Injectable alginate hydrogels for cell delivery in ... - Semantic Scholar

Acta Biomaterialia 10 (2014) 1646–1662

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Review

Injectable alginate hydrogels for cell delivery in tissue engineering q Sílvia J. Bidarra a,⇑, Cristina C. Barrias a, Pedro L. Granja a,b,c a

INEB – Instituto de Engenharia Biomédica, Universidade do Porto, Rua do Campo Alegre, 823, 4150-180 Porto, Portugal FEUP – Faculdade de Engenharia da Universidade do Porto, Departamento de Engenharia Metalúrgica e de Materiais, Rua Dr. Roberto Frias, s/n, 4200-465 Porto, Portugal c ICBAS – Instituto de Ciências Biomédicas Abel Salazar, Universidade do Porto, Rua de Jorge Viterbo Ferreira, 228, 4050-313 Porto, Portugal b

a r t i c l e

i n f o

Article history: Available online 12 December 2013 Keywords: Tissue engineering Regeneration Cell delivery Injectable Alginate

a b s t r a c t Alginate hydrogels are extremely versatile and adaptable biomaterials, with great potential for use in biomedical applications. Their extracellular matrix-like features have been key factors for their choice as vehicles for cell delivery strategies aimed at tissue regeneration. A variety of strategies to decorate them with biofunctional moieties and to modulate their biophysical properties have been developed recently, which further allow their tailoring to the desired application. Additionally, their potential use as injectable materials offers several advantages over preformed scaffold-based approaches, namely: easy incorporation of therapeutic agents, such as cells, under mild conditions; minimally invasive local delivery; and high contourability, which is essential for filling in irregular defects. Alginate hydrogels have already been explored as cell delivery systems to enhance regeneration in different tissues and organs. Here, the in vitro and in vivo potential of injectable alginate hydrogels to deliver cells in a targeted fashion is reviewed. In each example, the selected crosslinking approach, the cell type, the target tissue and the main findings of the study are highlighted. Ó 2013 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.

1. Injectable cell delivery systems In many clinical scenarios, where normal physiological conditions or homeostasis are compromised, there is a need for tissue transplantation or implantation. The ideal paradigm in tissue engineering consists in introducing cells or tissues grafts, native to the injured area, to foster the regenerative process. In this context, cellbased therapeutic approaches can thus be considered a vital tool in regenerative medicine strategies. They rely on the successful delivery of living cells to the target location, where they can produce a desired therapeutic effect by paracrine delivery of biomolecules (growth factors, cytokines, hormones, etc.) or replace lost cells with donor cells that can integrate and regenerate the damaged tissues [1,2]. Cells used for cellular therapies can be previously manipulated to produce a missing substance, such as a specific protein that is absent in a metabolic disease [3]. To ensure that an adequate number of cells reach the target tissue, cell-based approaches have usually been based on the delivery of high-density single-cell suspensions to the site of injury through injection. However, such direct cell injection often has a poor outcome due to large and rapid loss of cell viability (thus requiring

q Part of the Special Issue on Biological Materials, edited by Professors Thomas H. Barker and Sarah C. Heilshorn ⇑ Corresponding author. Tel.: +351 226074900; fax: +351 226094567. E-mail addresses: [email protected] (S.J. Bidarra), [email protected] (C.C. Barrias), [email protected] (P.L. Granja).

high cell densities, which makes it technically complex and also extremely expensive), reduced engraftment of delivered cells and limited control over cell fate, in terms of both site-specificity (with cells eventually migrating and affecting other sites) and cell differentiation [4–6]. Therefore, more effective cell transplantation methods, capable of sustaining the survival of implanted cells while maintaining their function and enhancing their incorporation into the host, are mandatory. One strategy to achieve this goal relies on the delivery of transplanted cells via a temporary support made of biocompatible materials that can be further biochemically and physically modified to improve cell delivery [7,8]. This strategy will provide cell protection, prolonged retention at the injury site and a more physiological three-dimensional (3-D) environment. Moreover, many studies have pointed to the importance of strategies that promote cell–cell and cell–matrix interactions, which impact considerably on cell morphology, viability and function [5,9]. For instance, for anchorage-dependent cells, these interactions define cell shape and organization, which in turn will regulate cell behavior, namely survival, differentiation, proliferation and migration [10]. Hydrogels are candidates of choice to mediate cell delivery and accommodate cells in a 3-D microenvironment due to their natural similarities to the extracellular matrix (ECM). One paradigmatic example is alginate, a tunable and versatile natural injectable hydrogel with huge potential as an artificial 3-D cellular matrix, which has already been explored in a myriad of studies as an injectable cell delivery system for a broad variety of biomedical applications [11].

1742-7061/$ - see front matter Ó 2013 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved. http://dx.doi.org/10.1016/j.actbio.2013.12.006

S.J. Bidarra et al. / Acta Biomaterialia 10 (2014) 1646–1662

For clinical applications, cell delivery through injectable materials may be a desirable method, since these systems offer specific advantages over preformed scaffold-based approaches. On the one hand, they can be applied using minimally invasive techniques, improving patient compliance and comfort, and eventually leading to faster recovery and hence lower healthcare costs. On the other hand, they present several additional appealing features, namely: (i) easy incorporation of therapeutic agents, such as proteins or cells, and their subsequent localized delivery; (ii) simplicity of implantation by injection; (iii) high contourability, providing adaptable filling of defects with irregular shapes and sizes; and (iv) site specificity as well as confined delivery [12–22]. A key requisite of cell delivery vehicles is the maintenance of cell viability throughout the injection procedure. For instance, when flowing through a syringe needle, cells can experience three types of mechanical forces that can lead to cell disruption: (i) a pressure drop across the cell; (ii) shearing forces due to linear shear flow; and (iii) stretching forces due to extensional flow [23]. Therefore, an injectable matrix is also required to exhibit adequate mechanical properties to protect injected cells and thereby ensure their survival [23,24]. Injectable cell-based systems can be prepared with different configurations, depending on the type of application. Distinct and often conflicting terms exist to designate these systems. In the present review, we propose a classification (Fig. 1) in which systems have been divided into four main categories: (1) surface attachment with a preformed microcarrier, where cells are attached to the outside surface of microcarriers or even to the surface of inner pores; (2) microencapsulation, where cells are dispersed in a liquid contained within a polymeric membrane or capsule (mainly used for immunoisolation); (3) matrix entrapment, where cells are embedded in a hydrogel matrix that creates a 3-D environment, which can be formed in situ or ex situ (as in the case of microparticles); and (4) multicellular aggregates, where cells self-aggregate spontaneously or upon induction (due to matrix properties) to present some integrity, despite sometimes being ‘‘scaffold-free’’, and can be manipulated as ‘‘microparticles’’ due to their size. These multicellular aggregates can be further injected

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using a hydrogel-based vehicle [5]. Although alginate hydrogels have been employed in all of these types of systems, this review will mainly focus on applications belonging to category 3, thus the term ‘‘entrapment’’ will be used throughout the text. The main purpose of this review is to demonstrate the alginate´s potential as an adequate 3-D microenvironment for cell delivery, in which cells are kept in direct contact with this synthetic ECM and, after transplantation, might become incorporated in the host tissue and actively participate in the regeneration process. Therefore, categories 1, 2 and 4 above will be described no further in this review. In strategy 1, cells are actually in a 2-D rather than a true 3-D environment; strategy 2 aims at isolating cells from the host environment that will most likely not be incorporated in the host; and in strategy 4 the 3-D environment is a result of the natural cellular aggregation and is not necessarily provided by the matrix. The present review focuses mainly on strategies where cell-laden alginate-based systems were designed to be delivered in a minimally invasive way and, after injection, allow the transplanted cells to be integrated in the host’s damaged tissue and actively participate in the regeneration process.

2. Hydrogels in tissue engineering Hydrogels are 3-D hydrophilic, cross-linked polymeric networks capable of absorbing a significant amount of water or biological fluids [27]. Chain crosslinking can be achieved by chemical or physical methods, and the large variety of crosslinking reaction schemes make it possible to control the gelation kinetics and the subsequent hydrogel properties [28]. Hydrogels are compliant and permeable structures, mostly constituted of water, which resemble the native ECM, providing an ideal 3-D microenvironment for culturing cells. As a consequence, hydrogels have emerged as a valuable platform for examining the effects of ECM properties on cell behavior [9,29–33]. Additionally, they can be further modified to improve their mechanical and biochemical properties, to better mimic the native ECMs, via physicochemical modifications of the gel-forming polymers and/or crosslinking

Fig. 1. Cell immobilization strategies for injectable cell-based therapies. Several strategies and carrier materials can be used for cell immobilization. These can be adapted for use in injectable cell-based therapies. Here, the different approaches were divided into four major categories: (1) surface immobilization; (2) microencapsulation; (3) matrix entrapment; and (4) multicellular aggregation. The bottom images correspond to injectable alginate-based systems incorporating cells. (A) Bone marrow stromal cells cultured for 5 days on the surface of calcium titanium phosphate microspheres, under standard osteoinductive conditions. Cells were stained with Alexa-fluor 488-phalloidin (F-actin) and counterstained with propidium iodide (DNA). (B) Human mesenchymal stem cells (hMSCs) entrapped and cultured for 21 days inside 2 wt.% RGD–alginate microspheres in basal conditions. (C) hMSCs entrapped and cultured for 8 days inside 2 wt.% PVGLIG/RGD–alginate hydrogels under basal conditions. Cells were stained with Alexa-fluor 488-phalloidin (F-actin) and counterstained with DAPI (DNA). (D) Mouse mammary epithelial cell line (EpH-4) within 2 wt.% RGD–alginate after 7 days. Cells were stained for E-cadherin (red) and counterstained with DAPI (DNA). Adapted from Refs. [25,26].

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compounds [14,24,34–40]. Furthermore, many hydrogels can be formed under mild conditions, creating the adequate conditions for cytocompatible cell entrapment. In situ forming hydrogels present the added benefit of injectability, and have been widely studied as cell carriers for in vivo tissue engineering [41]. As previously pointed out, injectable hydrogels offer several advantages for cell entrapment and subsequent delivery, for which purpose they should ideally combine a number of requisites, such as: (i) be in a sufficiently fluid state during administration, like a low-viscosity solution, or undergo shear thinning before administration; (ii) gelation should start or be completed shortly after injection; (iii) be biocompatible and biodegradable, and their products bioresorbable; and (iv) fulfill specific requirements according to the application envisaged (e.g. celladhesive capability for entrapment of anchorage-dependent cells). Injectable hydrogels can be obtained from either natural or synthetic polymers [42,43]. Naturally derived polymers are frequently selected since these hydrogels either include components of the extracellular matrix (e.g. collagen, fribronectin, fribrinogen) or present a chemical structure similar to natural glycosaminoglycans (e.g. alginate, hyaluronic acid, chitosan), offering an intrinsic advantage over synthetic hydrogels [33]. Since they are derived from natural sources, many of them contain cellular binding domains, thus allowing cell adhesion and/or present soluble signaling factors to become intrinsically bioactive and capable of regulate cellular behavior. However, they may exhibit batch-to-batch variations and, in some cases, it is difficult to modulate their usually poor mechanical properties [33,44]. Additionally, natural hydrogels may be inherently immunogenic or give rise to some immunogenicity due to the presence of contaminates, like proteins and endotoxins. In contrast with natural polymers, synthetic polymers allow better control over and reproducibility of their mechanical properties and chemistry [45]. On the other hand, they may present a low degradation rate in physiological conditions and, in some cases, their preparation involves the use of toxic chemicals. Among these, polyethylene glycol (PEG)-based polymers are widely used as injectable hydrogels for cell delivery, and considerable efforts have been made in order to improve their features, including incorporation of cell adhesion ligands or biodegradable units susceptible to cell proteolytic activity, making them viable alternatives [38–40].

3. Alginate hydrogels Alginates are natural anionic biopolymers typically extracted from brown seaweeds. They are unbranched polysaccharides consisting of 1,4-linked b-D mannuronic acid (M) and a-L-guluronic acid (G) units, which are covalently linked together in different number and sequence distributions along the polymer chain, depending on the alginate source [46]. The functional properties of alginate depend on its monomer composition (M/G) ratio and sequence [47]. For example, MG blocks (MGMGMGM) form the most flexible chains and G blocks (GGGGGGG) form stiff chains. Moreover, alginate can be prepared with a wide range of molecular weights (typically 101–103 kDa) [24,47]. Alginates are already used in several clinical applications, e.g. for treatment of heartburn and acid reflux (GavisconÒ, BisodolÒ, Asilone™), as wound dressing materials (AlgicellÒ, AlgiSite M™, ComfeelÒ Plus, KaltostatÒ, SorbsanÒ and Tegagen™) and as an appetite suppressant for long-term weight loss [48–50]. Alginate is further being currently assessed in several clinical trials, namely as an agent for weight control, in the treatment of type I diabetes and as temporary acellular scaffolds to attenuate adverse cardiac remodeling [51]. Additionally, tissue engineering alginate products are commercially available for 3-D cell culture, such as AlgiMatrix

from Invitrogen and NOVATACH peptide-coupled alginates from FMC Biopolymers. Alginate is considered to be non-immunogenic and has shown great potential as a cell delivery vehicle [46,50,52–54]. In Fig. 2 several examples of alginate as a cell delivery system are depicted. It is possible to visualize that alginate matrices allow the outward migration of different cell types, in this case human umbilical vein endothelial cells (HUVECs) (Fig. 2A and B) and human mesenchymal stem cells (hMSCs) (Fig. 2C and D), in different scenarios in vitro (Fig. 2A), ex vivo (Fig. 2B) and in vivo (Fig. 2C and D). Cell migration from the matrix is a vital process when a cell delivery strategy is envisaged. As previously pointed out, an adequate vehicle must provide cell protection during injection. Alginate has been shown to have a shielding effect on cells during ejection from a syringe needle. Aguado et al. [23]tested 1 wt.% alginate with three different molecular weights on four different cell types (HUVECs, rat MSCs, human adipose stem cells and mouse neural progenitor cells), and have demonstrated the protective effect of alginate hydrogels with optimized mechanical properties (G0 30 Pa). The authors showed that HUVECs in phosphate-buffered saline (PBS) or in non-crosslinked alginate have a significantly lower cell viability compared to in crosslinked alginates. Moreover, for all the studied cells, there was a significantly higher viability in alginate with G0 30 Pa than in PBS or in alginate with a higher modulus. In an earlier study by Kong and colleagues [24], the authors showed that it is possible to maintain the alginate’s gel-forming ability by adjusting its molecular weight, while decreasing its viscosity and thus better preserving cell viability (cell viability 80% within 2 and 3.5 wt.% alginate hydrogels). To obtain different molecular weights, the alginate was irradiated, resulting in solutions with different viscosities and degradation rates [47]. Alginate presents a number of benefits for general use in biomedical applications, coupled with a set of unique advantages for cell encapsulation and entrapment. Alginate allows the formation of hydrogels (with water) in situ, and the gelling process can be carried out using non-toxic solvents and under physiological conditions (namely in terms of pH and temperature), which provide easy cell encapsulation and entrapment. These hydrogels possess a soft nature, making them physically similar to most native tissues. Additionally, alginate mechanical properties can be tuned in order to encompass a range of stiffnesses that cover a variety of tissues. For example, the compression modulus of alginate hydrogels can range from 1 to 1000 kPa and the shear modulus from 0.02 to 40 kPa [55]. Alginate mechanical properties can be controlled by changing different parameters, such as the polymer source, molecular weight, concentration and chemical modifications, and/or the type and density of the crosslinking [46]. Because they are transparent, alginate hydrogels allow the routine analysis of entrapped cells using standard microscopical techniques and they further enable easy cell recovery without cell damage. Regarding batch-to-batch variation and immunogenic response, both can be avoided through the use of highly purified and wellcharacterized alginates (concerning the molecular weight and the G/C ratio) that are currently commercially available. For instance, highly purified or ultrapure alginate contains low levels of residual endotoxin, lower than 100 EU g1, and has been shown not to induce any severe foreign body reaction when implanted into animals [53]. As major drawbacks for these applications, alginate is slowly biodegradable and is not cell-interactive. However, these can be easily overcome by a number of chemical and biochemical modifications [14,15,35,56]. Although alginate chains cannot be cleaved by mammalian enzymes, ionically crosslinked alginate hydrogels disintegrate progressively in vivo due to the exchange between monovalent cations, such as Na+, present in the surrounding milieu and the


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Fig. 2. Outward cell migration from alginate hydrogels. In vitro: HUVECs migrating out from an alginate–RGD hydrogel disk and (A) adhering to the bottom of a TCPS culture plate. Ex vivo: (B) forming tubular-like structures that sprout into Matrigel. In vivo: an hMSC-laden alginate–RGD–PVGLIG hydrogel after subcutaneous implantation in an severe combined immunodeficiency (SCID) mice, showing (C) a paraffin section stained with safranin/light green (alginate stains in orange and the host tissue and entrapped cells stain in blue) where transplanted hMSC (arrows) can be observed within the disk, and (D) a paraffin section immunolabelled with an anti-human specific monoclonal antibody showing an hMSC still inside the hydrogel (dashed arrow) and another one at the hydrogel periphery (solid arrow), in close proximity with the host mice cells (autofluorescence green staining) (⁄ alginate; orange line – alginate matrix boundary).

divalent cations that crosslink the hydrogel. Additionally, alginate can be modified to become degradable in physiological conditions by partial oxidation of alginate chains using sodium periodate [56]. The periodate oxidation cleaves the carbon–carbon bond of the cis-diol group in the uronate residue and alters the chain conformation to an upon-chain adduct, which behaves similar to an acetal group susceptible to hydrolysis. The degradation rate can be further regulated by adjusting the molecular weight distribution of oxidized alginates without varying the number of oxidized uronic acids, chain inflexibility and gel-forming ability [35]. The molecular weight can be modified by c-irradiation, which breaks the alginate chain [47]. The oxidized binary hydrogels improved the formation of bone tissue compared to non-modified alginate, since a faster degradation occurs that facilitates the formation of new bone tissues [35]. Additionally, alginates are chemically versatile, allowing the easy incorporation of biochemical cues to engineer specific cell responses. For example, as cells have no specific receptors for binding to alginate, several methods to promote cell attachment to alginate matrices have been developed. These include the coupling of ECM proteins such as laminin, collagen and fibronectin [57–59]. However, since the coupling of an entire protein is difficult to control, can lead to non-specific interactions, may elicit an immune response and proteins are subject to proteolytic degradation [60], the decoration of biomaterials with cell recognition motifs, such as small immobilized peptides, has become popular. The arginine–glycine–aspartic acid (RGD) sequence was one of the first peptides to be used to promote cell adhesion on a biomaterial and is still one of the most widely employed. This tripeptide motif corresponds to the minimal essential cell adhesion peptide sequence identified in many ECM proteins, such as fibronectin, collagen, laminin, osteopontin and vitronectin, which are associated with integrins in cell surface membranes [60]. To promote cell adhesion, alginate can be functionalized with this peptide using the aqueous carbodiimide chemistry, as demon-

strated by Rowley et al. [34]. In this case, RGD peptides are linked via a stable covalent amide bond and this reaction occurs between the carboxyl group of the alginate and the N-terminus of the peptide. Several other combined modifications have been investigated to improve alginate properties. For instance, the same chemical approach was used to further tailor alginate hydrogels to a more sophisticated ECM-like 3-D cellular microenvironment. Alginate was grafted not only with RGD but also with a protease-labile crosslinking peptide (proline–valine–glycine–leucine–isoleucine– glycine, PVGLIG) that is cleavable by metalloproteinases (MMPs) produced by cells [14,15]. The resultant MMP-sensitive hydrogels can be partially remodeled by cell-driven proteolytic mechanisms, leading to increased cellular evasion/invasion, and are particularly appealing vehicles for cell-delivery strategies. In another example, alginate was chemically functionalized with cell signaling moieties such as galactose to improve hepatocyte cell recognition [61,62]. Since hepatocytes have a specific receptor that recognizes this ligand, cells entrapped within alginate bearing galactose residues present higher functionality and survival rate. Also, to mimic the rheological behavior of the nucleus pulposus in the intervertebral disk, Leone et al. [63]have covalently crosslinked alginate by amide bond formation, after activation of the carboxylic groups and their conversion into amide moieties. They were able to attain a rheological behavior very similar to the one of the human nucleus pulposus and, when chondrocytes were added, they were able to proliferate and produce ECM proteins. According to the aim of the strategy, several other alginate modifications have been proposed to improve its behavior. Several alginate derivatives, their properties and possible applications are described in detail in a recent review by Pawar and Edgar [64]. One of the greatest potentials of alginate is its mild and diverse gel-forming capacity. As described below, alginate hydrogels can be prepared by various crosslinking methods, which give rise to different delivery strategies with equally different outcomes.

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3.1. Formation of alginate hydrogels Alginate hydrogels can be formed using chemical or physical crosslinking strategies, ionic methods being the most widely used. An overview of some of the crosslinking strategies that have been reported is presented in the following sections, and in Fig. 3 some of these strategies are represented schematically. 3.1.1. Ionic crosslinking The most common method to achieve alginate gelation and crosslinking is through the exchange of sodium ions from guluronic acid units with divalent cations such as calcium (Ca2+), strontium (Sr2+) and barium (Ba2+) [67] (Fig. 3A). As alginates present different affinities towards the different divalent ions, they give rise to gels with different stability, permeability and strength, depending on the cation used [68]. The stability of these hydrogels depends on the exchange between monovalent cations from the surrounding environment and the divalent cations from the hydrogel, resulting in diminished mechanical properties, which can limit their use for cell studies over long periods of time [48]. This process depends on several alginate features, such as concentration, source, degree and type of crosslinking, and molecular weight [69]. Moreover, the type of physiological conditions (in vitro or in vivo) will have an effect on hydrogel stability. Alginate stability is also variable in vitro. For instance, calcium-crosslinked alginate hydrogels rapidly lose stability in 0.9 wt.% sodium chloride, due to the exchange of calcium by non-gelling monovalent sodium ions [69], as well as in solutions containing chelators such as citrate or phosphate [70], which act as de-crosslinking agents by removing calcium ions. Calcium alginate hydrogels can nevertheless remain stable for weeks in cell culture medium, which generally contains sufficiently high calcium concentrations to counteract such effects [71]. Ionic crosslinking of alginate can be further obtained by external or internal gelation. 3.1.1.1. External gelation. External gelation of alginate hydrogels using soluble salts of divalent cations, such as calcium chloride (CaCl2), as ionic crosslinker agents is one of the most frequently used procedures, as it is a very simple process that provides immediate and non-toxic cell entrapment [50]. However, although gelation occurs almost instantaneously, this process frequently results in unbalanced crosslinking density and a polymer concentration gradient within the formed hydrogel [72]. This method is commonly used to obtain gel beads by dripping a sodium alginate solution into an aqueous solution of calcium ions. Alginate microbeads have been widely used as cell delivery systems for a range of applications aimed at tissue regeneration

[17,19,20,73–95]. These microbeads are generally produced by coaxial airflow, which controls the size of the droplets by blowing them from a needle tip into a CaCl2 bath before they fall due to gravity. For a given alginate solution (concentration, type and viscosity), the size of the beads depends on the airflow, the needle diameter and solution flow rate [48]. To attain smaller beads (150 lm), other techniques must be used, such as electrostatic bead generation. In this case, a voltage is applied between the needle and the electroconductive solution underneath [96]. Therefore, it is possible to change the droplet size by adjusting the voltage. Nowadays, different types of microbead generators are commercially available. In a recent work, a novel methodology involving the use of superhydrophobic substrates was reported to efficiently produce spherical hydrogels entrapping rat MSCs isolated from bone marrow and fibronectin without the need of a precipitation bath [97]. Alginate beads were formed by applying microdrops of lowviscosity alginate onto superhydrophobic surfaces, then adding drops of CaCl2 on top of each of them to crosslink the alginate. This methodology has a number of advantages over the conventional techniques, such as reduced mechanical forces and particle aggregation, which are coupled with decreased cell loss. Overall, alginate beads for tissue engineering should allow good permeability for nutrients and oxygen while providing optimal cell survival conditions for several days. Moreover, these beads allow the fixation of cells in the damaged tissue site, thus preventing cell removal via the bloodstream [97]. Several examples in which alginate beads were used for tissue regeneration are discussed below. Although clearly demonstrating the enormous potential of alginate microbeads for cell entrapment, these works report the use of different conditions. Thus, the different alginate concentrations, molecular weights and even cell concentrations used are highlighted. 3.1.1.2. Internal gelation. Internal gelation strategies are being widely investigated with a view to promoting in situ hydrogel formation. This way, a polymeric solution combined with cells can be injected in a liquid state and will then form a hydrogel at the site of interest. Although these strategies have not been as extensively explored as external gelation, there is an increasing number of reports in the literature describing the use of internal gelation of alginate hydrogels for cell delivery in different applications aimed at tissue regeneration [14,18,49,98,99]. The most common strategy for promoting the internal gelation of alginate hydrogels consists in using divalent cation salts of low solubility, which enable the gelation rate to be slowed down and hence afford better control over the gelation time. Calcium carbon-

Fig. 3. Schematic representation of alginate hydrogel formation. (A) Ionic crosslinking with calcium-induced chain–chain association of guluronate blocks forms the junction zones responsible for gel formation (egg-box model) [65]. (B) Covalent crosslinking with PEG-diamines by carbodiimide chemistry. The hydrogel is composed of two different polymers with different mechanical properties: PEG chains provide some elasticity, while the alginate chains provide mechanical strength [66].


ate (CaCO3) and calcium sulfate (CaSO4) have been widely used for this purpose. They have low solubility in pure water at neutral pH, but are soluble under acidic conditions, allowing its uniform distribution in the alginate solution before gelation occurs [72,100]. Free calcium ions can later be released from these salts by slightly decreasing the pH with glucone-d-lactone (GDL), thereby allowing gradual gelation. In the case of CaCO3, the CaCO3/GDL molar ratio can be set at 0.5 to yield a neutral pH [14,72]. It is notable that, although there is a slight decrease in pH, the cell viability is not affected [14,18]. As an alternative to GDL, a photoacid generator (PAG) has been proposed that dissociates under UV light, releasing H+ ions, which will react with CaCO3 to create Ca2+ [101]. Despite the great potential of these light-derived hydrogels as cell delivery systems for several biomedical applications, their cytocompatibility needs to be investigated further. Another way to promote photoactivated internal gelation consists in the use of water-soluble Ca2+ chelators (‘‘cages’’), which can be mixed with alginate solutions and, upon light exposure, will undergo an irreversible molecular change that decreases their affinity to Ca2+. This will result in the release of Ca2+, which will subsequently trigger crosslinking [102]. In contrast to the use of insoluble calcium salts, this methodology allows homogeneous alginate hydrogels to be obtained even in concentrated alginate solutions (e.g. 10 wt.%). Using this method, improvements in mechanical properties and homogeneity have been observed over comparable alginate concentrations. 3.1.2. Covalent crosslinking Covalent crosslinking of alginate hydrogels can be achieved by a variety of different methods, and usually provides more stable and mechanically stronger gels than ionic crosslinking. It is beyond the scope of the present article to review them all; instead, we focus on the strategies used for cell entrapment and delivery. Generally, once a material is covalently crosslinked, it no longer meets the injectability criteria. There are, however, a few exceptions, such as photocrosslinking alginates [103] and shape-memory alginate scaffolds [104], which will be explained later on. The major disadvantages of these methods are their increased complexity and the eventual toxicity of the reagents used. Covalently crosslinked alginate hydrogels can be synthesized with a wide range of mechanical properties using, for example, PEG-diamine molecules with different molecular weights as crosslinkers [66] (Fig. 3B). The elastic modulus of the crosslinked alginate changes according to the molecular weight of the PEG molecules. The hydrogel properties can be further regulated by multifunctional crosslinking molecules, which provide a wider range and tighter control over degradation rates and mechanical stiffness, as demonstrated by Lee et al. [105]. In their work, PAG hydrogels were formed with either poly(acrylamide-co-hydrazide) as a multifunctional crosslinking molecule or adipic acid dihydrazide (AAD) as a bifunctional crosslinking molecule. This multicrosslinking strategy led to the formation of stronger hydrogels. 3.1.2.1. Photocrosslinking. Injectable photocrosslinkable alginates have been proposed for tissue engineering applications and have been designed to allow better control over mechanical properties, swelling ratios and degradation rates than ionically crosslinked alginates [50,103,106,107]. Their cell-signaling ability can also be tailored by incorporating biochemical signals such as growth factors and cell adhesive peptides [108,109]. Photocrosslinkable alginate can be delivered in a minimally invasive way and then rapidly crosslinked in physiological conditions in situ following a brief exposure to ultraviolet (UV) light [110]. This process occurs under mild conditions and therefore can be performed in direct contact with cells. For that purpose, alginate was modified with 2-aminoethyl methacrylate using

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carbodiimide chemistry and these methacrylated alginates were subsequently photocrosslinked using UV light with a photoinitiator [103]. Photocrosslinkable alginate with controlled degradation and cell adhesive properties have attracted great interest with regard ton tissue engineering applications [103,109,111,112], although they are still relatively new and seem to require further studies to assess their effectiveness. 3.1.2.2. Shape-memory alginate scaffolds. A novel type of covalently crosslinked alginate, capable of retaining its shape and forming a macroporous structure, has been developed. The concept involves a hydrogel that allows control over its size and shape. Basically, a hydrogel is formed into the desired shape, then is collapsed in vitro for storage and handling purposes and, finally, is reformed in its original shape in vivo by rehydration [6,104]. It is noteworthy that, although these covalently cross-linked alginate hydrogels present great potential as injectable bulking agents, they exhibit less contourability than other alginate hydrogels. On the other hand, they display greater shape definition, with a wider range of physical and mechanical properties. To obtain these shape-memory scaffolds, alginate was covalently crosslinked with AAD and the hydrogels were formed by standard carbodiimide chemistry using 1-ethyl-(dimethyl aminopropyl) carbodiimide, 1-hydroxybenzotriazole and AAD [104,113]. Thornton et al. [104]were able to produce macroporous alginate hydrogel scaffolds with a predefined geometry, which were then dehydrated and compressed into small, temporary forms. The compressed scaffolds were then minimally invasively delivered to the dorsal subcutaneous space of a mouse through a catheter. Next, they were successfully rehydrated in situ by the injection of PBS. Overall, this work demonstrated the potential of these covalently crosslinked hydrogels as an injectable delivery system. More recently, Wang and colleagues [6] showed that these shape-memory scaffolds have the potential to serve as a synthetic matrix for skeletal muscle cell survival, proliferation and migration. These results confirm the potential of these shape-memory alginate scaffolds as cell delivery systems for tissue regeneration, although this strategy needs further exploration. 4. Alginate-based injectable cell delivery systems 4.1. Bone regeneration In the fields of orthopedics and oral and maxillofacial surgery, bone regeneration remains a clinical challenge, despite the introduction of various bone augmentation techniques and bone graft materials. The current standard treatment is based on the use of bone graft materials that are divided into two major groups: natural and synthetic bone grafts [114]. Natural bone grafts include auto-, allo- (human donors) and xenografts (other species). Autologous bone graft is considered the ‘‘gold standard’’ for bone repair and regeneration by many surgeons, mainly due to lack of immunogenic reaction and optimal biological performance in terms of osteogenicity, osteoinductivity and osteoconductivity [115]. However, such bone grafts present several limitations, namely limited availability, the risks associated with harvesting, the additional surgical procedure required, donor site morbidity, post-operative pain and infection. On the other hand, allografts and xenografts are widely available and there is no need for additional surgery. However, with both types of graft there is the risk of immunoreaction and, since they undergo several processing techniques, their osteoinductive and osteoconductive potentials are reduced [115– 117]. To overcome the above-mentioned limitations of natural grafting, a large number of synthetic grafts have been designed

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over the past decades. These materials present several advantages, such as wide availability, lack of antigenic response and easy tailoring to a specific application. However, their biological performance regarding the induction and support of bone growth is inferior to those of natural grafts [115]. One way to overcome this is to add cells and/or growth factors to the materials, to produce more biologically effective systems capable of restoring, maintaining or improving bone function [118]. In the subsequent paragraphs several promising alternative strategies to the current therapies are described in which injectable alginate-based cell delivery systems have been used with the aim of promoting bone regeneration (Table 1). The first studies to report the culture of bone cells in a 3-D alginate-based microenvironment were carried out using unmodified hydrogels and committed cell types, such as the mouse calvarial3T3 osteoblastic cell line (MC3T3) or primary chick embryo osteoblasts. In one of the first reports, Majmudar et al. [122]observed that osteoblasts were kept viable for as long as 8 months within alginate beads. Using an in situ gelling strategy, Kuo and Ma [72]were able to control the gelation rate of alginate hydrogels, and successfully entrapped MC3T3 pre-osteoblastic cells. Due to the importance of cell–ECM interactions in cell adhesion, proliferation and differentiation, Alsberg et al. [119]explored for the first time the use of RGD-modified alginate for the entrapment of MC3T3-E1. They observed that RGD–alginate enhanced in vivo bone formation after 16 and 24 weeks compared to unmodified alginate. Later, Evangelista et al. [20]demonstrated the importance of these peptides not only in promoting cell adhesion but also in enhancing osteogenic differentiation. MSCs or bone marrow stromal cells (BMSCs) are an important cell population, being a mandatory first element for an unimpeded bone repair process, so have shown great potential in bone tissue engineering [123]. Markusen et al. [124]demonstrated that hMSCs isolated from bone marrow could be entrapped within alginateGRGDY (glycine-arginine-gycine-aspartic acid-tyrosine) beads with viability higher than 80%. Duggal et al. [125] later investigated the phenotype and gene expression of hMSC from bone marrow and adipose tissue in a 3-D RGD–alginate system as compared to a 2-D culture. They observed that hMSCs within the beads acquired a compact morphology, maintained high cell viability and almost ceased their proliferation. They also observed that hMSCs in 3-D RGD–alginate matrices presented more similarities with hMSCs cultured in two dimensions than with uncultured freshly isolated

hMSCs by studying gene expression through a microarray analysis (an Affymetrix GeneChip Human Gnome U133A 2.0 Array platform, which contains 22,000 probes that represent 14,500 genes associated with early development, intracellular signaling, cell shape and differentiation). This study demonstrated that 3-D culture within RGD–alginate did not present the same gene expression profile as their cells of origin. In the search for a bone tissue engineering application, Bidarra et al. [17]reported the successful entrapment of bone marrow-derived hMSCs within 2 wt.% RGDmodified alginate microbeads cultured under dynamic conditions for 21 days, as shown in Fig. 4. Upon osteogenic stimulation, the hMSCs were able to differentiate along the osteoblastic lineage and also to stimulate neighboring endothelial cells to form tubelike structures on matrigel, thus demonstrating their pro-angiogenic ability. To assess this ability, hMSCs within RGD–alginate beads were entrapped within matrigel and, after matrigel gelation, endothelial cells were added on top. After 48 h, tube-like structures were counted and compared to those formed by endothelial cells seeded on matrigel alone. The combination of stem cells with growth factors within a matrix may lead to increased bone formation. To explore this possibility, Simmons et al. [120]investigated whether BMSCs entrapped within 2 wt.% irradiated alginate with RGD could enhance bone regeneration, in the absence or presence of bone morphogenetic protein (BMP)-2 and/or transforming growth factor (TGF)-b3. The authors observed that only BMSCs transplanted in the presence of both factors produced significant bone tissue. Therefore, the entrapment of progenitor cells and growth factors in a single degradable matrix may provide a more potent strategy for bone tissue formation. To amplify the MSCs potential, in some cases, these cells have been genetically modified with genes that codifiy for factors that have a crucial role in bone regeneration, such as BMP-2. Rat bone marrow-derived MSCs containing a BMP-2 transgene, when injected directly into articular osteotomies in nude rats, were capable of accelerating bone healing compared to MSCs without the BMP-2 gene or MSCs with BMP-2 injected within alginate [121]. In this example, the authors observed that the alginate carriers induced the formation of a chondroid mass that impeded bone healing, therefore suggesting that alginate may not be a favorable vehicle for these genetically modified MSCs. It is noteworthy that a high molecular weight alginate without cell adhesion peptides was used in this study, and no other formulations were tested,

Table 1 Examples of injectable alginate hydrogels proposed for cell delivery in bone tissue engineering applications (these data are not intended to be all-inclusive). Alginate supplier

Alginate characteristics

Alginate concentration (wt.%)

Hydrogel formation (bead diameter, lm)

Cell type (cell density, 105 cells ml1)

Refs.

Sigma–Aldrich FMC BioPolymer FMC BioPolymer; Sigma FMC BioPolymer

n/a Ultrapure Ultrapure MVG; with RGD

1.2 n/a 2

EG with CaCl2 (352) EG with CaCl2 (2000) IG with CaSO4

Rabbit MSC (10) Human PDLSC and GMSC (10) Swine MSC (50)

[74] [75] [98]

MVG Protonal LF20/40, HMW and LMW, oxidized with RGD MVG Protonal LF20/40, HMW and LMW, oxidized with RGD MVG with RGD

2

EG with CaCl2 (700)

hMSC (200)

[17]

2

EG with CaCl2 (200–500)

HOC and HUVEC (200)

[19]

2 2

MC3T3 (in vitro, 220) (in vivo, 315) MC3T3 (200)

[119]

MVG Protonal LF20/40,HMW and LMW, oxidized with RGD MVG irradiated with RGD Pronova SLG 100 n/a Intermediate guluronic content

EG with CaCl2 (1400) IG with CaSO4 EG with CaCl2 (1000)

2 1.2 2.5 1.5

IG with CaSO4 EG EG with CaCl2 EG with CaCl2 (300)

BMSCs (200) Rat MSC (0.5) Rabbit MSC (10) Murine ATSC (10)

[120] [121] [77] [78]

FMC BioPolymer FMC BioPolymer FMC BioPolymer FMC BioPolymer FMC BioPolymer Sigma–Aldrich Sigma Aldrich

[20]

ATSC, adipose tissue stromal cells; BMSC, bone marrow stromal cells; EG, external gelation; GMSC, gingival mesenchymal stem cells; hMSC, human mesenchymal stem cells; HMW, high molecular weight; HOC, human osteoprogenitor cells; HUVEC, human umbilical vein endothelial cells; IG, internal gelation; LMW, low molecular weight; n/a, not available; PDLSC, periodontal ligament stem cells.


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Fig. 4. Osteogenic differentiation of hMSCs inside alginate hydrogel microspheres: (A) hMSCs within 2 wt.% RGD–alginate microspheres with a cell density of 2 107 cells ml1 remain with a high viability after 21 days in culture, (B) as demostrated by the Live/Dead assay (viable cells stain green), (C) express high levels of ALP (a common osteogenic marker; ALP + cells stain pink) activity when cultured under standard osteoinductive conditions [17].

which could presumably yield more favorable results. In another study, using miniature swine bone marrow-derived MSCs expressing BMP-2, the authors compared three different types of alginates (regular alginate, ultrapure alginate and RGD-modified alginate) and also collagen type I as injectable cell delivery systems [126]. The crosslinking of the alginate–cell suspension was achieved by internal gelation with CaSO4. The authors reported a significantly better behavior of collagen type I, although detailed experimental information is not available, such as the molecular weight of the alginate, the concentration of RGD or even the source of alginate. Moreover, the alginate was not modified to improve its biodegradability, such as recurring to alginate oxidation, which alters chain conformation to an upon-chain adduct that is susceptible to hydrolysis in physiological conditions. Due to their flexibility, alginate beads with cells can also be formed in the presence of antibiotics, giving rise to a delivery system that is simultaneously capable of helping in the treatment of infected defects. For example, Ueng and colleagues [77] entrapped rabbit bone marrow-derived MSC with vancomycin within 2.5 wt.% alginate for the treatment and prevention of surgical bone infections. They were able to demonstrate the sustained release of vancomycin over 14 days in vitro and the released concentration exceeded the minimum inhibitory concentration against Staphylococcus aureus (determined by an antibiotic disk diffusion method applied to the nutrient broth). Moreover, when alginate beads loaded with vancomycin and rabbit MSCs were implanted in bone cavities created in rabbit lateral femoral condyles, they were able to contribute to the newly formed bone and a substantial mineral deposition occurred. Nevertheless, as highlighted by the authors, further studies need to be performed in order to specifically investigate the effect of the antibiotic in the combined system using an infected bone model. In more recent years interesting new cell sources have been attracting attention in the tissue engineering field. For example, using adipose-derived cells from mice, Abbah et al. [78] observed, in vitro, more cell proliferation and alkaline phosphatase (ALP) activity within 1.5 wt.% alginate than in cells growing in a monolayer. It is noteworthy that this is one of the few studies where significant cell proliferation in three dimensions was observed, with the cell number almost doubling during the first week in culture. Other authors used adipose-derived cells and platelet-rich plasma (PRP) from rabbits within 1.2 wt.% alginate to produce beads. They observed that these cells were capable of enhancing vascularization and mineralization in vivo [74]. To assess the possibility of using dental-derived stem cells for craniofacial applications, since they might be more likely to differentiate into craniofacial tissues, Moshaverinia and colleagues [75] explored the potential of two such types of stem cells (human periodontal ligament stem cells and human gingival MSC) entrapped within 1 wt.% oxidized alginate beads. They demonstrated for the first time that these cells were able to differentiate, in vitro, into

osteogenic and adipogenic lineages, thus providing a promising strategy for bone tissue engineering. Another strategy to enhance bone regeneration relies on the use of co-cultures of bone cells and endothelial cells. In this regard, Grellier et al. [19]entrapped primary human osteoprogenitor cells together with primary HUVECs within 2 wt.% RGD-modified alginate microbeads and showed that the co-transplantation of these two cell types in femoral defects of mice promoted a significant increase in mineralization after 6 weeks of implantation, as depicted in Fig. 5.

4.2. Cartilage regeneration Adult articular cartilage has a very limited reparative capacity, hence, in adults, cartilage defects do not regenerate. Efforts to achieve repair of these lesions have been limited by the challenge of stimulating the resident cells to form new cartilage. Among the many different clinical strategies to induce the formation of cartilaginous tissue, the implantation of autologous chondrocytes and stem cells have been explored in the pursuit of cartilage repair [127,128]. There are many examples of strategies involving the use of alginate as a cell delivery vehicle for cartilage regeneration. Regarding the presence of adhesion peptides in the alginate matrix, it has already been demonstrated that these motifs support chondrocyte and MSC adhesion, but in a 3-D environment MSC-induced chondrogenesis may be inhibited when cells interacted with RGD peptides [129]. This study points to the importance of designing

Fig. 5. Effect of human osteoprogenitors and endothelial cells co-immobilized within RGD–alginate microspheres on mineralization in a bone defect. Bone defects were performed in the femoral metaphysis of nude mice, and RGD–alginate microspheres containing no cells, only human osteoprogenitors (HOPs) or HOPs coimmobilized with HUVECs were implanted. Quantification of mineralization revealed by von Kossa staining 3 and 6 weeks post-implantation. After implantation for 6 weeks, mineralization of co-immobilized HOPs and HUVECs was significantly higher than with HOPs immobilized alone. ⁄p 6 0.05; ⁄⁄p 6 0.01 [19]. Copyright 2009 Elsevier Ltd.

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Table 2 Examples of injectable alginate hydrogels proposed for cell delivery in cartilage tissue engineering applications (these data are not intended to be all-inclusive). Alginate supplier





Refs.

FMC BioPolymer FMC BioPolymer n/a Sigma–Aldrich Sigma–Aldrich Sigma–Aldrich Sigma–Aldrich FMC BioPolymer

LMW (150 kDa) LMW (150 kDa) n/a Low viscosity n/a n/a n/a LF 20/40 Protonal; low viscosity M/G = 0.49; medium viscosity M/G = 1.6 LF 20/40 Protonal; 19,600 g mol–1 M/G ratio = 1–1.4; ultrapure M/G ratio = 1.56; 75,000–100,000 g mol–1; viscosity = 0.25 Pa s LF 20/40 Protonal; 19,600 g mol–1 n/a n/a n/a n/a n/a n/a

2 2 1.2 1.2 1.2 1.2 1.2 1.5 and 2

EG EG EG EG EG EG EG EG

ASCs (250) ASCs (250) TGF-b3 CHOs (100) CHOs and MSCs (5) Rabbit MSCs (10) IGF-I/FGF-2 rabbit CHOs (40) Synovium MSCs (70) Calf CHOs (330)

[79] [80] [130] [82] [83] [84] [85] [90]

2 1 1.2

Photocrosslinked EG with CaCl2 EG with CaCl2 (676)

Bovine CHOs (100) Rabbit BMSCs (250) Human CSP (120)

[109] [86] [131]

2 1 2 0.8, 1.2 and 2 1.2 1.2 1.2

Photocrosslinked EG with CaCl2 (2000) EG with CaCl2 EG with CaCl2 (3000) EG with CaCl2 EG with CaCl2 EG with CaCl2

Bovine CHOs (100) Rat CHOs (5) Human CHOs (200) Calf CHOs (25,000 and 100,000 cells per bead) FGF-2 rabbit CHOs (20) IGF-I rabbit CHOs (20) Rabbit CHOs (20)

[103] [88] [132] [89] [133] [134] [135]

FMC Biopolymer Sea Matrix n/a FMC BioPolymer Fluka FMC BioPolymer FMC BioPolymer Sigma–Aldrich Sigma–Aldrich Sigma–Aldrich

with with with with with with with with

CaCl2 CaCl2 CaCl2 CaCl2 CaCl2 CaCl2 CaCl2 CaCl2 (800)

ASC, adipose stem cells; BMSC, bone stromal marrow cells; CHO, chondrocytes; CSP, mesenchymal progenitor cells from the subchondral bone marrow; EG, external gelation; FGF-1, fibroblast growth factor 1; G, guluronic acid; HMW, high molecular weight; IG, internal gelation; IGF-1, insulin-like growth factor 1; LMW, low molecular weight; M, mannuronic acid; MSC, mesenchymal stem cells; n/a, not available.

appropriate scaffolds and biomaterials for cartilage engineering applications. In fact, the majority of the studies with MSCs for chondrogenesis do not use modified alginate, which is in agreement with the fact that they are not anchorage-dependent cells. In the majority of the studies published, cells are delivered in alginate beads (Table 2). It is noteworthy that one major parameter that differs between studies is the cell source used by the authors. Independently of the cell source, alginate has shown great potential in cartilage regeneration. Therefore, the following examples are discussed considering the cell origin. Several studies report the use of fully differentiated chondrocytes entrapped within an alginate matrix. For instance, primary rat articular chondrocytes entrapped within 1 wt.% alginate were able to proliferate and maintain their phenotype [88]. In fact, these cells showed a significantly higher proliferation rate in 3-D than in 2-D cultures, and showed continued growth for 40 days. For cartilage regeneration, alginate hydrogels seem to provide an adequate environment since chondrocytes are not anchorage-dependent cells and their natural microenvironment consists in a hydrated matrix. Dobratz and colleagues [132] prepared an injectable in situ crosslinking system to produce an engineered cartilage by injecting a CaCl2 solution immediately after mixing a 2 wt.% alginate solution with primary perichondrium-derived human chondrocytes. This minimally invasive system allowed for an in vivo molding that maintained its shape for at least 38 weeks after injection. Jeon and co-authors [103,109]developed a photocrosslinkable, biodegradable alginate system with controlled cell adhesive properties in which bovine articular chondrocytes were successfully entrapped. In this study, chondrocytes within RGD–alginate exhibited a higher proliferation rate and greater chondrogenic differentiation compared to the cells entrapped within unmodified hydrogels. It is noteworthy that these hydrogels contained TGFb1. This new photocrosslinkable biomaterial constitutes a promising approach to cartilage tissue engineering. In an attempt to overcome some of the limitations of chondrocyte transplantation, cells are being genetically modified to enhance repair. For instance, porcine TGF-b3-secreting chondrocytes combined with porcine synovial-derived MSCs within alginate beads were able not only to induce MSCs differentiation, but also to prevent chondrocyte dedifferentiation [130]. In addition,

the combination of both types of cell presented the highest level of chondrogenesis in vivo. In other examples, rabbit chondrocytes were modified to overexpress two growth factors simultaneously or separately, namely insulin-like growth factor (IGF)-I and fibroblast growth factor (FGF)-2 [84,133–135]. In these works, the 1.2 wt.% alginate beads entrapping these genetically modified cells promoted important cartilage repair in vivo. IGF-I is an anabolic growth factor of chondrocytes that has been shown to stimulate the production of proteins from the cartilage ECM (e.g. collagen type II and proteoglycans), whereas FGF-2 is mitogenic for articular chondrocytes [136]. Normal chondrocytes produce IGF-I and, in the 3-D alginate environment, these cells continue to produce this factor [89]. Moreover, Yoon et al. [89]demonstrated that the association of higher cell densities and higher matrix concentrations leads to an enhancement in IGF-I production, since IGF-I expression by calf chondrocytes was increased with smaller cell-to-cell distances (higher cells densities). MSCs have been explored as an alternative cell source in cartilage regeneration strategies. The chondrogenic potential of hMSCs isolated from subchondral bone marrow was assessed after entrapment within 1.2 wt.% alginate beads and under chondrogenic stimulation [87]. Although further in vivo studies are necessary, the authors demonstrated the relevance of the calcium–alginate system for stem cell injection by obtaining full chondrogenic differentiation within the alginate matrix. The chondrogenic potential of bone marrow stem cells isolated from rabbits was also assessed by entrapping these cells and comparing their behavior in ultrapure (UP) and commercial grade sodium alginate [86]. The authors observed a significant enhancement of cell proliferation and differentiation in the UP compared to the commercial grade alginate. When UP beads with cells were implanted, a significant improvement in cartilage repair was observed compared to commercial grade alginate, which points to the importance of purification alginate in such applications. Dashtdar et al. [83]demonstrated that, independently of their state of differentiation, transplanted rabbit bone marrow-derived MSCs entrapped within 1.2 wt.% alginate beads promoted greater healing compared to untreated controls in articular cartilage defects in rabbit. In a comparative analysis, the effectiveness of rabbit allogeneic MSCs was evaluated against that of autologous chondrocytes in repairing focal cartilage defects in an in vivo rabbit model [82]. When transplanted within 1.2 wt.%


alginate beads, the two cell types promoted similar chondrogenic differentiation and produced selective up-regulation of cartilagespecific genes. These results support the potential use of MSCs for articular cartilage repair. In other studies, MSCs were modified with the purpose of producing TGF-b3 – a commonly used growth factor to induce MSCs chondrogenesis – and simultaneously suppressing collagen type I expression [85,136,137]. Although further studies are needed to optimize the efficiency of this strategy, encouraging in vitro results were obtained when these cells were entrapped within 1.2 wt.% alginate. Overall, the treatment of articular cartilage defects with stem cells of mesenchymal origin has achieved satisfactory outcomes and MSCs seem to present several advantages compared to chondrocytes, which have limited proliferation capacity and, in some cases, even seem to lose their cartilaginous phenotype during expansion [82]. However, although adult stem cells reveal such great potential, the majority of the studies in cartilage regeneration have used differentiated chondrocytes as the cell source [81,90,138,139] and, despite their limitations, most of the clinical trials that use these cells have good repair outcomes (i.e. the cells were able to form new cartilage). Adipose-derived stem cells (ASCs) have emerged as another promising adult cell source for cartilage regeneration because of their easy accessibility and chondrogenic potential [140]. Additionally, ASCs have demonstrated the ability to secrete a wide range of trophic factors that can stimulate endogenous cartilage regeneration. Lee and co-authors [79,80] optimized the cell culture conditions for entrapped rat ASCs within alginate microbeads and, after implantation in a focal cartilage repair site, tissue infiltration and perichondrium were observed. 4.3. Intervertebral disk regeneration Intervertebral disk (IVD) degeneration is a major health concern worldwide. Replacement of the nucleus pulposus (NP) of the IVD, a natural hydrated structure, with injectable biomaterials represents a potential treatment strategy for IVD degeneration. Naturally derived materials such as alginate have been used as models to study the behavior of NP cells in a 3-D environment [141–146]. In several studies, NP cells were cultured in 3-D alginate beads to assess their behavior under the influence of different conditions [147–150]. For instance, when bovine NP cells within 1.2 wt.% beads were cultured in the presence of bovine lactoferricin there was a decrease in inflammatory mediators involved in degenerative disk disease (cytokine interleukin-1 and endotoxin lipolysaccharide) that mediate the suppression of prostaglandin (PG) accumulation [147]. As a consequence, there was an up-regulation of PG synthesis, aggrecan expression and pericellular matrix formation. In another example, alginate beads with NP cells were used as a model to evaluate the toxicity of some compounds commonly used in spinal surgery [148,149]. In a different study, the influence of different medium compositions, one with human PRP and another with TGF-b1, on chondrogenic differentiation of hMSCs comparison with NP cells was investigated [150]. In comparison with TGF-b1, PRP did not induce adequate chondrogenic differentiation. These types of studies are key for better understanding how cell biology influences disk degeneration and how this process could be manipulated. For instance, Stephan et al. [142]evaluated a range of cell densities entrapped in alginate under different culture conditions (standard culture conditions with 3.15 g l1 glucose and 10% serum or without glucose and/or 20% serum), and demonstrated that NP cell growth and survival were both influenced by cell density and the availability of serum or nutrients such as glucose. For instance, additional serum (20%) promoted cell proliferation and the formation of large cell clusters, and even larger cell clusters were created without glucose. Additionally, lower cell densities led to higher cell

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proliferation and cluster formation. Bron and colleagues [143] analyzed the effect of different alginate concentrations (2, 4 and 6 wt.%) and preparation conditions (external and internal gelation) on the gene expression of components of the IVD’s ECM. They found that 2 wt.% alginate disks prepared by external gelation closely matched the viscoelastic properties of the nucleus pulposus. It is noteworthy that alginate was not modified to promote cell adhesion, hence cells were not able to actively respond to matrix stiffness via focal adhesion sites. Therefore, no pronounced effect of matrix stiffness was observed on ECM synthesis. Using an alternative crosslinking approach, Chou et al. [111]studied the potential of photocrosslinked alginate for NP cell encapsulation and in vivo repair, and further compared it with ionically crosslinked alginate. Photocrosslinked alginate maintained structural integrity, as illustrated in Fig. 6, and presented a higher number of NP cells producing ECM components (type II collagen and chondroitin sulfate proteoglycan) compared to ionically crosslinked alginate in vitro. Additionally, increases in gene expression (type II collagen and aggrecan) and ECM accumulation (type II collagen and sulfated glycosaminoglycan) were obtained by photocrosslinked alginate after 8 weeks in vivo. At this time the matrices presented an equilibrium Young’s modulus of around 4.31 kPa, which is close to the native NP equilibrium modulus reported in other studies [183,184] that employed a similar unconfined compression testing protocol (5–6.7 kPa). Generally, these studies underscore the potential of alginate as a synthetic ECM for IVD engineering. 4.4. Adipose tissue regeneration Contour defects due to loss of soft tissue, largely composed of subcutaneous adipose tissue, are associated with tumor resection, trauma and congenital abnormalities. Current treatment for soft tissue reconstruction use autologous fat transplantation, alloplastic implants and autologous tissue flaps, which present a number of challenges and limitations, such as donor site morbidity and volume loss over time, implant migration and foreign body reaction [49,151]. Several studies have tried to overcome these limitations through the use of human adipose stem cells (hADSCs) in an attempt to engineer adipose tissue in vivo. hADSCs have several attractive features, including a high expansion potential in culture, the ability to be readily induced to differentiate into adipocytes in vitro, easy isolation and isolation in large quantities. The feasibility of using an injectable hydrogel such as alginate as a delivery vehicle for these cells for contour improvement has been demonstrated in a few studies (Table 3). Kang and co-authors [99] showed that an RGD-modified alginate matrix enhanced adipogenic differentiation, demonstrating that the RGD peptides function not only as cell adhesion sites, but also as adipogenic stimulator for ADSCs. In another study, the behavior of hADSC was analyzed within alginate matrices prepared using two different crosslinking agents: calcium chloride and calcium gluconate [151]. Both matrices maintained cell viability (higher than 80%) and supported adipogenic differentiation (positive staining for neutral lipid accumulation and increased levels of perilipin expression). It is noteworthy that an enhancement of biological performance was detected in the case of alginate crosslinked with calcium gluconate, which could be explained by the larger pore sizes, as qualitatively observed by scanning electron microscopy. In contrast to calcium chloride, calcium gluconate provides the slow release of small amounts of Ca2+, and this leads to the formation of a more homogeneous matrix with larger pores. This type of matrix probably promotes a more efficient diffusion of oxygen and nutrients within the alginate matrix while allowing a higher cell–cell contact, and thus might present advantages for application in adipose tissue engineering. In another example,

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Fig. 6. Methacrylate alginate hydrogels after 4 weeks in vivo: (A) ionically and (B) photocrosslinked hydrogels. After 4 weeks in vivo, all hydrogels were encapsulated in a thin fibrous capsule consistent with a mild foreign body response. However, ionically crosslinked alginate hydrogels underwent dissolution while photocrosslinked hydrogels remained intact, with no apparent changes in structural integrity and shape. Arrows indicate location of the hydrogel [111]. Copyright 2009 Elsevier Ltd.

Table 3 Examples of injectable alginate hydrogels proposed for cell delivery in adipose tissue engineering applications (these data are not intended to be all-inclusive). Alginate supplier





Ref.

FMC BioPolymer

Ultrapure MVG; 40% M/60% G; Irradiated, oxidized and RGD modified; 3 LMW:1HMW LMW

2

IG with CaSO4

hADSC (20)

[49]

1

hADSCs (n/a)

[151]

hADSCs (5)

[99]

APC (3T3-L1 cells)

[112]

Mice ADSCs (20)

[152]

n/a FMC BioPolymer FMC BioPolymer n/a

200,000–300,000 g mol RGD modified LF 10/60 Protonal RGD modified

2 3

EG with CaCl2 and calcium gluconate (thin layer) IG with CaSO4 (disks, 8 mm diameter and 1 mm thickness) Photocrosslinked

Low viscosity

1.2

EG with CaCl2

–1

ADSCs, adipose-derived stem cells; APC, adipose progenitor cells; EG, external gelation; G, guluronic acid; HMW, high molecular weight; IG, internal gelation; LMW, low molecular weight; M, mannuronic acid; n/a, not available.

photocrosslinked methacrylated RGD–alginates with different mechanical properties were investigated as 3-D hydrogel matrices for a mouse pre-adipocytic cell line (3T3-L1) [112]. Hydrogels with different stiffness were obtained, with representative elastic moduli of different physiological conditions: 3.3 (physiological), 7.9 (intermediate) and 12.4 kPa (pathological). The authors observed that increased matrix rigidity enhanced cell proliferation and angiogenic capacity, while adipose differentiation was inhibited. Collectively, the obtained results highlight the role of matrix stiffness in the design of biomaterial scaffolds for adipose tissue engineering. Jing and co-workers [152] isolated ADSCs from mice that were differentiated into the adipogenic lineage in vitro and then entrapped within 1.2 wt.% alginate crosslinked with calcium chloride. The cells–alginate system was then implanted subcutaneously and new adipose tissue was formed after 8 weeks, demonstrating the potential of combining predifferentiated cells and alginate for adipose tissue engineering. Kim et al. [49]further demonstrated the potential of cryopreserved hADSCs entrapped within 2 wt.% RGD–alginate hydrogel to produce living adipose tissue in vivo via a minimally invasive way, since crosslinking occurred in situ due to the presence of CaSO4. 4.5. Cardiac regeneration Cardiovascular disease is one of the most common causes of death in the world [153]. In the majority of cases heart failure,

and consequently death, are due to myocardial infarction (MI) associated with the left ventricle (LV) [154]. In this clinical area, cardiac tissue engineering offers promising post-MI treatments, based on the use of injectable alginate hydrogels [155,156]. In 2008, Landa et al. [157]demonstrated for the first time that in situ forming 1 wt.% alginate hydrogel improves left ventricular remodeling and function in a rat model of recent and old infarcts. This group later tested the same system in a large animal model of MI (swine) [158]. They were able to show that an intracoronary injection of a calcium crosslinked alginate solution was feasible in a larger infarcted area, and that alginate diffuses from the leaky coronary microvasculature and deposits in the infarcted myocardium. In an attempt to improve alginate function, alginate was further modified with adhesion peptides (RGD and YIGSR (tyrosine-isoleucine-glycine-serine-arginine)) [159]. However, when injected, a reduction in the beneficial effect of the alginate in the infarcted area was observed. A possible explanation for this may lie in the possible physical changes that occurred due to alginate modification (e.g. increased viscosity leading to reduced biomaterial spreading). Using a different in situ approach, Yu et al. [160]were able to deliver 1.5 wt.% alginate, grafted or not with RGD and calcium chloride using a Duploject (Baxter) applicator, which provided a simultaneous mixing and delivery of both components into the infarcted area of the LV, in a chronic rodent model of ischemic cardiomyopathy. The authors observed that both alginates (modified and unmodified) could reshape the aneurysmal LV (as illustrated


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Fig. 7. Effect of injected alginate in a chronic rat infarct model. Trichrome staining at 5 weeks post-injection. (A) PBS-injected and (B) alginate-injected heart. The trichrome staining demonstrated that the aneurysmal LV geometry of alginate-treated animals resembled normal control hearts, in contrast to the thin-walled, dilated LV of the PBStreated animals. The black arrow in (B) indicates residual alginate in situ. Scale bar = 1 mm [160]. Copyright 2008 Elsevier Ltd.

in Fig. 7) and improve LV function, and that RGD–alginate could particularly enhance angiogenesis. This group later adopted a different strategy by developing a cell delivery system where hMSCs were entrapped and then injected within RGD–alginate microbeads smaller than 100 lm [51,161]. The RGD–alginate microspheres induced angiogenesis and delayed the negative remodeling of the LV by preventing the infarct wall thinning and chamber dilation after an infarct. Moreover, RGD–alginate microbeads were able to enhance the retention and survival of transplanted stem cells. Interestingly, the authors were unable to observe in vivo any differences between RGD–alginate alone and with cells. It is noteworthy that they observed hMSCs within RGD–alginate microspheres in the MI after 1 week, demonstrating the effectiveness of these microspheres for the delivery of cells to the myocardium. In this context it is worth mentioning that, following the encouraging results in animal models, an alginate hydrogel has become the first injectable acellular biomaterial to enter clinical trials for treating MI. In fact, the ongoing clinical trial using this biomaterial, with the market name of IK-5001, is currently recruiting patients to a phase II clinical trial [51]. 4.6. Regeneration of other tissues and organs Alginate hydrogels are also being investigated for their ability to enhance regeneration in other tissues and organs. However, their application as injectable cell delivery vehicles is still poorly explored. In the central nervous system, the benefits of other hydrogels (matrigel, collagen, fibrin, agarose) to deliver cells has already been demonstrated in comparison with preformed scaffolds and cells alone [2,162]. However, only few studies using alginate have been reported to date. For example, alginate microbeads have been explored as an entrapment system to facilitate the neuronal differentiation of embryonic stem cells (ESCs), which could be further scaled up for the production of large numbers of differentiating cells [163]. For that purpose, a range of alginate concentrations (1.2, 1.7, 2.2 and 2.5 wt.%), with a final cell concentration of 5 106 cells ml1, were evaluated. With the right stimuli, the authors found that 2.2 wt.% alginate was the most conducive to ESC differentiation into astrocytes and neuronal lineage cells. In another example, Schwann cells over-expressing brain-derived neurotrophic factor (BDNF) entrapped in alginate beads were able to increase auditory neuron survival when transplanted into the cochlea of a deaf guinea pig [164]. This cell-based neurotrophin treatment could be considered a valuable option for the delivery of neurotrophic factors to reduce or prevent auditory neuron degeneration in sensorineural hearing loss. In a similar approach,

fibroblasts engineered to produce BDNF survived, grew and expressed bioactive BDNF while entrapped within alginate microcapsules, and were shown to be capable of successfully guiding the neurite growth of dorsal root ganglia [165]. In a different strategy, highly anisotropic alginate-based capillary hydrogels, introduced in a spinal cord lesion, were able to induce directed nerve regrowth without a major inflammatory response [57]. As previously mentioned, one of the first applications of alginate hydrogels in tissue engineering involved immunoisolation – more specifically, the transplantation of encapsulated pancreatic islets in an effort to cure type I diabetes. The purpose of these systems was to isolate the cells from the surrounding environment while producing insulin in diabetic patients. Despite its great clinical interest, these systems have not been deemed suitable for clinical applications yet due to several problems, mainly related with immunogenicity [166]. Since the 1970s several efforts to improve this type of system have been made, namely in what concerns the alginate composition and purity, the use of different coating materials and cell sources, as well as the site of implantation [166–168]. Recently, 3-D alginate matrices have demonstrated great potential in the development of a bioartificial liver [169]. Cells from a human-derived liver cell line (HepG2) were entrapped in 2 wt.% alginate and cultured in a fluidized bed bioreactor. Cells were reported to remain viable at high cell densities and proliferate to form compact cell spheroids with extensive cell-to-cell contacts and adequate cell function (lactate production, glucose and oxygen consumption, gene and protein expression). The authors hypothesized that HepG2 cultured in an optimized form could be used to temporarily improve the function of deficient liver to allow enough time for its repair or for transplantation. In another study, alginate matrices were used to evaluate the effect of matrix stiffness on the behavior of a human hepatocyte cell line, C3A [170]. As demonstrated, hepatocytes stopped growing and de differentiated in stiff hydrogels (3 wt.% alginate), whereas in soft hydrogels (0.7 wt.% alginate) cells maintained their differentiated phenotype, expressing higher xenobiotic metabolism and increased albumin secretion. Overall, the understanding of the 3-D mechanical effects of the microenvironment on C3A cells may provide a platform for the future design of liver tissue engineering. Other cells from a human hepatoma cell line (Huh-7) were successfully immobilized within alginate beads for 7 days, and throughout this period cells proliferated and organized into aggregates [91]. Hepatocytes established a 3-D architecture with cell polarity, cell junctions and bile canaliculi. The functionality of these cells was assessed by the production of albumin and the exhibition of CYP1A (one of the major drug metabolizing enzymes). Moreover, the authors observed the expression of specific receptors of hepatitis C virus,

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which suggests that this system could be further used for the development of a physiologically relevant model for in vitro viral studies. RGD–alginate scaffolds have also been explored in the field of skeletal muscle tissue engineering as vehicles for myoblasts [171,172]. Although not injectable, these multifunctional scaffolds were able to release cells and growth factors simultaneously, and consequently accelerate the muscle regeneration process. In another example, a covalently crosslinked shape-memory alginate scaffold combined with IGF-1 and skeletal muscle cells also showed great potential for cell and growth factor delivery by a minimally invasive technique for the regeneration of skeletal muscle [6]. Scaffolds were able to release IGF-1 over several days in vitro and allowed skeletal muscle cells to survive, proliferate and migrate outwards over 28 days. Alginate hydrogels combined with cells have also been investigated for skin regeneration strategies. Since skin is composed of two different layers, one rich in keratinocytes (epidermis) and the other rich in fibroblasts (dermis), a bilayered structure should offer a more effective treatment. Considering this, Hunt and colleagues [95]immobilized fibroblasts within alginate beads and then keratinocytes were cultured on the surface, forming a bilayered composition. On the one hand, the viable fibroblasts within the matrix did not proliferate but were able to secrete ECM, which would be able to replace the alginate matrix after degradation. On the other hand, keratinocytes were able to grow uninhibited by fibroblast overgrowth to form a stratified epidermal layer. Although in vivo studies need to be addressed, this study reveals the potential of alginate for skin regeneration. In a more general approach, due to their immunomodulatory capacity, MSCs entrapped in alginate were able to preserve their anti-inflammatory function and provide a controlled delivery vehicle to attenuate inflammation and, consequently, promote tissue repair in vivo [173–176]. 4.7. Vascularization In attempting to achieve long-term clinical success of engineered tissues, it is now widely recognized that vascularization should be promoted [177]. In point of fact, to mimic the functionality and complexity of natural tissues, such as the heart, liver, kidney, muscle, nerve and bone, therapeutic stimulation of vessel growth is important during the regeneration process. However, this is still ignored in many tissue engineering approaches. The use of alginate-based strategies for therapeutic vascularization

has been recently reviewed by Gandhi and colleagues [178], and some of the previously mentioned studies with injectable alginate matrices have taken this into account with good outcomes [17,74]. One of the approaches proposed focused on stimulating vascularization via cell-released soluble factors. For example, the local delivery of MSCs in alginate microspheres per se enhanced the vascularization of tissue engineered constructs, since these cells secreted different angiogenic factors that were able to diffuse out of the matrix into the surrounding environment [179]. hMSCs delivered using this technology were shown to augment the restoration of vascular supply in ischemic tissue, highlighting the potential of alginate-entrapped hMSCs as a strategy to enhance paracrine-mediated vascular recovery after hindlimb ischemia [179]. In another study, a subpopulation of mouse bone marrowderived cells encapsulated in 1.5 wt.% alginate was able to influence venular remodeling in a mouse dorsal window chamber skinfold model (a circle of tissue that is dissected and covered with a round coverslip) [93]. Moreover, the in vivo delivery of rabbit bone marrow-derived MSCs entrapped in 2 wt.% alginate was demonstrated to improve hindlimb perfusion in an endovascular model of peripheral arterial disease [92]. An alternative strategy has been the use of alginate matrices as vehicles for endothelial-like cells that can themselves be incorporated into new blood vessels. In fact, the use of cell-based therapies involving the use of mature endothelial cells or progenitor endothelial cells has proven to be instrumental in the treatment of cardiac infarction, for example, as well as in other situations where vascularization is deficient [153,180,181]. Although these cell sources have not yet been fully explored in alginate-based systems, a few studies have been performed with encouraging outcomes. For example, an in situ formed 1 wt.% RGD–alginate hydrogel matrix with a bimodal distribution of high and low molecular weight and low stiffness was shown to recreate an adequate 3-D microenvironment for vascular cells, survival, migration and proliferation in vitro. As depicted in Fig. 8, cells within RGD–alginate were able to spread and form a cellular network, in contrast to cells within unmodified alginate [18]. Moreover, this specific matrix promoted the organization of cells in 3-D networks, which constitutes a step forward in the development of an injectable strategy to enhance vascularization. In a more complex approach, the co-entrapment of different cell types might be even more effective. In a recent in vitro study, the co-entrapment of endothelial cells with other cell types, namely ADSCs, in a photocrosslinked RGD–alginate were shown to be capable of accelerating vascularization. This effect was explained by the increasing proliferation and tube-like structure

Fig. 8. Endothelial cells within alginate hydrogels. HUVECs were immobilized with 1 wt.% alginate with a cell density of 2 107 cells ml1 of alginate: (A) without RGD petides and (B) with RGD peptides 48 h post-immobilization. Cells were stained with Alexa-fluor 488-phalloidin (F-actin) and counterstained with Hoechst 33342 (blue), and visualized under a Leica SP2 AOBS SE laser scanning confocal microscope (Leica Microsystems). The images represent an overlay of different plans [18].


formation of endothelial cells in response to the elevated concentration of pro-angiogenic factors (e.g. vascular endothelial growth factor) secreted by the co-cultured cells [112]. In in vivo studies, Silva et al. [4,182]used RGD–alginate scaffolds to create a depot of human vascular progenitor cells for the treatment of ischemic murine hindlimb musculature. The authors observed a reverse hindlimb ischemia and an increase in the number of vessels compared to direct cell injection. These results demonstrate the importance of the delivery approach to the sustained release and repopulation of the surrounding tissue, although the system was not injectable.

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Acknowledgements This work was financed by FEDER funds through the Programa Operacional Factores de Competitividade – COMPETE, and by Portuguese funds through FCT – Fundação para a Ciência e a Tecnologia, in the framework of the projects Pest-C/SAU/LA0002/2011 and BIOMATRIX (PTDC/SAU-BEB/101235/2008 and FCOMP-01-0124FEDER-010915). S.B. is the recipient of a post-doctoral fellowship from FCT-POPH (SFRH/BPD/80571/2011). C.B. has a research position funded by FCT-POPH-FSE (Ciência 2008). Appendix A. Figures with essential colour discrimination

5. Conclusions and perspectives Despite the recent advances and promising results of cell-based therapies, substantial challenges remain in controlling the fate of transplanted cells. The success of these strategies might be significantly improved with the use of adequate cell carriers. Alginatebased matrices stand out as ideal materials for such applications. Alginate can form tunable and versatile hydrogels that are particularly attractive due to their easy and mild gelation, attained via a variety of crosslinking approaches. These highly hydrated and permeable matrices offer unique advantages for the development of 3-D cellular microenvironments. To increase their ECM-like features, alginates that are intrinsically ‘‘bioinert’’ can be modified with a variety of biological cues to promote cell–matrix interactions. Another distinguishing characteristic of these materials is their potential injectability, essential to deliver cells directly into the site of injury through a minimally invasive approach. As pointed out, injectable alginate systems are already having a remarkable impact in many biomedical applications. The examples described in this review illustrate their potential for therapeutic cell delivery, in different areas of the tissue-engineering field. In some of the studies, the success of this strategy has been clearly demonstrated. In this context, alginate hydrogels play two important roles: they not only provide a protective vehicle for transporting cells into the body, but also provide a temporary ECM-like supportive matrix that guides 3-D cell organization and might even modulate cell fate, promoting cellular integration and consequently tissue repair and regeneration. Despite these promising insights, it is also clear that alginatebased materials still have to evolve considerably to become the ideal matrix for tissue regeneration. The alginate hydrogels presently used clinically for wound healing, heat burns, acid reflux and weight control applications play quite a passive role. Although, in tissue regeneration, important steps are being made in the field of cardiac regeneration, in the other areas described there is still a considerable gap between research and the clinical application. One of the aspects that could contribute to decreasing this gap would be better characterization of the alginate sources for better comparison between studies. There were several relevant works available where the alginate used has not been characterized in terms of molecular weight, amount of G and M units, endotoxins content, peptide concentration or even alginate concentration. In line with the development of the field itself, a burgeoning number of cutting-edge discoveries are leading to the design of improved alginate hydrogels, with several interesting properties in terms of injectability, in situ crosslinking strategies, mechanical properties, biodegradability and biomimetic behavior. At the same time, there has been an important investment in the exploration of new cell sources and complex cell–cell and/or cell–drug combinations. All of these efforts and their achievements will certainly contribute to improving the current status of cellular therapies in the near future.

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