environment is essential for cell survival because it mediates transport of nutrients and waste products to and from the extracellular space; syneresis may interfere with this process.162 For example, poly(NIPAAm) homopolymer gels exhibit shrinking behavior to the extent that they are impractical for cell encapsulation.49 However, several formulations of poly(NIPAAm) have shown promise for cell delivery in various TE applications, largely though the incorporation of hydrophilic segments such as AAc167 or PEG168 that counteract syneresis. It is worth noting that that LCST and mechanical properties of the gel are usually also a function of AAc/PEG functionalization169 and high concentrations (>25 wt %) can cause LCST to increase over 37 °C.29 Additionally, MC grafted with poly(NIPAAm) yields a reversible temperature-responsive hydrogel with little syneresis.170 Sá-Lima et al.162 explored the capacity of this system to encapsulate murine chondrocytes and found the material to be nontoxic to fibroblast and chondrocytes in vitro while enhancing GAG production. Furthermore, the authors provide evidence of transplanted cells maintaining chondrocytic lineage and signs of PG- and GAG-containing ECM production after 4 weeks, although it is worth noting that the poly(NIPAAm) portions of this material are nondegradable. Subsequent in vivo studies are required for this material to further determine its viability for TE applications. Another temperature-responsive formulation of poly(NIPAAm-co-AAc) physically mixed with very high-molecular-weight HA was tested as an injectable material for rabbit chondrocytes and transforming growth factor β-3 (TGF-β3) delivery.171, 172 In addition to excellent biocompatibility and biodegradability, HA also exhibits superior fixation to bone and cartilage tissue relative to collagen, alginate, and chitosan.171 HA has therefore received considerable research attention for related TE applications. Na et al.171 found that chondrogenic differentiation and GAG/collagen production were directly correlated to both the presence of HA and TGF-β3 (controls included matrices with no TGF-β3 or HA) in vivo.

Hydrogels of poly(NIPAAm-co-AAc) with peptide sequences (arg-gly-asp (amino acid sequence) (RGD); Arg-Gly-Asp, bioactive tripeptide sequence commonly found within ECM proteins173) grafted on poly(AAc) polymer chains were evaluated by Chung et al.174 to create a degradable semi-IPN mediated through oligopeptide cross-linkers that are specifically cleaved by the MMP family of collagenases for bone regeneration. This class of material is unique due to the ability to independently tune matrix properties (mechanical strength, degradability, LCST, etc.) and parameters that offer biological cues (peptide/bioactive molecule incorporation). Other copolymers of poly(NIPAAm), such as, with methoxy PEG methacrylate (mPEGMA), producing poly(NIPAAm-co-mPEGMA) with PEG grafts have shown the capacity to maintain MSC cell viability and promote proliferation in 3D in vitro cultures.29 Pollock et al.29 demonstrated that the LCST and rheological properties of the gels are dependent on the copolymer composition, PEG graft concentration, and molecular weight, without significantly affecting swelling properties. The tunable nature of such materials that produce highly reproducible matrix properties are extremely attractive for a wide range of TE applications.

Ion-Mediated Gelation Systems

Ion-mediated gelation systems are another class of injectable hydrogels that respond to local or controlled ionic alterations to gel in situ. Alginates are biocompatible, relatively low cost, and form under relatively mild conditions, producing ionotropic gels suitable for various medical applications. Commercially available alginates are usually extracted from brown seaweed algae and contain blocks or alternate sequences of L-guluronic (GA) and D-mannuronic (MA) acids. The viscosity, stiffness, and pore size of alginate hydrogels are dependent on the MA:GA ratio and the molecular weight. Larger number of GA blocks result in higher stiffness due to diaxial linkages, whereas a high-MA:GA ratio results in a smaller average pore size.175 These systems also exhibit structural properties that are a function of temperature. At lower temperature, diffusion of cross-linking agents (such as, Ca2+) are slower, creating a more ordered 3D hydrogel structure that may be advantageous for some TE applications.176

Alginates, however, are not inherently bioactive and undergo relatively slow degradation under physiological conditions. The issue of bioactivity gathered attention from several authors where immobilized peptide177 (RGD) or controlled release of growth factor122 (VEGF) showed improved control over cellular behavior in alginate hydrogels. Means of mediating faster degradation rates have been also been investigated. In one case, gamma irradiation was used to reduce the molecular weight from 300 to 25 kDa when exposed to 100 kGy radiation.178 However, this strategy has inherent complications due to the presence of possibly cytotoxic (without radioprotectants179) ionizing radiation. Partial oxidation of alginates has been shown to provide control over degradation rates via hydrolysis.180 Using this strategy, Kim et al.181 designed an RGD-functionalized alginate-based hydrogel for adipose-derived stem cells delivery. This specific formulation degraded under physiological conditions where 80% mass loss can be expected in a period of 40 days.123, 181 The resulting hydrogels demonstrated newly generated adipose tissue 10 weeks after injection with evidence of well-organized vascularization.181 RGD-modified alginate-based hydrogels have also been tested in vivo for improved left ventricle function and angiogenesis in rodent models with ischemic cardiomyopathy.182In vitro studies with human umbilical vein endothelial cells demonstrated significantly higher cell counts, adhesion, and proliferation rates in RGD-modified substrates relative to alginate only controls after 7 days of culture. The in vivo studies demonstrated a statistically significant increase in left ventricle function and arteriogenesis, as well as improved systolic and diastolic wall thickness using both RGD-modified and unmodified alginate gels, relative to PBS controls. It is important to note that significant differences were not observed between the two types of alginates for the above criterion. However, an increase in arteriole density was significantly higher in the RGD-modified alginate compared with both alginate and PBS controls.182

Alginates have been combined with calcium phosphates for bone TE,147 producing an injectable system with appropriate mechanical strength (similar to cancellous bone), while promoting favorable cell behavior. Another interesting system uses calcium-loaded temperature-sensitive liposomes with alginate to initiate gelation at physiological temperatures.183 Above the lipid chain melting temperature, permeability of Ca2+ through the liposome membrane increases.184 The melting can be modulated to be lower than 37 °C through alterations in the structural properties of the membrane bilayer185. Thus, when subjected to physiological conditions, Ca2+ efflux from liposomes leads to alginate cross-linking.

Self-Assembling Peptides

Self-assembling peptides are attractive materials due to their ability to form scaffolds under physiological conditions. (arg-ala-asp-ala) arginine-alanine-aspartic acid-alanine (RADA)-16I is one commonly used self-assembling peptide that has been evaluated for brain,186 spinal cord,187 and cardiac188 TE applications. In aqueous medium, RADA-16I forms a stable β-sheet structure. The addition of monovalent cations, as observed under physiological conditions, results in the formation of a complex nanofiber scaffold.189 RADA-16I covalently coupled to bone marrow homing motif (BMHP1) injected in spinal cord of rat models led to an increase in expression of ECM remodeling proteins and nervous tissue regeneration in rat SCI models.190 Another study reportedly increased the stability of the bioactive molecule and NSC adhesion by incorporating a spacer segment containing four glycine residues between the core of RADA-16I and BMHP1 relative to systems containing shorter spacers in vitro.191 Similarly, a bioactive RADA-16I system composed of RADA-16I covalently coupled with the peptide sequence RGDSP supported MSC-derived cardiac stem cells differentiation into cardiomyocytes.192 RGDSP was conjugated to the self-assembling peptide (SAP) using solid-phase synthesis extension at the C-terminus of SAP polymer chains without significant alterations in mechanical or structural properties. These scaffolds not only showed high biocompatibility, but also demonstrated cytoprotective properties for transplanted cells in vivo, inhibiting apoptotic and necrotic cellular activity under oxygen and glucose deprivation compared with nonfunctionalized RADA-16I and simple cell transplantation.193 Moreover, they report improved cardiac function (quantified through echocardiography) with both formulations of RADA-16I, with the functionalized SAP showing statistically significant improvements over controls and other experimental groups.193

ELPs, derived from the naturally occurring ECM protein, elastin, are another common form of SAP. Elastin is a polymer of tropoelastin, and it produces highly cross-linked insoluble structures, although the exact mechanism of polymerization has yet be described in detail.194 The polymer has two general functional domains, one is heavily hydrophobic (rich in valine, proline, and glycine residues) and the other is responsible for mediating cross-links (between lysine residues catalyzed by lysyl oxidase).195 This reversible polymer exhibits exceptional stability and sensitivity to temperature, pH, and ionic strength, although it exhibits low rates of degradation.195 ELPs are an analog of elastin with repeating five peptide sequence VPGXG, where X is any peptide other than proline, which leads to a large number of possibilities for interesting gel properties.196, 197 By changing the sequence of the monomer, ELPs can be designed to have different gelation temperatures and viscoelastic properties. Furthermore, they can be efficiently produced in large scale through genetic modification in both eukaryotic and prokaryotic (E.coli) cell lines with reproducible peptide sequence and MWs.198 ELPs have been widely used for TE applications such as cartilage,199–201 vascular grafts,202, 203 soft tissue,204, 205 prevention of scarring after laminectomy,206 and ocular207 TE.

Chemically Cross-Linked Hydrogels

Enzymatically Triggered Systems

A considerable amount of research interest is placed on developing hydrogels that undergo cross-linking mediated by enzymes (both endogenous and exogenous enzymes). Such reactions can allow for a biologically compatible in situ gelling process that does not require conditions that might negatively impact the bioactivity of specifically functionalized polymers. Additionally, these reactions are usually very specific and thus avoid side reactions and limit concerns of toxicity.

Fibrin is a classic example of an enzymatically triggered injectable hydrogel that has been used in a variety of TE applications.13–16 Fibrin constitutes the native hemostatic agent and provides a polymer lattice for clot formation during coagulation. The coagulation cascade is stimulated through either the extrinsic or intrinsic pathway that ultimately activates thrombin, which in turn converts soluble circulating fibrinogen into insoluble self-assembling fibrin polymers.208 The plasma transglutaminase, factor XIIIa, then catalyze the formation of an isopeptide bond between specific glutamine and lysine residues on fibrin monomers to generate a stable cross-linked network.208 Multiple research groups have harnessed this native enzymatic hydrogel for various TE applications. For example, co-delivery of neurotrophin-3 and platelet-derived growth factor from fibrin matrices significantly enhanced neural progenitor cell survival and differentiation in a spinal cord injury mouse model.14 In this case, Johnson et al. used bidomain affinity peptides (ATII) to improve NSC viability and direct cell fate to produce populations of primarily neurons or oligodendrocytes. In vivo studies demonstrated statistically significant improvement in NPC survival, proliferation, and favorable differentiation behavior using fibrin scaffolds containing both NT-3 and PDGF after 2 weeks compared with controls.14 Other groups have successfully altered the polymerization dynamics,13 proteolytic degradation,15 or cross-linking mechanisms16 to tune fibrin for specific cell/physiological applications.

Another example of enzymatically triggered gelling system is described by Jin et al. where they characterized hydrogels of dextran conjugated to tyramine that may be applicable for a variety of TE applications, such as cartilage regeneration.108, 164 They report a system with tunable gelation time (5 s to 9 min depending on tyramine functionalization and concentration of supporting enzyme horseradish peroxidase and H2O2), mechanical properties (storage modulus ranging from 3–41 kPa), and degradation (through incorporation of diglycolic anhydride).108 One apparent downside to these systems is that the fast gelation periods are only valid in the presence of H2O2, which may induce cytotoxicity.209 Although a series of studies conducted by Kurisawa et. al. have shown favorable MSC attachment on gelatin-hydroxyphenylpropionic acid and HA-based hydrogels cross-linked using horseradish peroxidase (HRP) and H2O2, relative to culture plates.210, 211 Using the same cross-linking reaction, dextran has also been conjugated to heparin and applied for cartilage TE.163 Use of horseradish peroxidases has not been limited to dextran. Tyramine-modified HA gels have been reported to maintain suitable bioactivity and biocompatibility in vivo.212 Polysaccharide-based systems such as chitosan213 and alginate214 that incorporate enzymatic reactions have also been proposed for cartilage TE and cell transplantation applications, respectively.

Schiff Base-Mediated Gelation

Schiff base reactions can be used to achieve in situ gelling without the need for any chemical cross-linking agents,166 although they can nonspecifically react with amines as mentioned previously. For example, water soluble N-succinyl-chitosan and aldehyde-functionalized HA can be used to synthesize a composite hydrogel with variable gelation time, microstructure, mechanical strength, and degradation for cell delivery that may be suitable for various TE applications.166 The investigators focused on cartilage repair and showed improved bovine articular chondrocyte survival, proliferation, and maintenance of chondrocyte phenotype in vitro. The 50:50 ratio of N-succinyl-chitosan to aldehyde-functionalized HA demonstrated relatively quick gelation time (∼1 min), remained well hydrated due to the presence of hydrophilic groups, showed favorable compressive modulus (25 kPa), and provided statistically significant enhancement of cell viability relative to other formulation ratios.166

In another study, dextran, oxidized using sodium periodate, was cross-linked with adipic acid dihydrazide producing a hydrogel that may be suitable for various TE applications.215 Dextran is an attractive material due to its availability, flexibility in terms of molecular weight, and the presence of a large number of hydroxyl groups that allow for chemical modification. In this study, the resulting hydrogels were nontoxic, exhibited relatively quick gelation kinetics (<4 min), and tunable degradation, as well as mechanical strength (7–32 kPa) dependent on the cross-linking density.215 In a different study, chondroitin sulfate was multifunctionalized with methacrylate (disaccharides substituted using glycidyl methacrylate; 12% efficiency) and aldehyde groups (hydroxyl groups oxidized using sodium periodate to produce aldehydes; 70% efficiency) for cartilage TE.216 The methacrylate groups then can polymerize to produce a cross-linked structure, whereas the aldehyde functional groups react with amine groups located on host tissue through Schiff base interactions to produce a bioadhesive material that improve structural stability between implanted hydrogel and host tissue.216, 217 The in vivo studies suggest improved cartilage tissue regeneration in goat animal models treated with the adhesive material and hydrogel relative to untreated defects.216 However, the controls for this particular goat study did not include a group treated with hydrogel alone, and therefore, the exact effect of the adhesive is unclear. In addition, Schiff base reactions have been used to cross-link HA to cellulose derivatives218 and to produce HA-lipid conjugates219 for various applications.

Alternative Mechanisms for Chemically Cross-linked Gels

Polymers have also been designed to undergo chemical cross-linking and polymerization on exposure to visible or ultraviolet radiation (i.e., mediated by photoinitators that cross-link vinyl group-containing polymer strands).220 PEG, for example, has been functionalized with photoactivated cross-linking end groups such as polyrotaxanes or methacrylates (an initiator such as 2,2-dimethoxy-Z-phenyl-acetophenone is required for methacrylate cross-linking) to produce hydrogels.221, 222 These systems also have their drawbacks that include cytotoxicity due to the photoinitiators and its related cross-linking reactions which can be exothermic, causing local cell death.223 Gelation induced though disulfide bonding has been tested using thiol-functionalized HA derivatives. The materials exhibited good biocompatibility and encapsulated fibroblast cells were showed promising viability and proliferation in vitro.224 However, the gelation time for the material is too high (∼30 min dependent on formulation, pH, and temperature) for most injectable TE applications.224, 225 Another means of mediating in situ chemical gelation is through Michael addition between thiols and acrylates as discussed in previous sections. For example, thiolated HA were conjugated to unsaturated esters and amides of various PEG strands (PEGDA showing the highest rate of reaction) and evaluated in vivo (mice models) seeded with fibroblast cells.225 Their in vitro studies showed that the material is biocompatible, promoted fibroblast proliferation (10 × population in 28 days compared with controls) and production of ECM proteins (collagen type I).225 The in vivo studies demonstrated that material gelation was sufficiently rapid (∼9 min), and significant tissue necrosis or damage to surrounding cells was not observed. Although the authors looked at degradation rates in vitro (between 6–27% in 25 h in the presence of a different bovine testicular hyaluronidase concentrations),225 data regarding in vivo degradation was not provided.

Diels-Alder (DA) reactions have also been investigated for various biological applications due to their high yields, specific reaction (between a diene and a dienophile) that occurs in mild aqueous conditions.226, 227 No catalysts are required for the reaction and it produces no byproducts. Recently, a DA reaction-based bioconjugated HA hydrogel was characterized by Tan et al.228 Although the gels are biocompatible and degradable, the gels required over 20 min to set and reached a storage modulus under 20 Pa, which is very weak.

Other Means of Gelation

Pregelled/Preformed Injectable SystemsCeramic materials such as hydroxyapatite and β-tricalcium phosphate (β-TCP) have been widely used for bone TE due to their biocompatibility and osteoconductive properties. Besides being noninjectable, the osteoinductive (directing osteoprogenitor cell differentiation) and osteogenic (promoting formation of new bone tissue) capabilities of these materials are limited.229 Injectable systems for bone regeneration have been fabricated using a variety of natural biomaterials such as alginate,230 collagen,231 and fibrin.232 Although these materials exhibited favorable biocompatibility, their ability to conform to defect shape is limited both in vitro and in situ, while also they lacking the mechanical strength required for bone TE.233 An improved strategy would incorporate the osteogenic capabilities of particular cells and the osteoinductive properties of signaling molecules in an injectable system which mimics the structural and mechanical characteristics of bones.234 Matsuno et al.233 attempted to develop such a material composed of β-TCP microbeads and alginate which acts as a scaffold for osteogenic cells and a carrier for growth factors. They passed alginic acid solution through β-TCP microbeads coated with calcium chloride to initiate instantaneous alginate hydrogel formation around the beads, producing an injectable matrix with appropriate 3D structure and mechanical properties with a compressive strength of 69 kPa (data collected in dry conditions).233 Compressive strength for bones are considerably higher (cancellous bone >2 MPa; cortical bone >100 MPa).235 Their invivo studies indicated a formation of calcified bone-like tissues, suggesting that this composite system is capable of supporting bone growth using a minimally invasive procedure. Because the hydrogel/microbead material is polymerized before insertion into the defect area, monomer concentration becomes a limiting factor. In this case, the authors were limited to using a maximum 2 wt % monomer before the material became too viscous to be injectable. Data regarding the dependence of new tissue formation to size and concentration of microbeads would be interesting, but were not explored.

Microspheres have also been studied as means for cell transplantation due to the ease of cell seeding before injection and their ability to promote adequate nutrient transfer and permit host cell infiltration after implantation. Hydrogel microsphere formulations include nonporous chitosan (175–250 μm diameter beads) for adipose-derived stem cell delivery have been demonstrated to induce migration from the bead surface to a simulated ECM environment made of a collagen gel matrix.236 A different study characterized the in vitro viability of gelatin microspheres (70–300 μm in diameter) cross-linked with genipin (reactive toward amine groups) as a vessel to encapsulate human fetal osteoblast cells (maximum cell attachment attained in 2 days with maintained cell viability after 7 days incubation).237

Shear-Thinning Gels

Shear-thinning gels also have the potential of being preformed ex vivo with appropriate physical properties and delivered to the desired location, usually by injection, followed by quick recovery of the gel's original properties. For example, Gupta et al.238 developed a fast in situ gelling (sol-gel transition in 2 min) system composed of biodegradable blend of HA (2%) and MC (7%) for SCI designed to be delivered intrathecally. In this case, their blend of HA-MC degraded to 90% of its original mass in 14 days. However, the degradation rate slowed drastically thereafter. An interesting system composed of chitosan and silicate laponite nanodisks with PEO grafted on its surface (SLD-PEO) was characterized as an injectable material for TE.239–241 The exact gelling mechanism has yet to be described in detail, but hydrogen bonding is hypothesized to play a large role in this system.242 The system was tested in vitro with 3T3 fibroblast cells, and it was found to be biocompatible, with cells showing a more spread morphology on gel films with higher chitosan content.239 Studies were also conducted with mouse preosteoblast cells, and the material was capable of maintaining cell viability and promoted differentiation into osteoblasts in vitro.240 This nanocomposite hydrogel holds promise for various TE applications; however, the long-term viability of these materials (biocompatibility, degradation, and inflammatory response) are yet to be determined conclusively through in vivo studies.

INJECTABLE HYDROGELS FOR SPACE-FILLING APPLICATIONS

  1. Top of page
  2. Abstract
  3. INTRODUCTION
  4. INJECTABLE HYDROGELS FOR DRUG DELIVERY
  5. INJECTABLE HYDROGELS FOR TE AND REGENERATIVE MEDICINE
  6. INJECTABLE HYDROGELS FOR SPACE-FILLING APPLICATIONS
  7. CONCLUSIONS
  8. REFERENCES AND NOTES
  9. Biographical Information
  10. Biographical Information
  11. Biographical Information
  12. Biographical Information

Because injectable hydrogels assume the shape of the environment into which they are placed, they are suitable choices for some materials which neither release drugs nor serve as TE matrices, but instead provide a structural or space-filling function. For plastic surgery or tissue filling applications, minimally invasive administration and minimized scar formation make injectable materials ideal. Using embolic agents for diverting or eliminating, flow in blood vessels can be useful for treatment of arteriovenous malformations or aneurysms. Similar materials might also be useful for contraception. Injectable hydrogels have also been investigated as tissue barriers for adhesion prevention or as bulking agents to reinforce weakened tissue.

There are several design considerations for hydrogels to be used in space-filling applications. Typically, the primary consideration is that hydrogels tend to be relatively weak, and so both the modulus and the viscoelastic character (i.e., phase angle) of the gel must be suitable for the application. For a material to be resilient and resistant to creep in response to constant or periodic stresses, it must lose very little of the energy put into it—that is, the material must be highly elastic (phase angle near zero). Deswelling is likely to be problematic in applications such as endovascular embolization where full occlusion of a space is a necessary function of the material. Swelling in the same applications should also be limited to avoid exerting undue forces on weakened vessel walls. Other considerations include bioadhesion, solute transport through the device, biocompatibility, sterilization, deliverability (such as through a catheter), delivery time, gelation time, and degradability (and if so, degradation time).

In this section, injectable space-filling hydrogels reported in the recent literature will be presented according to the type of cross-linking responsible for in situ gelation (physical, chemical, or multiple). Although many materials such as prefabricated gels and silicone are used as injectable fillers in cosmetic surgery243–246 and as tissue bulking agents,247–249 this section will focus on materials which undergo an increase in mechanical properties after administration.

Physical Gels

Physical gels form associative networks without covalent cross-links connecting the constituent molecules. Although chemical gels tend to be highly elastic, some physical gels tend to be viscoelastic,53, 250 which is disadvantageous for most space-filling applications. For example, physical gels of poly(N-isopropylacrylamide) are formed due to hydrophobic interactions between side groups.251–253 During deformation of the material, hydrophobic interactions between side groups can be “broken” and then reformed between different combinations of polymer side groups. The interchangeable nature of the hydrophobic interactions can result in a viscous or viscoelastic gel.36, 53, 169 The phase angle can be controlled by a number of factors including concentration,29, 250 polymer architecture,169 light pre-cross-linking,48, 56in situ cross-linking,37, 254 or molecular weight.29

One application of injectable hydrogels is intramyocardial injection following myocardial infarction slow the progression of heart disease by reducing stresses on the infarcted tissue.255, 256 Degradable physical gels of poly(NIPAAm-co-AAc-co-HEMA-PTMC) were shown to lead to improved contractility, ventricular wall thickness, and capillary density at 8 weeks after injection.257 Although no rheological data was reported, the gels were initially able to be stretched to ∼200% strain. However, it is unclear whether the gels were effective after 8 weeks and especially after degradation of the gel. If not, gel would need to be reinjected to be effective for more than a few months, which would be impractical. Self-assembling peptide hydrogels have been investigated for myocardial injection as well, but gels alone did not lead to improved outcome as measured by cardiac function or ventricular wall thickness.131, 258, 259 These hydrogels did show encouraging results in a rat model when releasing growth factors in a TE approach,258, 259 but the data from hydrogels without growth factors suggests that the gels' mechanical properties are insufficient to provide an improved outcome.

Copolymer architecture was used by Lin and Cheng169 to make injectable gels of elastic modulus between 1000–2500 Pa using block and star copolymers with a central PEG group and terminal poly(NIPAAm) groups. The materials exhibited reversible gelation due to the poly(NIPAAm) groups above the LCST of 26–29 °C. These gels show limited syneresis, a fast injection time, and some elasticity (tan δ between 0.24 and 0.62). Stereocomplexing degradable star copolymers containing eight-arm PEG and terminal segments of either poly(D-lactic acid) (PDLA) or PLLA were shown to exhibit lower gelation concentrations and higher strengths when equimolar amounts of each enantiomer were present compared with PEG-PLLA alone and compared with triblock copolymers.260 These gels were relatively strong, with the gels having up to 14 kPa storage modulus and low phase angle.

Lightly cross-linked temperature-responsive gels based on poly(NIPAAm) have been developed for space-filling applications. Below the LCST, light cross-linking allows for the material to behave as a liquid, and yet above the LCST allows for improved elasticity characteristic of a chemical gel. In general, temperature-responsive materials with light precross-linking tend to be weaker than temperature-responsive materials which undergo cross-linking in situ because the cross-link density is limited to a low fraction of co-monomers. However, these polymers do not have reactive groups present, which may improve their biocompatibility. A lightly cross-linked copolymer of poly(NIPAAm-co-butyl methacrylate) lightly cross-linked with PEG-diacrylamide groups56, 57 was investigated for embolization of cerebral aneurysms alone or in combination with protection devices such as stents and coils.261 Because of the low LCST of 13–18 °C, a cooling jacket had to be continuously flushed with 4 °C saline to prevent gelation inside the microcatheter containing the polymer before delivery. Complete occlusion was observed after 14 days when the polymer was coadministered with a stent or both stent and coil, but partial recanalization was observed in aneurysms with neither, likely due to the low modulus of the polymer (storage modulus ∼1000 Pa). It is also difficult to assess the effectiveness of the polymer alone because the study was only carried out for 14 days and the untreated control aneurysms did not rupture.261

Vernengo et al.262 used a similar polymer design of poly(NIPAAm) lightly cross-linked with PEG-dimethacrylate of various molecular weights for nucleus pulposus replacement. These materials are injectable and are reported to maintain their mass for up to 90 days in vitro, with compressive moduli over 50 kPa seen in several of the formulations, particularly those with lower molecular weight PEG (below 4.6 kDa). However, using this range of PEG resulted in equilibrium water content that was near or below 50 wt % when the polymers were formulated at 15 wt %. Therefore, the gels shrank to ∼30% of their initial volume within 90 days. This degree of shrinking could be problematic for a load-bearing material, due to concentration of stress on the shrunken implant, absence or deformation of the material potentially allowing for bone–bone contact, or even risk of the implant being displaced.

Another material design used for intramyocardial injection involves the use of a dextran and PCL-based degradable macromolecular cross-linker subsequently copolymerized with NIPAAm to yield a precross-linked injectable hydrogel.263 The materials had storage modulus up to 1500 Pa depending on the cross-linker content. This material was injected into infarcted heart tissue in rabbits 4 days postinfarct.264 After 4 weeks, they showed increased ejection fraction and lower left ventricular end diastolic diameter (LVEDD). Although no material was observed at 4 weeks in the histology due to degradation of the cross-linker, it is unclear if this material has longer term effects either on cardiac function or biocompatibility due to its partial nondegradability.

Ionic cross-linking of alginate gels that occurs based on local tissue calcium concentration has also been used as an injectable bulking agent in myocardium.265 Alginate solution was injected 7 days postinfarct in rats. After 8 weeks, animals receiving alginate injections showed increased wall thickness and lower LVEDD. It was shown by staining that most of the material was not present in the infarct by 6 weeks. However, as with other myocardial bulking agents, it is unclear how the heart performs well past degradation.

In situ Cross-Linked Gels

Covalently cross-linked materials based on liquid or injectable precursors are useful for space-filling applications because they form highly elastic network polymers with phase angles near zero. These materials are not as prone to creep as physical gels, and therefore tend to be suitable for applications that require load bearing or long-term exposure to stress. The precursor materials can potentially be delivered through a catheter without solidification, and gelation time can be easily controlled.

In situ photopolymerization of a PEG-based cross-linker and NVP has been used to form a thin barrier on the interior of blood vessels to reduce thrombosis following vascular injury.266 The photoinitiator was first adsorbed onto the vessel surface followed by polymerization, creating a hydrogel with a nearly uniform thickness conforming to the surface of the vessel wall. The liquid precursor materials can also penetrate into the tissue, leading to adhesion between the gel and the tissue. Gel formation took place within seconds, and gel thickness was well controlled by adjusting the polymerization time. The gels significantly reduced thrombosis and reduced neointimal thickening by 80% after 2 weeks in a rabbit balloon injury model.

In situ cross-linking between two soluble or liquid precursors has been used in a variety of bioinert or structural materials. One of the strongest is a reverse-emulsion (water-in-oil) system developed by Vernon et al.267 consisting of three mass equivalents of a mixture of two organic precursors (one functionalized with thiols and the other with acrylates) mixed with one equivalent of aqueous buffer, resulting in 75 wt % gels. Reverse-emulsion gels made of tetrafunctional pentaerythritol tetrakis(3-mercaptopropionate) (QT) and pentaerythritol triacrylate (TA) had ultimate strength in compression exceeding 6 MPa and ultimate deformation of 37%. Stabilization with a surfactant was required for gelation to provide high surface area between the oil phase containing the precursors and the alkaline aqueous phase, which activates the reaction. A similar system based on QT and poly(propylene glycol) diacrylate (PPODA) has recently been developed for intracranial aneurysm embolization (Fig. 6).268, 269 The gelation kinetics of this system are desirable for embolization because there is a time after mixing during which the material is still injectable as a liquid with a phase angle near 90° (purely fluid vs. elastic) and low viscosity. Then, gelation occurs quickly, resulting in a material with low phase angle near 0° (purely elastic). A number of factors affect the gelation time and final properties of these materials, including mixing time and the composition of the aqueous phase.268In vivo studies of the PPODA-QT system in swine showed neointimal growth at 1 month when the material was used to incompletely fill (80–90 vol %) aneurysms to avoid re-entry of material into the parent vessel.269 Based on the high strength of these gels, they were able to be deployed successfully without a second device such a stent to retain the gel inside the aneurysm. Also the deliverability time of these strong gels is much lower and the delivery process easier than for the solvent-exchange Onyx system (poly(ethylene-co-vinyl alcohol) delivered in DMSO)270, 271 currently approved in the US.

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Figure 6. Top panel: PPODA/pentaerythritol tetrakis(3-mercaptopropionate) (QT) waterborne polymer gels forming in vivo in a swine model of a vascular defect. Bottom panel: time lapsed frames demonstrating injection and gelation of the PPODA/QT system in a glass aneurysm model.

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Natural-synthetic hybrid in situ cross-linking materials have been developed for bioinert applications as well. Cloyd et al.272 screened a variety of gel types for mechanical matching of human nucleus pulposus and concluded that in situ cross-linking hydrogels based on hyaluronic acid and PEG-grafted chitosan was an appropriate match. These materials had moduli in compression between 5 and 20 kPa. Hybrid hydrogels based on HA and MC cross-linking by aldehyde reaction with hydrazides showed moderate cytotoxicity in vitro, but were effective in preventing peritoneal adhesions in vivo in a rabbit cecal injury-side wall defect model.218 Gels with MC were shown to be stronger than in situ cross-linking HA gels, but all of the gels tested were weak, with shear moduli below 500 Da. These gels exhibited slower degradation than HA-based gels, dependent on the type of cellulose derivative used. This material design is suitable for prevention of peritoneal adhesions because complete surface coverage and biocompatibility are primary concerns, while the strength requirement is relatively low.

So-called “fibrin glue” forms on the conversion of fibrinogen to fibrin by thrombin, and so gels can be made on demand by mixing thrombin and fibrinogen solutions together—for example, using a double-barreled syringe. These gels have good biocompatibility, are degradable, and can form gels rapidly. Fibrin glues have long been used as surgical sealants,273–275 and more recently have been evaluated for intramyocardial injection after infarction. Mukherjee et al.276 evaluated the injection of a fibrin-alginate composite gel into infarcted myocardium in pigs and found that infarct expansion was prevented for 2 weeks after injection. Promising results have also been obtained using fibrin glue in the same application both alone and seeded with skeletal myoblasts in smaller animal models.277, 278

Physical–Chemical Gels

Gels that cross-link both physically and chemically have potential to combine the advantages of fast gelation typical of temperature-responsive physical gels with the strength and elasticity of a covalently cross-linked material. Combining poly(NIPAAm-co-HEMA-acrylate) with either thiol-functionalized poly(NIPAAm)279, 280 or thiol-functionalized cross-linkers50 yields stronger gels with improved elastic properties at low frequency compared with physical gels. Alternatively, thiol-functionalized poly(NIPAAm) forms strong and elastic physical–chemical gels on mixing with PEGDA below the LCST, which, when heated, reach storage moduli near 1 MPa, an exceptionally high value for a hydrogel.281 Alternatively, NIPAAm-based macromers can be polymerized in situ to obtain physical–chemical gels.282 Yet some frequency response is still observed in physical–chemical gels when allowed to cross-link above the LCST, and the gel moduli plateaus at lower values when the gels are heated to 37 °C more quickly,280 which indicates that cross-linking may be incomplete when the polymers must cross-link at temperatures above the LCST. These materials also are nondegradable.

Degradable physical–chemical gelling systems based on Michael addition reactions between poly(organophosphazene)s with other poly(organophosphazene)s or PEG-based cross-linkers have recently been reported.254, 283, 284 Some of these gels with eight-arm PEG-based cross-linkers254 showed the highest storage moduli (over 40 kPa) and slow degradation, with over half the gel mass remaining after 90 days after subcutaneous injection in mice. FT-IR data also suggests that the gels are capable of cross-linking at 37 °C over about 40 min. Although the poly(organophosphazene) acrylated precursor was found to be somewhat cytotoxic even at a very low concentration of 0.05 wt %, the same concentrations of the cross-linked hydrogel led to cell viability similar to cells without polymer.

CONCLUSIONS

  1. Top of page
  2. Abstract
  3. INTRODUCTION
  4. INJECTABLE HYDROGELS FOR DRUG DELIVERY
  5. INJECTABLE HYDROGELS FOR TE AND REGENERATIVE MEDICINE
  6. INJECTABLE HYDROGELS FOR SPACE-FILLING APPLICATIONS
  7. CONCLUSIONS
  8. REFERENCES AND NOTES
  9. Biographical Information
  10. Biographical Information
  11. Biographical Information
  12. Biographical Information

Injectable hydrogels are an important class of materials in light of the drive in the clinic toward minimally invasive procedures. Developments in polymer science provide opportunity for more sophisticated injectable hydrogels with varied and useful properties. As is demonstrated in this review, exciting research is underway for many different systems. Although a number of injectable hydrogel systems have arrived at the clinic and have impacted the health care of many patients, a significant challenge for future work remains translating the multitude of promising materials reported in the research literature into innovative and impactful products in the clinic by demonstrating their biocompatibility and clinical efficacy.

REFERENCES AND NOTES

  1. Top of page
  2. Abstract
  3. INTRODUCTION
  4. INJECTABLE HYDROGELS FOR DRUG DELIVERY
  5. INJECTABLE HYDROGELS FOR TE AND REGENERATIVE MEDICINE
  6. INJECTABLE HYDROGELS FOR SPACE-FILLING APPLICATIONS
  7. CONCLUSIONS
  8. REFERENCES AND NOTES
  9. Biographical Information
  10. Biographical Information
  11. Biographical Information
  12. Biographical Information

Biographical Information

  1. Top of page
  2. Abstract
  3. INTRODUCTION
  4. INJECTABLE HYDROGELS FOR DRUG DELIVERY
  5. INJECTABLE HYDROGELS FOR TE AND REGENERATIVE MEDICINE
  6. INJECTABLE HYDROGELS FOR SPACE-FILLING APPLICATIONS
  7. CONCLUSIONS
  8. REFERENCES AND NOTES
  9. Biographical Information
  10. Biographical Information
  11. Biographical Information
  12. Biographical Information
Thumbnail image of

Derek Overstreet is currently a graduate student in Biomedical Engineering at Arizona State University. He received his Bachelor's degree (2008) in Biomedical Engineering from Case Western Reserve University in Cleveland, OH. His research is focused on the development of improved temperature-responsive injectable hydrogels for drug delivery and embolization applications.

Biographical Information

  1. Top of page
  2. Abstract
  3. INTRODUCTION
  4. INJECTABLE HYDROGELS FOR DRUG DELIVERY
  5. INJECTABLE HYDROGELS FOR TE AND REGENERATIVE MEDICINE
  6. INJECTABLE HYDROGELS FOR SPACE-FILLING APPLICATIONS
  7. CONCLUSIONS
  8. REFERENCES AND NOTES
  9. Biographical Information
  10. Biographical Information
  11. Biographical Information
  12. Biographical Information
Thumbnail image of

Dipankar Dutta received his B.S. (2010) from the Fischell Department of Bioengineering at the University of Maryland, College Park. He is currently pursuing his Ph.D. under the supervision of Dr. Sarah Stabenfeldt at Arizona State University. His research focuses on the development of novel biomaterials for tissue engineering in traumatic brain injury applications.

Biographical Information

  1. Top of page
  2. Abstract
  3. INTRODUCTION
  4. INJECTABLE HYDROGELS FOR DRUG DELIVERY
  5. INJECTABLE HYDROGELS FOR TE AND REGENERATIVE MEDICINE
  6. INJECTABLE HYDROGELS FOR SPACE-FILLING APPLICATIONS
  7. CONCLUSIONS
  8. REFERENCES AND NOTES
  9. Biographical Information
  10. Biographical Information
  11. Biographical Information
  12. Biographical Information
Thumbnail image of

Sarah E. Stabenfeldt is an Assistant Professor in the School of Biological and Health Systems Engineering at Arizona State University. She received a B.S. in Biomedical Engineering from Saint Louis University (2002) and a Ph.D. in Bioengineering from Georgia Institute of Technology (2007). Dr. Stabenfeldt's research interests focus on regenerative strategies for repair of neural injury including neural tissue engineering.

Biographical Information

  1. Top of page
  2. Abstract
  3. INTRODUCTION
  4. INJECTABLE HYDROGELS FOR DRUG DELIVERY
  5. INJECTABLE HYDROGELS FOR TE AND REGENERATIVE MEDICINE
  6. INJECTABLE HYDROGELS FOR SPACE-FILLING APPLICATIONS
  7. CONCLUSIONS
  8. REFERENCES AND NOTES
  9. Biographical Information
  10. Biographical Information
  11. Biographical Information
  12. Biographical Information
Thumbnail image of

Brent Vernon is an Associate Professor of Biomedical Engineering and the Graduate Program Chair in the School of Biological and Health Systems Engineering at Arizona State University in Tempe, AZ. He has degrees in Biomedical Engineering from Arizona State University (B.S.E., 1993) and the University of Utah (Ph.D., 2000). He obtained further training as a Postdoctoral Research Associate at the ETH/University of Zurich. His research focus is injectable polymers for drug delivery.