Tanya Portnov, Tiberiu R. Shulimzon and Meital Zilberman ...meitalz/Articles/Review 8.pdf · 92 T....

Rev Chem Eng 2017; 33(1): 91–107

*Corresponding author: Meital Zilberman, Department of Biomedical Engineering, Faculty of Engineering, Tel Aviv University, Tel Aviv 69978, Israel; and Department of Materials Science and Engineering, Tel Aviv University, Tel Aviv 69978, Israel, e-mail: [email protected] Portnov: Department of Materials Science and Engineering, Tel Aviv University, Tel Aviv 69978, IsraelTiberiu R. Shulimzon: The Pulmonary Institute, Chaim Sheba Medical Center, Tel Hashomer 52656, Israel

Tanya Portnov, Tiberiu R. Shulimzon and Meital Zilberman*

Injectable hydrogel-based scaffolds for tissue engineering applications

DOI 10.1515/revce-2015-0074Received December 14, 2015; accepted May 19, 2016; previously published online June 25, 2016

Abstract: Hydrogels are highly hydrated materials that may absorb from 10% to 20% up to hundreds of times their dry weight in water and are composed of three-dimensional hydrophilic polymeric networks that are sim-ilar to those in natural tissue. The structural integrity of hydrogels depends on cross-links formed between the pol-ymer chains. Hydrogels have been extensively explored as injectable cell delivery systems, owing to their high tissue-like water content, ability to mimic extracellular matrix, homogeneously encapsulated cells, efficient mass trans-fer, amenability to chemical and physical modifications, and minimally invasive delivery. A variety of naturally and synthetically derived materials have been used to form injectable hydrogels for tissue engineering applications. The current review article focuses on these biomaterials, on the design parameters of injectable scaffolds, and on the in situ gelling of their hydrogel systems. The last sec-tion of this article describes specific examples of catheter-based delivery systems.

Keywords: alginate; gelatin; hydrogel; injectable scaf-folds; natural biomaterials.

1 IntroductionThe multidisciplinary field of tissue engineering has gained much attention in recent years, as it holds great promise for improving the health and quality of life of mil-lions of people by developing functional substitutes for damaged or diseased tissues.

The most commonly used strategy in tissue engineer-ing is the seeding of cells into artificial, interconnected porous structures, called scaffolds, capable of supporting three-dimensional tissue formation. Cells can either be seeded on top of the scaffold in vitro, before implantation, or the patient’s own cells that infiltrate the implanted scaffold in vivo (Hou et al. 2004). Scaffolds may also be used as vehicles for the controlled and targeted release of therapeutic agents such as drugs and growth factors (Kretlow et al. 2007, Li et al. 2012).

The basic types of biomaterials used for the fabrica-tion of scaffolds can be broadly classified as follows:1. Synthetic polymers, such as the α-hydroxy acid poly-

mer family [e.g. poly(lactic-co-glycolic acid) (PLGA)] and polyanhydrides;

2. Natural polymers – complex sugars, such as hyalu-ronic acid (HA) and chitosan; and

3. Inorganic materials – bioactive ceramics, such as hydroxyapatite and tricalcium phosphate (Kohane and Langer 2008).

The role of scaffolds in tissue regeneration and repair is critical, designed to perform some or all of the following functions: (i) promote cell-biomaterial interactions, cell adhesion, and extracellular matrix (ECM) deposition; (ii) permit the sufficient transport of gases, nutrients, and regulatory factors to allow cell survival, proliferation, and differentiation; (iii) biodegradable at a controllable rate that approximates the rate of tissue regeneration under the culture conditions of interest; and (iv) provoke a minimal degree of inflammation or toxicity in vivo (Bra-hatheeswaran et al. 2011). These key scaffold character-istics can be tailored to the application by the careful selection of the polymers, additional scaffold compo-nents, and the fabrication technique.

Some common fabrication techniques include solvent casting and particulate leaching, gas foaming, thermally induced phase separation, freeze drying, and nanofiber electrospinning (Harris et al. 1998, Hutmacher 2001, Pham et al. 2006, Brahatheeswaran et al. 2011). These methods give rise to constructs in the form of meshes, fibers sponges, and foams, which are typically formed outside the body and must then be surgically implanted. To achieve the desired shape maximally matching the

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92 T. Portnov et al.: Injectable hydrogel-based scaffolds

defect, computer-generated molds that are based on clini-cal imaging techniques must be used. Such molds are reliant upon the accuracy of the images, increase cost, and may delay patient treatment (Moglia et al. 2013).

Recently, many tissue engineering designs using injectable, in situ forming systems have been reported for a broad spectrum of applications, including bone repair (Tran et al. 2011, Amini and Lakshmi 2012) and cardiac (Johnson and Christman 2013, Radhakrishnan et al. 2014) and cartilage (Balakrishnan and Banerjee 2011) tissue regeneration. In general, injectable scaffold formulations consist of cells and/or growth factors and solidifiable pre-cursors, which before administration may be in the form of solution, paste, microparticles or nanoparticles, beads, or thread-like materials, able to flow through a small gauge-needle or a catheter. Following injection and solidi-fication, the in situ forming matrix provides a temporary three-dimensional template on which the cells can adhere to form a functional new tissue.

Unlike tissue engineering approaches that use pre-fabricated scaffolds, injectable systems have drawn great interest within the field as a unique therapeutic method for difficult to reach areas of the body, showing the ability to conform to any shape irrespective of the defect geom-etry. From a clinical perspective, the minimally invasive procedure of injection may dramatically reduce patient discomfort, risk of infection, scar formation, hospitaliza-tion time, and treatment cost (Hou et al. 2004). Further-more, owing to the nature of these systems, cells and bioactive molecules can be easily incorporated in the scaf-fold solution either by mixing before injection or by simul-taneously injecting the cells/bioactive molecules together with the scaffold, allowing for their homogeneous distri-bution within the scaffold matrix.

Injectable scaffolds may be ceramic based (usually calcium phosphate- and calcium sulfate-based cements for orthopedic and dental applications; Tran et al. 2011, Rahman et al. 2012), formulated as polymeric micro-spheres (Cruz et al. 2008, Rungseevijitprapa and Bod-meier 2009, Voigt et al. 2012), or hydrogel based. Among the three, the latter systems have received the highest attention from the research community and were there-fore chosen to be the focus of the current review paper. The paper outlines the major requirements and chal-lenges faced in developing injectable hydrogel-based scaf-folds for tissue engineering. It covers the most commonly used natural and synthetic biomaterials and the various gelation mechanisms used to transform them into in situ gelling systems while highlighting their tissue engineer-ing applications. In addition, some catheter-based hydro-gel systems are discussed in brief. Finally, the challenges

and perspectives about the future of injectable hydrogels are given.

2 Design parameters of injectable scaffolds

The requirements of a scaffold are numerous and often depend on the application and on the environment into which the scaffold is ought to be injected. However, there are certain properties a material must fulfill to be consid-ered for tissue engineering use as an injectable scaffold. These parameters include both classical physical param-eters (e.g. mechanics and degradation) and biological performance parameters (e.g. cell adhesion). Other highly desirable features concerning the scaffold processing are ease of fabrication and scalability for cost-effective indus-trial production.

2.1 Mechanical properties

Once the scaffold is produced and placed, the formation of tissues with desirable properties relies on scaffold material mechanical properties on both the macroscopic and micro-scopic levels. Macroscopically, the scaffold must withstand biomechanical load and provide temporary support to the cells. On the microscopic level, evidence suggests that cell growth, differentiation, and ultimately tissue formation are dependent on the mechanical input to the cells at the implantation site (Tran et al. 2011, Mazaki et al. 2014).

The adequate mechanical performance of a scaffold depends on specifying, characterizing, and controlling the material mechanical properties, including elasticity, compressibility, viscoelastic behavior, tensile strength, and failure strain (Drury and Mooney 2003). For hydro-gels, these properties are affected by the cross-linking density, which is dictated by the rigidity of polymer chains, cross-linker characteristics, gelling conditions (e.g. tem-perature and pH), and swelling as a result of hydrophilic/hydrophobic balance (Lee et al. 2001, Wu et al. 2014). The degree of porosity will also have a substantial effect on the mechanical properties, with the stiffness of the scaffold decreasing as porosity increases.

2.2 Porosity

The porosity (i.e. void fraction) of the scaffold is another important design requirement, which plays a critical role



T. Portnov et al.: Injectable hydrogel-based scaffolds 93

in the outcome of a tissue-engineered graft. Cells seeded inside the scaffold rely heavily on the void spaces within the construct for cellular in-growth, vascularization, and the exchange of nutrients and waste products (Kretlow et al. 2007). In addition, the extent of ECM secretion also increases by increasing the pore size (Annabi et al. 2010). Thus, the injected graft should have a highly organized, porous, and interconnected structure. Pores may be formed in hydrogels by phase separation during synthe-sis or they may exist as smaller pores within the network (Hoffman 2012). A typical porosity of 90%, as well as a pore diameter of at least 100 μm, is known to be compul-sory for cell penetration and a proper vascularization of the ingrown tissue (Rezwana et al. 2006).

2.3 Biodegradation

Biodegradability is an essential parameter for many inject-able biomaterials, whether the gels are originated from natural resources or are synthetically created. Ideally, the degradation rate of a scaffold should match the rate of new tissue development, and this time is dependent on the tissue type to be engineered. Hydrogel degradation can be caused by hydrolysis, enzymatic cleavage, and/or dissolution. Synthetic hydrogels are mostly degraded through hydrolysis. The degradation rate and mechanical properties of cross-linked gels are typically coupled to each other, as both typically rise and fall together with the cross-linking density of the gels (Lee et al. 2001). In a study by Zhu and Ding (2006), for example, the in vitro degradation of hydrogels with different lengths and types of oligoesters was examined gravimetrically. Hydrogels made from the shorter oligoester blocks presented relatively slow degra-dation because of the smaller probability of hydrolysis of oligoester blocks. The degradation of hydrogels containing oligo(ε-caprolactone) was slower than that of those con-taining oligolactide, and this was attributed to the more hydrophobic nature of oligo(ε-caprolactone) versus that of oligolactide. The network cross-linking degree, which was controlled through the concentration of the accelera-tor, was also shown to affect biodegradability. The rate of enzymatic degradation will depend both on the number of cleavage sites in the polymer and on the amount of availa-ble enzymes in the scaffold environment (Mann et al. 2001).

2.4 Injectability

To guarantee injectability, the system should be in a sol state before administration. The sol is desired to be of

sufficiently low viscosity to allow the use of a small gauge-needle, which alleviates the patient’s pain. The viscosity could be affected by several factors. In a study by Burdick and Anseth (2002), the initial macromonomer concentra-tion was shown to affect the overall cross-linking density of the network and consequently the material’s viscosity. Interestingly, osteoblast viability was observed to decrease with the increase in macromonomer concentration as well (Burdick and Anseth 2002). Other parameters affecting viscosity in cross-linked scaffolds include the composi-tion and concentration of the initiator and the molecular weight of the cross-linking agent. Injection forces were also found to depend on injection speed, needle size, and injection site (Rungseevijitprapa and Bodmeier 2009).

2.5 In situ solidification parameters

The precursors of injectable scaffolds should undergo a mild solidification process preferably under, or close to, physi-ological conditions to keep high cell viability and molecular bioactivity as well as to avoid damage to the surrounding tissues. Toxic organic solvents or harsh processing condi-tions, such as high temperature, should be avoided. Ideally, the solvent used should be physiological saline, cell culture medium, or a biologically acceptable organic solvent (e.g. N-methyl-2-pyrrolidone or dimethyl sulfoxide). The temper-ature or pH at the site of implantation should not be signifi-cantly altered during the solidification process.

In situations where there is no defined boundary to the site of delivery, the scaffold must also be cohesive to enable accurate positioning at the required location. Con-sequently, the solidification time must be balanced and tuned according to the route and site of administration. It should not be very long, otherwise, if injected into the body, the biomaterial will presumably diffuse to the surround-ing tissue ahead of hydrogel formation. On the contrary, the gelation should not be very short either. Otherwise, the surgeons may not have sufficient time to perform the injection. A reasonable hydrogel gelation time, when administered by simple injection, is approximately 10 min (Zhu and Ding 2006, Li et al. 2012). However, intramyocar-dial injection procedures, using a catheter technique (as further elaborated in Section 5), for instance, can some-times be more than 1 h long, requiring the biomaterial to remain liquid throughout the entire injection procedure.

2.6 Biocompatibility and cytotoxicity

Injectable systems must be biocompatible and nontoxic to both cells and tissues. Naturally derived polymers




frequently demonstrate adequate biocompatibility, whereas synthetic polymers may elicit significant nega-tive responses from the body. Therefore, one may have some restrictions when preparing hydrogels from syn-thetic polymers for these applications (Lee and Mooney 2001). Nevertheless, in all cases, the cytotoxicity of scaffold constituents including macromonomers, cross-linking agents, functional groups, and initiators should be examined, as these components may be cytotoxic to surrounding cells during the cross-linking reaction. In addition, the cytotoxicity of any substances leached from the cross-linked polymer should be determined, as unre-acted residues and byproducts from the polymerization reaction released from the scaffold may prove detrimen-tal to cell viability. For biodegradable scaffolds, the cyto-toxicity of the degradation products should also be tested (Shin et al. 2003). The injectable precursors should also be able to undergo sterilization before injection, with the sterilization process having no significant impact on the chemical properties of the resulting scaffold.

2.7 Cellular response

The materials used in tissue engineering applications should mimic the ECM to provide mechanical support and regulate cell behavior including cell anchorage, segrega-tion, communication, and differentiation.

Unfortunately, with the exception of collagen, which is a natural ECM protein, most cells do not have recep-tors to hydrogel-forming polymers and thus cannot adhere. Furthermore, because of the hydrophilic nature of hydrogels, ECM proteins such as laminin, fibronectin, and vitronectin typically do not readily adsorb to the gel surface (Drury and Mooney 2003). To promote cell adhe-sion nevertheless, the commonly used strategy is to cova-lently couple an entire ECM protein or peptide sequences, capable of binding to cellular receptors, to the polymer. The most common peptide used in this approach is the amino acid sequence arginine-glycine-aspartic acid (abbreviated as Arg-Gly-Asp or, more commonly, RGD) derived from numerous ECM proteins (Alsberg et al. 2001, Mann et al. 2001, Burdick and Anseth 2002). In the case of growth factor incorporation, the injectable scaffold acts as a reservoir that releases the molecules at the repair site for the length of time necessary to create an environment con-ducive to tissue regeneration (Whitaker et al. 2001). Due to the short half-lives of growth factors, an important con-sideration is how to retain their bioactivity and effectively release them to the site with optimal dosage and timing. Alternatively, growth factors can be covalently attached to

polymer scaffolds to promote cell migration or production (Suzuki et al. 2000).

3 Materials for injectable scaffoldsInjectable hydrogels can be created by natural, syn-thetic, or their hybrid biomaterials. Based on the material origin, hydrogels can be classified into three major types: natural, synthetic and synthetic/natural hybrid hydro-gels. This section describes the properties of the natural and synthetic hydrogel forming polymers used for design-ing injectable scaffolds.

While synthetic biomaterials are appealing for tissue engineering because their chemical and physical prop-erties (e.g. block structures, molecular weights, and degradable linkages) are typically more controllable and reproducible than those of natural materials, the latter are the primary focus for injectable materials, owing to their biocompatibility, inherent biodegradability, and critical biological functions.

3.1 Naturally derived materials

Representative members include collagen, gelatin, fibrin, alginate, chitosan, HA, chondroitin sulfate, and agarose.

3.1.1 Collagen

Collagen is the main protein component of most extracel-lular environments in the human body. Furthermore, its superior biocompatibility, bioactivity, and biodegradability have made it one of the most extensively investigated bio-material scaffolds for tissue engineering (Lee and Mooney 2001). To this day, 29 distinct collagen types have been characterized, all displaying a typical helical structure with a defined pattern of amino acids, when a type I collagen is currently the gold standard in the field of tissue engineer-ing (Parenteau-Bareil et al. 2010). Collagen-based biomate-rials can originate from two fundamental techniques. The first one is a decellularized collagen matrix preserving the original tissue shape and ECM structure, whereas the other relies on extraction from biological tissues (e.g. bovine and porcine skin), purification, and polymerization of colla-gen and its diverse components to form a functional scaf-fold (Parenteau-Bareil et al. 2010). Both techniques can be submitted to various cross-linking methods, as detailed in Table 1, to enhance their mechanical and enzymatic resist-ance properties for implantation purposes.




3.1.2 Gelatin

Gelatin, which is an irreversibly hydrolyzed form of col-lagen, contains inherent peptide sequences that facilitate cell adhesion and enzymatic degradation (Koshy et al. 2014). Additionally, its low cost, lack of immunogenicity, and safety record in medicine (hemostatic agent, blood volume expander, etc.) makes gelatin an attractive inject-able biomaterial. Gelatin hydrogels can be obtained via a simple thermally induced gelation (cooling an aqueous gelatin solutions to < 30°C) or chemical cross-linking. In spite of their advantages, studies have found that scaf-folds composed of gelatin alone suffer from mechanical weakness. Hence, gelatin is often used in combination with other polymers such as chitosan, alginate, and others, in which its role is to enhance scaffold biocompat-ibility (Pok et al. 2013, Balakrishnan et al. 2014, Mazaki et al. 2014).

3.1.3 Fibrin

Fibrin, another commonly used fibrillar protein, is involved in the coagulation cascade (Sierra 1993), wound

healing responses, and promotion of angiogenesis (Naito et al. 2000). Fibrin glue (also called fibrin sealant) is a commercially available two-component product consist-ing of the fibrin precursor (fibrinogen) and the activating enzyme (thrombin), which self-assemble upon contact to create a fibrin gel (Christman et al. 2004). The remarkable advantages of the fibrin gel are its bioactivity, elasticity (can resist stretching to more than five times its resting length without breakage), soft compliance at small strains, and impressive stiffening to resist larger defor-mations (Janmey et al. 2009). Furthermore, the fibrino-gen can be obtained autologously (isolated from human plasma), avoiding the potential risks of foreign body reac-tion and infections (Zhao et al. 2008). The main disad-vantage of fibrin is its increasing instability and solubility over time in vitro and in vivo due to fibrinolysis. However, it has been reported that, by varying the fibrin param-eters, such as fibrinogen concentration, thrombin con-centration, and ionic strength, it is possible to generate gels with different appearance, mechanical properties, and stability (Balakrishnan and Banerjee 2011). Recently, some impressive and significant progress has been made using the fibrin gel as an injectable scaffold in cardiol-ogy (Christman et al. 2004, Ryua et al. 2005) and cartilage regeneration (Peretti et al. 2006).

Table 1: Injectable hydrogel materials and their in situ gelation mechanisms.

Hydrogels Polymer Solidification mechanism References

Synthetic hydrogels

PEG-based/PEO Free radical/chemical/thermal/Michael-type addition cross-linking

Liu et al. 2010, Nguyen and Lee 2010

PVA Freeze-thaw/free radical/chemical cross-linking

Ossipov et al. 2008, Samavedi et al. 2014

PPF or OPF Free radical/chemical cross-linking Temenoff et al. 2003, Tan and Marra 2010PEO-PPO-PEO, PLGA-PEG-PLGA, PEG-PLLA-PEG

Thermal cross-linking Cao et al. 1998, Cortiella et al. 2006

pNIPAAm Thermal cross-linking Cho et al. 2004, Liu et al. 2004, Ohya and Matsuda 2005, Chen and Cheng 2008

Peptides Self-assembly/ionic/pH-responsive Guvendiren et al. 2012, Li et al. 2012Natural hydrogels

Collagen Free radical/chemical/thermal cross-linking

Ohan et al. 2002, Suuronen et al. 2006, Wu et al. 2007, Parenteau-Bareil et al. 2010

Fibrin Chemical/thermal gelation Christman et al. 2004, Peretti et al. 2006, Zhao et al. 2008

Gelatin Chemical/Michael-type addition/thermal cross-linking

Nickerson et al. 2006, Li et al. 2012

Alginate Free radical/chemical/ionic cross-linking

Smeds and Grinstaff 2000, Lee and Mooney 2012, Li et al. 2012, Balakrishnan et al. 2014

Chitosan Free radical/chemical/Michael-type addition/thermal/pH-induced cross-linking

Chiu et al. 2009, Li et al. 2012, Mekhail and Tabrizian 2014

HA Free radical/chemical/Michael-type addition/thermal cross-linking

Leach et al. 2004, Yeo and Kohane 2008, Jin et al. 2010, Tan and Marra 2010

Chondroitin sulfate Chemical cross-linking Strehin et al. 2010Agarose Thermal cross-linking Varoni et al. 2012




3.1.4 Alginate

Alginate, also known as alginic acid, is an anionic, unbranched polysaccharide isolated from algae or bac-terial biofilms. Alginate is composed of (1,4)-linked β-D-mannuronate (M) and α-L-guluronate (G) sugar monomer blocks in varying proportions and sequen-tial arrangements, with the exact M and G composition being dependent on the algae or bacteria source. Alginate polymers have a high affinity for divalent cations (in the order Mg2+ < < Ca2+ < Sr2+ < Ba2+) and can form a cross-linked network when these divalent cations associate with the G blocks to form cross-links between the polymer chains (as further elaborated in Section 4.2.3). Alginate has been established as a biocompatible and nonimmunogenic polymer. Despite its advantageous features, alginate may not be an ideal candidate for tissue engineering because it does not specifically degrade (Lee and Mooney 2012). One approach to make alginate degradable in physiologi-cal conditions (depending on the pH and temperature of the solution) includes the oxidation of alginate chains to a low extent (~5%) using sodium periodate (Bouhadir et al. 2001). Another potential limitation in using alginate gels in tissue engineering is the lack of cellular interaction. To enhance cell ligand-specific binding properties, alginate has been covalently coupled with lectin and RGD (Lee and Mooney 2001, Balakrishnan and Banerjee 2011, Depan and Misra 2013).

3.1.5 HA or hyaluronan

HA or hyaluronan is a glycosaminoglycan that is widely distributed throughout the ECM of all connective tissues, where it plays an important role in lubrication, cell dif-ferentiation, and cell growth. HA offers many advantages as a tissue scaffold due to its viscous properties, ability to retain water, biodegradability, biocompatibility, and bioresorbability. Furthermore, HA contains functional groups (carboxylic acids and alcohols) along its back-bone that can be used to introduce functional domains or to form a hydrogel by cross-linking (Collins and Birkinshaw 2013).

3.1.6 Chitosan

Chitosan is a deacetylated derivative of chitin, a poly-saccharide found in the shells of crustaceans. The dea-cetylation of chitin produces D-glucosamine units that are randomly distributed within the polymeric chain

containing N-acetyl-D-glucosamine monomers. Due to its cationic nature, chitosan dissolves in slightly acidic solutions. Chitosan has been one of the most widely investigated polymers in the field of tissue engineering, and investigators have been developing a multitude of formulations that can undergo gelation in vivo (Mekhail and Tabrizian 2014). Its extensive use is attributed to its many desirable physicochemical characteristics such as biocompatibility, biodegradability, mechanical strength, hydrophilicity, good adhesion, and versatility in fabri-cation/modification (Pillai et al. 2009). Moreover, it is readily available and inexpensive. Chitosan lacks biologi-cal adhesion molecules that can promote cell attachment. Instead, cell attachment on chitosan occurs via two mech-anisms: (1) electrostatic attractions between the anionic cell membranes of most cell types and the slightly cationic nature of chitosan at physiological pH and (2) the adsorp-tion of serum proteins on the surface of chitosan, which in turn can be recognized by integrins and promote cell attachment (Chen et al. 2012). Many researchers attempted to improve attachment by blending or chemically modify-ing chitosan with other components. The incorporation of HA, for example, in the cross-linked chitosan network was shown to significantly increase cell proliferation and car-tilaginous ECM production by encapsulated chondrocytes (Park et al. 2013).

3.2 Synthetic materials

Examples of synthetic materials discussed hereunder include poly(ethylene glycol) (PEG), polyvinyl alcohol (PVA), poly(propylene fumarate) (PPF), oligo(PEG fuma-rate) (OPF), poly(N-isopropylacrylamide) (pNIPAAm), and poloxamers.

Cross-linked PEG and the chemically similar poly(ethylene oxide) (PEO) hydrogels, both hydrophilic, are widely considered biocompatible and biodegradable and are Food and Drug Administration (FDA) approved for several medical applications (e.g. sealants such as Dura-Seal™ for application over sutures in spinal surgery and ReSure®, a sustained ophthalmic drug delivery system). They are especially appealing due to their high water-absorbing capacity, which mimics native tissues (Liu et al. 2010). Biodegradable PEG hydrogels can be obtained via copolymerization with degradable polymers such as poly(lactic acid) (PLA), polyglycolic acid (PGA), and poly-caprolactone (PCL; Nguyen and Lee 2010). Furthermore, many naturally occurring biopolymers such as HA are also generally examined in combination with PEG hydro-gels (Leach et al. 2004).




PVA is another hydrophilic biocompatible and bio-degradable polymer that gels spontaneously in aqueous media. Improved stability has been achieved by freeze thawing and covalent cross-linking. The degradation rate of PVA hydrogels can be controlled by varying polymer molecular weight, concentration, and cross-linking density (Samavedi et al. 2014). By copolymerization and blending with natural and synthetic polymers, PVA has been used to create a vast array of hydrogels with diverse characteris-tics (Ossipov et al. 2008, Samavedi et al. 2014).

It is important to note that several hydrophobic syn-thetic polymers have also been used as injectable in situ gelling scaffolds. One such example is of a biodegradable and biocompatible PLGA scaffold. PLGA is dissolved in a biocompatible water-miscible solvent. When the polymer solution is injected into an aqueous environment, the solvent diffuses into the surrounding aqueous environ-ment, causing PLGA to harden. A porogen and a small amount of diH2O are mixed into the PLGA solution to allow for the formation of a porous structure throughout the resulting scaffold upon injection (Krebs et al. 2009). Another biodegradable and biocompatible polymer used as biomaterial is PCL. It is a semicrystalline, linear, ali-phatic polyester and has high tensile strength and unique elastic properties (Pok et al. 2013). However, the biodeg-radation of PCL occurs at a slow rate in vivo ( > 6 months), limiting its utilization as tissue engineering scaffold. Recently, researchers reported the development of an injectable PCL-based gel blended with a biodegradable natural polymer, sebacic acid, which degrades faster than pure PCL without loss of mechanical resistance (Salgado et al. 2012).

Certain peptide sequences self-assemble into nanofi-brillar hydrogel networks via charge interactions, hydro-gen bonding, hydrophobic interactions, π-stacking, or stereointeractions between the specific amino acid building blocks (Li et al. 2012). Peptide hydrogels enable the formation of scaffold materials that mechanically resemble the native ECM and can be formed at the site of implantation under physiological conditions, which are easily degraded by the body, and are biocompatible toward many cells (Li et al. 2012).

4 In situ gelling of hydrogel systems

Injectable hydrogels with potential applications in tissue engineering can be classified into physical and chemi-cal gels according to their cross-linking mechanisms of

solidification. For chemically cross-linked gels, network formation is achieved by covalently binding two or more monomers, whereas physical gelation arises from either polymer chain entanglements or physical interactions such as hydrogen bonds and ionic or hydrophobic interac-tions, in response to environmental stimuli, for example, temperature, pH, stress (Guvendiren et al. 2012), and ions.

Chemically cross-linked networks provide higher cross-linking densities to the polymer network and allow for the fabrication of scaffolds with enhanced mechani-cal properties compared to physically cross-linked net-works. However, the toxicity of the chemical cross-linking agents used may adversely affect cell behavior and the incorporated bioactive molecules and therefore should be carefully selected. On the contrary, the physical gelation of the network may avoid the use of cross-linking agents and organic solvents but shows limited physical perfor-mance (Tran et al. 2011) and relative instability (syneresis of the hydrogel), especially under the environment of a body medium, and the biodegradation rates are difficult to control well in vitro and in vivo (Zhu and Ding 2006). Recently, an attractive approach to combine physical and chemical gelation has been proposed by Boere et al. (2014). The researchers developed a dual hardening in situ cross-linkable hydrogel formed by temperature-induced physical cross-linking and chemoselective cross-linking by native chemical ligation without the addition of a catalyst. Combining thermogelation with native chemi-cal ligation allows the formation of an immediate physi-cal network that can be strengthened in time by chemical cross-links.

4.1 Chemically cross-linked hydrogel systems

4.1.1 Free radical polymerization

The polymerization operates through a free radical mech-anism triggered by redox or a photoradical initiator. Fol-lowing its generation, the initiating free radical reacts with the functional groups of the macromers, thereby adding them to the growing polymer chain, and a gel is formed. The solidification process is determined by a number of factors including the type and concentration of the initia-tor and precursors, intensity of visible or UV light, nature of the solvent system, and temperature. One of the draw-backs of radical polymerization systems is the generally exothermic nature of their solidification reactions, which may release heat that can lead to the necrosis of surround-ing tissues and harm encapsulated cells.




Examples of photopolymerizable macromers for tissue engineering applications, as summarized in review articles by Nguyen and West (2002) and later on by Di and Yoshihiro (2014), include PEG acrylate derivatives, PEG methacrylate derivatives, PVA derivatives, modified polysaccharides such as HA derivatives, and dextran methacrylate.

In particular, PEG-based photocurable hydrogels have been investigated for use in encapsulating cell applications, owing to their good biocompatibility, hydro-philicity, and ease of modification. Burdick and Anseth (2002), for example, have synthesized PEG-based hydro-gels as matrices for the photoencapsulation of osteo-blasts for bone regeneration. The majority of osteoblasts were observed to have survived the photoencapsulation process when PEG hydrogels were formed with 10 wt% PEG but exhibited a decrease in viability with the increase in macromer concentration, attributable to the overall cross-linking density of the network and the material’s viscosity. Similarly, Bahney et al. (2011) defined in their research nontoxic conditions for photoencapsulation of human mesenchymal stem cells (hMSC) in PEG diacrylate (PEGDA) scaffolds using a visible-light photoinitiator system.

A photocurable hydrogel composed of PEGDA covalently conjugated to a denatured fibrinogen is cur-rently commercially available (CE marked) under the name GelrinC™ (developed by Regentis Biomaterials). GelrinC™ is acellular injectable implant that degrades over the course of 6–12 months in synchronization with the growth of functional cartilage tissue.

Researchers (Wu et al. 2014) are also developing injectable, photocurable, biodegradable hydrogels based on methacrylated PEG-co-poly(glycerol sebacate) copoly-mers (PEGS-M). Alterations in the degree of methacryla-tion of PEGS-M hydrogels were shown to yield a tunable range of mechanical properties, swelling properties, and biodegradation rate in vitro. Bone marrow-derived mes-enchymal stem cells were conveniently encapsulated in PEGS-M hydrogels in situ by photo-cross-linking.

In another study, transdermal photopolymeriza-tion was used to convert poly(methylene oxide)-dimeth-crylate (methacrylate being the polymerizable group) and poly(methylene oxide) polymer suspension containing chondrocytes to a solid hydrogel, which was subcutane-ously injected to mice. Chondrocytes survived implan-tation and polymerization and formed neocartilage (Elisseeff et al. 1999).

Injectable hydrogel based on chitosan derivative/PEG dimethacrylate/N,N-dimethylacrylamide was also prepared by photopolymerization. The hydrogel showed

excellent mechanical behavior, good in vitro biocompat-ibility, and no cytotoxicity toward growth of human bone sarcoma cells, indicating its potential as a bone tissue engineering matrix (Ma et al. 2010).

Nonetheless, photopolymerization cannot be carried out uniformly in a large or thick system, especially in many clinical applications in which the light penetration depth is quite limited and light distribution is nonhomo-geneous. Therefore, photoinitiated systems are often less favored over redox and thermally activated systems.

One such cytocompatible and water-soluble radical initiation systems is composed of N,N,N,N-tetramethyl-ene-diamine (TEMED) acting as an accelerator and ammo-nium perosulfate (APS) acting as the initiator. It has been tested with PEG/oligo (hydroxyl acid) copolymer. The temperature change of the obtained hydrogels during the process of the cross-linking reaction was not significant at a reasonable gelation rate (Zhu and Ding 2006). PPF and OPF are another examples of biodegradable hydrogels extensively developed for use in tissue engineering appli-cations. PPF and OPF macromers are composed of bio-compatible blocks such as PEG and fumaric acid, which can be cross-linked through the unsaturated C = C double bond in the fumarate group and hydrolytically degraded through its ester bonds. Besides being photo-cross-linka-ble, these hydrogels are formed by polymerization under the initiation of APS/TEMED redox system (Temenoff et al. 2003). Analyzed by in vitro cytotoxicity assay and in vivo implantation, the PPF and OPF hydrogels have shown minimal or negligible cytotoxicity and are histocompat-ible (Tan and Marra 2010).

Other water-soluble methacrylated polymers that were polymerized using the APS/TEMED initiation system include dextran, albumin, (hydroxyethyl) starch, polyas-partamide, PVA, HA (Hennink and Van Nostrum 2012), and chitosan (Hong et al. 2007).

4.1.2 Chemical cross-linking

To circumvent the limitations of free radical systems, some chemical cross-linking approaches free of initiators have also been suggested. According to this approach, soluble polymer chains are modified with a pair of molecules that have specific affinity to each other. When these modi-fied polymeric chains are injected simultaneously, they undergo rapid cross-linking to give rise to a covalent cross-linked network. In these types of reactions, the solidifica-tion process is determined by the strength of the affinity of the cross-linkers. Glutaraldehyde and carbodiimides, for example, are two of the most commonly used cross-linker




reagents. They have been extensively used to cross-link collagen and gelatin-based biomaterials (Suuronen et al. 2006, Wu et al. 2007).

In situ HA hydrogel was prepared by chemical cross-linking upon mixing of one HA derivative with a hydrazide moiety and another HA derivative with an aldehyde. The cross-linked matrix showed good biocompatibility in vitro and in vivo and has been used in the prevention of peri-toneal adhesions in rabbit experiments (Yeo and Kohane 2008).

Recent studies identified that genipin, a natural product extracted from the gardenia fruit, can be used (0.5–3.5 wt%) to cross-link functional amine groups present in natural tissues and biomaterials with very minimal cytotoxic effects compared to studies performed with glutaraldehyde (Jin et al. 2004). Genipin has been investigated to modulate mechanical stiffness of colla-gen, chitosan, and gelatin (Nickerson et al. 2006, Lei et al. 2014).

A potential disadvantage of chemically cross-linked hydrogel systems is that the two prepolymers may have to be kept separated until use (e.g. in a double-barreled syringe) and that the gelation process could occur within the delivery device (e.g. needle; Kohane and Langer 2008).

Another approach to prepare injectable hydrogels as tissue engineering matrices is the use of the Michael-type addition reaction to cross-link precursor macromers. Cross-links are formed via the addition reactions between nucleophiles (e.g. thiol groups) and electrophiles (e.g. vinyl/acrylate groups; Mather et al. 2006). Michael addi-tion reactions have been popular for hydrogel preparation due to their controlled reaction time, ability to form dif-ferent types of bonds, and relatively mild reactivity with biomolecules (Li et al. 2012).

Michael addition is also applied to prepare HA hydro-gels. For example, solutions of thiolated HA and PEG vinyl-sulfone (PEG-VS) were cross-linked via Michael addition under physiological conditions (Jin et al. 2010). Gelation times varied from 14 to < 1 min depending on the molecu-lar weight of thiol-modified HA and PEG-VS, content of free thiols, and total polymer concentration. Gel-cell matrices were degraded in approximately 3 weeks. The hydrogels had a good biocompatibility to encapsulated chondro-cytes. On day 21, glycosaminoglycans and collagen type II were found to have accumulated in the hydrogels, sug-gesting their high potential for cartilage tissue engineer-ing (Jin et al. 2010). A similar study was conducted with thiolated chitosan and PEGDA (Teng et al. 2010). In vitro cell studies indicated that these hydrogels had good com-patibility towards human diploid fibroblasts and A549 cells.

4.2 Physical gelation

4.2.1 Thermogelling systems

Thermosensitive hydrogels are especially attractive due to their spontaneous gelation, free of any requirement of extra chemical treatment, above the lower critical solu-tion temperature (LCST), which is designed to be below body temperature. Typical thermal gelling polymers include copolymers of (N-isopropylacrylamide), PEG-based amphiphilic block copolymers, gelatin, agarose, and cellulose. The gelation of these polymers is related to the balance between intermolecular forces and hydropho-bic section aggregation, when the molecular weight of the hydrophobic section determines the sol-to-gel phase tran-sition temperature. The conformational changes that take place at the critical solution temperature are reversible, and gels can return to solution after the thermal stimulus that caused their gelation is removed.

Chenite et al. (2000) have demonstrated that chitosan solution neutralized with β-glycerolphosphate (GP) form in situ gelling systems. Gelation temperature was found to inversely correlate with the chitosan degree of deacetyla-tion. The ideal deacetylation degree for tissue engineering applications was found to be 91%, which provided a gela-tion temperature close to 37°C. The chitosan/GP gel was shown to be capable of delivering an osteogenic mixture of transforming growth factor-β (TGF-β) family members and encapsulating chondrocytes in vitro for normal cartilage regeneration over 3 weeks. A thermosensitive chitosan/GP blend (BST-CarGel®) developed by Piramal Healthcare Canada is currently commercially available for the repair of articular cartilage via arthroscopy or mini-open surgical techniques (Shive et al. 2006). Chemical modifications of chitosan have also been explored to render it thermosen-sitive. Bhattarai et al. (2005) incorporated PEG into chi-tosan, achieving gelation in physiological pH values. The amount of grafted PEG was critical for the formation of a thermosensitive hydrogel. Grafting < 40% (w/w) PEG on the chitosan backbone did not render chitosan soluble at a neutral pH, whereas grafting PEG at weight ratios higher than 55% (w/w) increased hydrophilicity and weakened the hydrophobic interactions between chitosan chains, thus preventing gelation.

Poloxamers, a family of Tri-block biodegradable copol-ymers, such as PEO-poly(propylene oxide) (PPO)-PEO [commercially called Pluronic® (F-127)], PLGA-PEG-PLGA, PEG-poly-L-lactide (PLLA)-PEG, PCL-PEG-PCL, poly(ε-caprolactone-co-lactide) (PCLA)-PEG-PCLA, and PEG-PCL-PEG, have been widely studied by many researchers as thermogelling polymers by micelle formation (Nguyen




and Lee 2010). At low temperatures, above the critical micellization concentration (CMC), the amphiphilic block polymer chains assemble first into micelles and bridged micelles, and then the ordered packing of bridged micelles is triggered at a higher temperature and a macroscopic gel is formed (Tan and Marra 2010). Pluronic® (F-127) was tested as a scaffold material for lung tissue engineering (Cortiella et al. 2006). The Pluronic® (F-127)/cell constructs resulted in the development of tissue with less inflamma-tory reaction in vivo than that of PGA-based constructs. Pluronic® (F-127) was also employed in combination with chondrocytes as an injectable scaffold for the creation of autologous tissue-engineered cartilage in the shape of a human nipple (Cao et al. 1998). However, Poloxamer hydrogels are not considered as ideal implanted materi-als. The sol-gel transition of these block copolymers is not suitable for the injection of hydrogels into deep ana-tomical sites in the body due to premature gelation inside the conduit or catheter used to deliver the hydrogel (Bae et al. 2005). In addition, the aliphatic polyesters of PLA and PLGA tend to undergo rapid degradation of the block copolymer, producing acidic monomers such as lactic or glycolic acids, which are known to be harmful to bioactive proteins or cells. To circumvent these drawbacks, Shim et al. (2006) synthesized a pH-sensitive and thermosensi-tive block copolymer by adding pH-sensitive sulfamethaz-ine oligomers (SMOs) to either end of a PCLA-PEG-PCLA block copolymer. The sulfonamide-modified block copol-ymer maintained its sol phase for about 2 h at body temperature at pH 8.0 but rapidly formed a gel in physi-ological conditions (pH 7.4 and 37°C) within only 5 min. Moreover, the buffering effect of the SMOs substantially suppressed the degradation rate of the block copolymer in physiological conditions. This sulfonamide-modified block copolymer also presented enhanced biocompatibil-ity both in vitro and in vivo.

pNIPAAm is nonbiodegradable and exhibits transi-tion from a hydrated, expanded state at low temperature to a collapsed state (coil-to-globule phase transition) at about 32°C (its LCST) in pure water. Copolymerization of NIPAAm with a more hydrophilic monomer increases the overall hydrophilicity of the polymer, and the stronger polymer-water interactions lead to an increase in the LCST. Likewise, copolymerization with a more hydro-phobic monomer results in a lower LCST than pNIPAAm. Moreover, the phase transition temperature is influenced by the presence of salts and pH to a certain extent (Klouda and Mikos 2008). The proximity of the ~32°C volume phase transition temperature (of pNIPAAm hydrogels with physi-ological temperature (37°C) makes these hydrogels envi-ronmentally tunable and therefore attractive as injectables

for tissue engineering applications. Over the past decade, various copolymers of NIPAAm have been developed. A copolymer of NIPAAm and water-soluble chitosan was tested for the chondrogenic differentiation of hMSC (Cho et al. 2004). The hydrogel showed a stable gelation at 37°C, and differentiation of hMSC into chondrocytes was observed both in vitro and in vivo. This hydrogel was rec-ommended for treating vesicoureteral reflux via an endo-scopic single injection technique. Liu et al. (2004) have grafted methylcellulose (MC) with NIPAAm, combining the thermogelling properties of both materials. They found that the addition of MC to NIPAAm polymers enhances the mechanical strength of the hydrogel with no syneresis. A thermoreversible injectable hydrogel, HA-γ-(chitosan-γ-pNIPAAm) [HA-(chitosan-pNIPAAm)], was prepared by double grafting of HA and carboxylic acid-terminated pNIPAAm to a chitosan backbone by covalent binding (Chen and Cheng 2008). MSCs proliferated well in these gels and were induced to differentiate into osteoblasts after 7 days, suggesting that the hydrogel is suitable for osteo-genesis of MSCs and bone regeneration. Gelatin is another biopolymer with thermoreversible properties. It solidifies at temperatures below 25°C, but when the temperature is raised above approximately 30°C the gel becomes liquid again. As the opposite thermal behavior is desired for bio-medical applications, researchers have combined gelatin with other polymers that show thermal gelation close to body temperature. Ohya and Matsuda (2005) have grafted gelatin with NIPAAm in an effort to produce a thermore-sponsive ECM analogue. Aqueous solutions showed a sol-gel transition at physiological temperature when the weight ratio of pNIPAAm to gelatin chains was higher than 5.8. Smooth muscle cells were suspended in medium solu-tions of pNIPAAm/gelatin and subsequently incubated at 37°C. It was shown that a low hydrogel concentration (5%, w/v) and a high pNIPAAm to gelatin ratio supported the highest cell proliferation and ECM production. Ohya and Matsuda suggested that this was due to increased hydro-phobicity caused by higher pNIPAAm ratios, which would lead to the formation of large aggregates. As a result, a higher porosity with larger pore size occurs, which com-prises a favorable cell environment.

4.2.2 pH-responsive systems

Like thermosensitive hydrogels, pH-responsive hydrogels take advantage of the physiological environment to trigger gelation. Chiu et al. (2009) employed a hydrophobically modified chitosan (N-palmitoyl chitosan) to develop a pH-triggered hydrogel system that showed a rapid




nanostructure transformation within a narrow pH range (pH 6.5–7.0). The hydrogel was subcutaneously injected into a rat model by application of shear stress (during injection) and quickly self-healed after the removal of shear. This in situ hydrogel system was shown to be non-toxic and in vivo degradable within 6 weeks.

4.2.3 Ion-mediated gelation systems

Another group of stimulus-responsive hydrogels uses ionic cross-linking as the trigger for gel formation. The naturally occurring alginates are the most predominant members of the group, exhibiting the ability to cross-link via di- or tri-valent cations, usually through the use of calcium chlo-ride as the ionically cross-linking agent. The Ca2+ ions bind to the guluronate blocks of the alginate chains, as their structure allows a high degree of ion coordination, which then form junctions with the guluronate blocks of adja-cent polymer chains, resulting in gel formation. A major limitation of the ion-mediated gelation systems is that the ions could be exchanged with other ionic molecules in aqueous environments, resulting in an uncontrolled deterioration of the original properties of the hydrogel. Furthermore, calcium ions released from these gels can also be immunostimulatory (Chan and Mooney 2013). For this reason, covalent cross-linking with various types of molecules and different cross-linking densities has been attempted to precisely control the mechanical and/or swelling properties of alginate gels (Lee et al. 2000).

For ionically cross-linked alginate hydrogels, gela-tion rate is a critical factor in controlling gel uniformity and strength when using divalent cations. Slower gelation produces more uniform structures and greater mechanical integrity (Kuo and Ma 2001). One approach to slow and control gelation is to use a buffer containing phosphate (e.g. sodium hexametaphosphate), as phosphate groups in the buffer compete with carboxylate groups of alginate in the reaction with calcium ions and retard gelation. Calcium carbonate (CaCO3), due to its lower solubility, can also slow the gelation rate and widen the working time for alginate gels (Cho et al. 2009). The gelation tempera-ture also influences the gelation rate and the resultant mechanical properties of the gels. At lower temperatures, the reactivity of ionic cross-linkers (e.g. Ca2+) is reduced, and cross-linking becomes slower. The mechanical prop-erties of ionically cross-linked alginate gels can vary significantly depending on the chemical structure of algi-nate. For example, gels prepared from alginate with a high content of G residues exhibit higher stiffness than those with a low amount of G residues (Drury et al. 2004).

To summarize, Figure 1 is a schematic illustration of injectable in situ gelling hydrogel-based systems for tissue engineering. Table 1 summarizes the most com-monly investigated injectable hydrogel polymers, both synthetic and natural, with respect to their in situ gelation mechanisms.

5 Examples for catheter-based delivery of injectable hydrogels

In situ cross-linkable hydrogels are highly desirable clini-cally, as they can be introduced into the body via a mini-mally invasive manner by conventional needle-syringe injection, endoscopes, or catheters.

In cardiac therapy, many biomaterials including fibrin, collagen, Matrigel, and alginate, have demonstrated preserved or improved cardiac function when delivered through a syringe into the targeted region of small animal models (Dai et al. 2005, Huang et al. 2005, Landa et al. 2008). However, for a biomaterial to be clinically relevant for heart treatment in large animals as well as in humans, it must be translated to a catheter delivery. Current cardiac catheter technology includes intracoronary and transen-docardial delivery techniques, both tested on large animal models and humans (Singelyn 2010). In small animal models, intramyocardial delivery is usually performed via direct epicardial injection during an open-chest surgical procedure, as the more complex delivery techniques are difficult to scale down for a mouse, rat, or rabbit. Figure 2 presents the various injection delivery techniques. Intra-coronary injections, which involve the use of a catheter access to the coronary vessels to deliver treatment, use the pathological phenomenon of leaky vasculature to deliver the liquid form of the material into the tissue from the bloodstream with no direct puncturing of the tissue (Johnson and Christman 2013). Recently, intracoronary delivery of acellular calcium cross-linked alginate hydro-gels was found not only feasible but also safe and effective in preventing left ventricular (LV) remodeling early after myocardial infraction (MI) in swine (Leor et al. 2009) and in dog models (Sabbah et al. 2008). Moreover, a first-in-man pilot study has shown that the intracoronary deploy-ment of acellular calcium cross-linked alginate scaffold is well tolerated by patients (n = 27) surviving from a mod-erate to large MI. These results have promoted the initia-tion of a multicenter, randomized controlled trial aiming to confirm the safety and efficacy of this new approach in high-risk patients after ST-segment-elevation MI (Frey et al. 2014).




Figure 1: A schematic representation of injectable in situ gelling hydrogel-based scaffolds for tissue engineering (in this case, bone repair).An aqueous mixture of gel precursors (A–C) and bioactive agents (cells and growth factors) is administered using a syringe into the bone and gelates once inside the body. The gel precursors can gelate due to (A) free radical polymerization mechanism, triggered by a photoradi-cal initiator, (B) chemical cross-linking or Michael-type addition reaction, or upon (C) physical stimuli (temperature, ionic concentration, or pH). After being delivered, the injectable hydrogels form tissue constructs (scaffolds), providing local biological and mechanical cues that may enhance tissue regeneration, simultaneously to their degradation.

Figure 2: Injection delivery methods: (A) direct epicardial injection: injection through a standard needle and syringe into the myocardial tissue, (B) transcoronary injection: injection via catheter access to the coronary vessels, and (C) transendocardial injection: injection into the myocardial wall via catheter access into the lumen of the LV.

On the contrary, transendocardial delivery, the injec-tion into the myocardial wall via catheter access into the lumen of the LV, is considered to be the preferred method of catheter delivery, as it does not require access to the coronary vessels. Recently, Singelyn et al. (2012) demon-strated that a decellularized myocardial matrix can be

delivered to the myocardium via a percutaneous transen-docardial approach employing the Myostar® intramyo-cardial needle-injection catheter (Biosense Webster, a Johnson & Johnson company). The material increased endogenous cardiomyocytes in the infarct area and main-tained cardiac function without inducing arrhythmias in a rat MI model.

For injections into the heart using intracoronary or transendocardial delivery techniques, the following requirements must be met:1. Hydrogel components must be premixed and flow

freely through an extremely long and narrow lumen.2. The injected material must stay in liquid form, while

being held at 37°C, for potentially more than 1-h-long procedure and only form a gel once it enters the tissue. In the case of transendocardial delivery, the material must be capable of being injected multiple times at the site of injury.

3. For both transendocardial and intracoronary deliver-ies, the material must be hemocompatible, as some leakage into the bloodstream is known to occur.

In an attempt to develop new approaches to catheter inject-able materials, Grover et al. (2013) investigated oxime cross-linked PEG, HA-PEG, and alginate-PEG based systems. The investigators reported on a significant difference between




pH-dependent gelation rates of these systems in vitro and those observed in vivo. For example, subcutaneous injection of the PEG-only oxime system resulted in hydrogel forma-tion within 20 min, over a broad range of pH values (4–10.5), whereas in vitro gelation at 37°C occurred from 30 min to more than 2 days, over the same pH range. A similar phe-nomenon was observed when the material was injected into the myocardial tissue of rats. According to the investigators, this will allow PEG hydrogels cross-linked by oximes to be held at physiological temperature for extended periods of time (tunable by pH), which facilitates the delivery of these materials using a catheter while rapidly gelling in vivo. A dif-ference in gelation rates in vitro versus in vivo has been seen with other injectable hydrogel systems as well. For example, certain ECM-based hydrogels have been shown to be inca-pable of forming hydrogels in vitro yet rapidly gelling in vivo (Christman et al. 2011). This indicates that slow or no gela-tion in vitro does not necessarily relate to gelation in vivo, where the environment is more complex.

In another study by Bastings et al. (2014), the research-ers used the pH-responsive sol-to-gel behavior of a synthetic supramolecular hydrogel, consisting of ureido-pyrimidi-none (UPy) units coupled to PEG, to develop an injectable system suitable for catheter delivery. The UPy-hydrogel was liquid above a pH of 8.5, with a viscosity low enough to enable passage through a 1-m-long catheter, but gelled instantaneously at neutral pH. In addition, the UPy-hydro-gels showed a self-healing behavior, namely, liquid-like under large deformations, but recovered within minutes upon deformation removal, thus demonstrating potential for good adjustment to the injection site. The hydrogel was loaded with growth factors (for controlled release) and delivered via a catheter to a porcine model of myocardial infarction, where it exhibited a reduction of infarct scar.

A fibrinogen-fibronect in vitro nectin hydrogel (FFVH) was developed and optimized with respect to its mechani-cal and biological properties for a scaffold able to receive reparative cells in target sites of a murine lung (Ingenito et al. 2010). Following in vitro tests, FFVH was able to promote cell engraftment in a sheep lung model of emphy-sema. Cells labelled with a fluorescent dye (PKH-26) were detected at the treatment sites after 1 month. Tissue mass [assessed by computed tomography (CT) imaging] and lung perfusion (assessed by nuclear scintigraphy) were increased at emphysema test sites. FFVH scaffolds promote cell attachment, spreading, and ECM expression in vitro and apparent engraftment in vivo, with evidence of trophic effects on the surrounding tissue (Ingenito et al. 2010).

In summary, catheter delivery of hydrogel-based scaffolds is currently under early stages of development, mainly evolving for purposes of cardiac repair, using

materials such as alginate, ECM, and PEG, without any cellular therapies. A variety of well-designed and user-friendly catheters are commercially available on the market nowadays (Sherman et al. 2006) and are crucial for a quick translation to the clinic. However, this mode of delivery also provides unique challenges and design parameters for the biomaterial.

6 Concluding remarks and future perspectives

Injectable hydrogels formed by in situ chemical or physical gelation mechanisms have been drawing enormous atten-tion in the past decade. The advantages of such systems have become especially obvious for tissue engineering applications, as they would allow replacing traditional surgical procedures with minimally invasive methods. Although the basic materials, design parameters, and in situ gelling techniques are well established and described in the current review article, there remain a number of challenges and room for significant improvements to fit specific applications. Numerous studies have been imple-mented in small animals, mostly by injecting the hydrogel subcutaneously rather than in the target tissue of interest, and fewer in large animal models, and only several mate-rials have been translated to patients (mostly in the fields of cartilage and cardiac regeneration).

The main challenges still lie in obtaining appropriate chemical and morphological cell cues, good mechanical properties, and gelation and degradation kinetics. As dis-cussed, some advances are already under way. For example, novel methods to control material degradation have been engineered into many hydrogels, and cell adhesion ligands have been attached to these materials. Various methods of cross-linking have also been implemented both to enhance the material biocompatibility and to control the mechanical properties. Nonetheless, how to integrate the advantages of each of the various materials wisely has been and will be the key factor for the development of novel injectable and biodegradable hydrogel scaffolds for tissue engineering.

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