10.2.3.1 Matrix-Core/Shell Microcapsules

Microcapsules can be classified into three categories (Figure 10.2): matrix-core/shell microcapsules, liquid-core/shell microcapsules, and cells-core/shell microcapsules (or conformal coating) [29]. Matrix-core/shell microcapsules are produced by placing polymer/cell droplets in a gelling solution. Basically, there are two methods to prepare the matrix-core/shell microcapsules. One method produces covalently crosslinked capsules, which attain an equilibrium swelling state that depends on the polymer–water interaction parameter and the crosslink density [33]. This type of capsule has significant advantages in terms of stability; however, the toxicity of the reactants and the time required to complete chemical crosslinking can be incompatible with cell survival. The other method involves the production of “reversible” or “physical” gels where the constituent polymer networks are held together by various degrees of molecular entanglements and secondary forces including ionic charges, hydrogen bonding, or hydrophobic interactions. All of these interactions are reversible and can be disrupted by changes in physical conditions or application of stress [33]. This class of microcapsules has been well studied due to its several advantages that include the maintenance of the cells in an aqueous environment where they are in contact with a soft and biocompatible material and for the long-term release of therapeutic molecules. Despite ability of the hydrogel matrix to support facile cell growth, the in vivo stability of electrostatic and hydrogen bonds maintaining the structure of this type of microcapsule is limited and the mechanical resistance of capsules also remains a concern [34].

10.2.3.2 Liquid-Core/Shell Microcapsules

Many technical improvements in microcapsule design and production have been introduced over the years. For example, the hydrogel network found within a capsule can be tuned to promote cell growth and protein production by including a liquid or semiliquid environment that supports higher diffusion of gases and nutrients than possible in a gel [35,36]. As a result, cells show increased survival and protein production. Liquid-core/shell microcapsules were first manufactured by dropping a cell/gelling solution into a polymer bath and have been pursued with different types of hollow polymeric microcapsules such as polymersomes, multilayered capsules, and hollow microspheres. Methods to fabricate this type of microcapsule have grown increasingly sophisticated beyond the self-assembly of microgels in droplets [37] to now include layer-by-layer polyelectrolyte deposition [38], interfacial polymerization [39], precipitation by phase separation [40], surface polymerization [41], copolymer vesicles [42], multiphase microfluidics [43], and a combination of molecular self-assembly and precipitation [44].

10.2.3.3 Cells-Core/Shell Microcapsules

Both of the techniques mentioned above (matrix-core and liquid-core shells) typically start with a scheme to generate a controlled-size droplet, followed by an interfacial process to stabilize the droplet and to obtain a solid microcapsule membrane around the droplet. Aside from these long-established approaches, new techniques are emerging that overcome shortcomings of the existing methods. For example, core/shell microcapsules that can be generated by conformal coating of cells been proposed [45,46]. Conformal coating constitutes a special case of microencapsulation where the term describes a method of forming a barrier directly around a small cell mass or a small piece of tissue to eliminate unutilized space in a microcapsule core. By closely surrounding the cell mass with the encapsulation membrane, this approach provides improved mass transport between the capsule exterior and the cell mass and also minimizes implant size by increasing the effectiveness of cell packing. The thickness of the poly(ethylene glycol) (PEG) layer formed on the cell surface in this strategy can be as thin as several nanometers thereby drastically reducing total graft volume compared with conventional microcapsules.

10.3.2.1 Alginate

Alginates are a frequently employed biomaterial for cell immobilization due to their abundance, easy gelling properties, and apparent biocompatibility; the significant breadth of literature available (reviewed in detail elsewhere [17,92]) drives home how important this compound has now become for cell and tissue encapsulation. Despite continuing relevance to the present day, alginate was first discovered in the late nineteenth century by Edward Stanford [93] and commercial production began in the 1920s through the harvest and processing of cell wall components of various species of marine brown algae [94]. This material was used in the pioneering study by Chang in 1964 [16] that laid a foundation for encapsulation as a strategy for immunoprotection of transplanted cells and the subsequent groundbreaking work of Lim and Sun in 1980 [18]. In the three decades since then, alginate has become a ubiquitous matrix material for cell, tissue, and macromolecule encapsulation with applications in many aspects of tissue engineering and regenerative medicine.

Alginates are linear block polymers consisting of α-L-guluronic acid (G block) and β-D-mannuronic acid (M block). The gelation of alginate from a solution containing divalent cations results in the formation of the calcium junctions of GG-GG, MG-GG, and MG-MG between alginate molecules. Calcium and barium are the most commonly used as the complexation ions to produce the gelled structure because of their selectivity and cooperative binding to the G blocks. The G and M content varies with the alginate source and can affect gel properties such as mechanical strength, biocompatibility, and permeability [95]. The main limitation of using alginate as an encapsulation matrix is poor mechanical stability in the presence of monovalent and certain divalent cations commonly found in cell culture media or a physiological milieu, which can have an adverse effect on hydrogel structure by destabilizing the polyelectrolyte interactions involved in gel integrity [36]. To increase the stability and to reduce the wall permeability (for immunoisolation) of the microcapsules, alginate microcapsules are often coated first with a polycation and then with another layer of alginate. Various polycations have been used for this purpose including poly-L-lysine (PLL), poly-L-ornithine (PLO), chitosan, lactose modified chitosan, and photopolymerized materials [96].

Cell adhesion cues have been incorporated into the alginate network either through the addition of Arg–Gly–Asp (RGD), a cell adhesion oligopeptide [97,98], or by copolymerizing alginate with gelatin as culture time increased and complete degradation of the gel occurred in 5 weeks [62]. In recent work, Liu et al. [99] used crosslinked gelatin and alginate derivatives to prepare cell sheets and spherical tissues wrapped in living cell sheaths. The microcapsules not only supported growth of the enclosed cells into spherical tissues but also provided a cell adhesive outer surface. Light-activated gelation of alginate was recently employed to encapsulate rat bone marrow cells through the release of calcium chloride from light-sensitive liposomes suspended in the alginate–cell solution [100].

10.3.2.2 Chitosan

Chitosan, a derivative of the extremely abundant chitin polysaccharide, has structural characteristics similar to mammalian glycosaminoglycans (GAGs) and has attracted much attention as an alternative polycation in preparing microcapsules. Chitosan is a biocompatible, biodegradable, and nontoxic cationic linear polysaccharide composed of β-(1-4)-linked D-glucosamine and N-acetyl-D-glucosamine units and is obtained by the extensive deacetylation of chitin by hydrolyzing the amino-acetyl groups using NaOH [101]. Chitosan can form hydrogels by ionic interactions [102] or through chemical cross-linking [103] with the bifunctional agent glutaraldehyde and it degrades via enzymatic hydrolysis [104].

Due to its weak mechanical properties and lack of bioactivity, chitosan is often combined with other materials, such as poly(lactic-co-glycolic acid) (PLGA) [105] to achieve more desirable mechanical properties for encapsulation. Chitosan has also been used to coat capsules to enhance stability and modify permeability as an alternative to PLL because it is regarded as being more biocompatible [106]. Under dilute acid conditions (pH<6), chitosan is positively charged and water soluble, while at physiological pH, chitosan is neutral and hydrophobic, leading to the formation of a solidified, physically crosslinked hydrogel. The addition of polyol salts enables encapsulation of cells at neutral pH, where gelation becomes temperature dependent [102]. The addition of the adhesion-promoting RGD tripeptide has been necessary to promote cell attachment to chitosan, allowing myoblast and cardiomyocytes to be successfully encapsulated in crosslinked gel made from this polysaccharide [73]. In a separate study, polylysine was grafted to chitosan to enhance the microenvironment for neural cell growth. When fetal mouse cortical cells were encapsulated in chitosan-g-polylysine gels, neurite extensions were observed in the 3D gel cultures [107]. Examples that illustrate the versatility of chitosan include the use of this material for microencapsulation of human red blood cells [108], HepG2 cells [109], BHK-21 cells [110], chondrocytes [111], bone marrow stem cells [112], and PC12 and R208F cells [71].

10.3.2.3 Hyaluronic Acid

HA, or hyaluronan, is a natural GAG that is nonimmunogenic in its high molecular weight form and also nonadhesive [113]. HA is a linear polysaccharide composed of repeating disaccharide units of N-acetyl-D-glucosamine and D-glucuronic acid and is found in many tissues, such as skin, cartilage, and the vitreous humor. HA is a component of the ECM in mammalian connective tissues and participates in many important biological processes during wound repair [81]. Colloids of HA can be gelled by prior chemical modification of hyaluronan with thiols [114], methacrylates [115], or tyramines [116], or can be crosslinked in situ using formaldehyde or divinyl sulphone. One of the key advantages of using HA gels for tissue engineering is that their degradation can be mediated by hyaluronidase, an enzyme secreted by a multitude of mammalian cell types [117]. These properties make HA an interesting biomaterial for cell-based scaffolds and several cell types (chondrocytes, fibroblasts, and murine embryonic stem cells (ESCs)) encapsulated into either HA/alginate beads [118] or HA-microfabricated devices have been observed to proliferate and behave correctly [81].

10.3.2.4 Collagen

Collagen gels, as well as other mammalian-derived protein-based polymers covered in the next section, are effective matrices for cellular growth because these materials contain many cell-signaling domains present in the in vivo ECM [119]. Collagen type I is one of the most studied materials for tissue engineering in part because of its low immunogenicity [120]. An advantage of collagen gels is that they can be created through natural means without chemical modifications; the resulting collagen matrices, however, do have drawbacks including mechanical weakness and low resistance to the enzymatic degradation. Various strategies, therefore, have been pursued to improve the stability and mechanical properties of collagen, including chemical (glutaraldehyde and water-soluble carbodiimide) [121], physical (UV irradiation and dehydrothermal) [122], and enzymatic [123] crosslinking approaches. As an example of the promise held by collagen microcapsule technology, self-assembled collagen microspheres provide superior hyaline cartilage reproduction for the microencapsulation of human mesenchymal stem cells (hMSCs) based on the stability of these devices, their ability to be injected in vivo, and their provision of a protective, growth- and migration-supporting environment for hMSC both in vitro and in vivo [75].

10.3.2.5 Protein-Based Gels

Proteins are polymers of amino acids that form structures attractive for cell encapsulation including β-sheet-rich fibrils that can be nanometers in width and micrometers in length that also further self-organize and entangle to form 3D hydrogels under appropriate conditions [124,125]. As an alternative to natural proteins, synthetic peptides hold great promise in designing hydrogel carriers with a wealth of bioactive signals programmed directly into the hydrogel matrix. For example, one could envision incorporation of cell adhesion moieties, biochemical cues to promote tissue deposition, and specific enzyme-sensitive sequences for cell-mediated degradation [126]. Several groups have explored synthetic peptides that are relatively short ~15- to 20-mer amino acid sequences designed to self-assemble into a hydrogel for cell encapsulations. One example was an amphiphilic peptide containing positively charged lysine residues that was demonstrated to successfully encapsulate mesenchymal stem cells (MSCs) [127]. Another example is provided by elastin-like peptides (ELPs) used to encapsulate human adipose-derived adult stem cells [128] and as a preosteoblastic cell carrier for bone regeneration [129].

10.3.3.1 General Considerations

In general, natural polymers, including those discussed above, have attractive features that provide them with therapeutic relevance. For example, natural polymers, as well as their degradation products, are usually not toxic or carcinogenic, nor do they elicit adverse reproductive and developmental effects and in most cases entrapment of living cells in matrices made from natural polymers can be accomplished under physiological conditions. In addition to natural biomaterials, however, many synthetic polymers as well as inorganic compounds have also been explored for cell encapsulation [130]. The mechanical and chemical properties of synthetic polymers can be tailored for different applications without having the immunogenicity-related concerns of some naturally occurring polymers (e.g., low molecular weight HA), which constitutes an important advantage of these materials over their natural counterparts [130].

Although synthetic materials provide researchers with large flexibility in material design, one drawback is that they do not have an intrinsic mechanism for interacting with cells and cell attachment typically relies on nonspecific adhesion, which tends to be weak [131]. This limitation, which restricts the use of synthetic materials in applications that require defined control over cell–matrix interactions, can be overcome in some cases by functionalizing these materials with bioactive molecules such as the RGD tripeptide adhesion motif. A number of synthetic hydrogels, e.g., poly(ethylene glycol), poly(acrylic acid), poly(vinyl alcohol), poly(N-isopropyl acrylamide), and their derivatives have been developed and successfully employed for use in tissue engineering and encapsulation of cells [49,132,133]. PEG will be discussed in more detail below as a well-studied representative of synthetic polymers used for cell encapsulation.

10.3.3.2 Poly(Ethylene Glycol)

PEG has been the most widely used synthetic material to create macromeric polymers for cell encapsulation. PEG hydrogels can be formed via the photoinitiated chain polymerization of functionalized PEG macromers, specifically dimethacrylated PEG. Photoencapsulation within these gels has been shown to be cytocompatible with multiple cell types, including fibroblasts, chondrocytes [83], osteoblasts [85], mesenchymal stem cells [86], aortic valvular interstitial cells [87], and neural cells [88]. Even though PEG gels are inherently cell repellent [117], certain cell types such as chondrocytes and osteoblasts encapsulated in PEG hydrogels survive without the addition of any adhesion or biological-signaling molecules [134,135]. However, many cell types are attachment dependent and require cell adhesion sites to survive necessitating chemical modification strategies to incorporate various peptides or other signaling molecules into the gels that reduce their cell-repulsion behavior. For example, hMSCs encapsulated in PEG hydrogels require the addition of a cell adhesion component, such as the oligopeptide sequence RGD, for cell survival [136]. Alternatively, strategies have developed natural hydrogel environments by mimicking the native ECM, which not only promote cell–matrix interactions with the hydrogel but also sequester important growth factors that can augment tissue growth [137].

PEG hydrogels are very stable in vitro and are refractory to degradation in biological systems. This drawback—when biodegradation of an implanted material is advantageous—has been addressed by exploiting the versatile nature of PEG polymer chemistry. For example, the incorporation of copolymers enables the network structure and stability of PEG-based materials to be modified conveniently [138]. In particular, polylactic acid (PLA) and poly(caprolactone) (PCL) have been widely used in poly(α-hydroxy esters) for creating biodegradable PEG macromers for cell encapsulation because degradation profiles can be controlled by tuning chemistry and size of the degradable block [139]; in one example, Martens et al. [140] demonstrated that biodegradation time was enhanced when PEG hydrogels contained labile PLA linkages. In another strategy, the degradation of PCL-b–PEG-b–PCL hydrogels was temporally controlled by exogenously delivering lipase to match tissue growth for cartilage tissue engineering [141].

10.4.1.1 Cell Encapsulation Technologies were Pioneered by Efforts to Treat Diabetes

It has long been recognized that diabetic patients, who suffer from a plethora of secondary complications such as blindness, neuropathy, and end-stage renal failure, benefit from insulin treatment. Relevant to this chapter, difficulties in obtaining and administering insulin provided impetus to many of the historical developments in the field of cell encapsulation dating back to 1964 when Chang [16] proposed the idea of using ultrathin polymer membrane microcapsules for the immunoprotection of transplanted cells and introduced the term “artificial cells” to define the concept of bioencapsulation. Based on efforts started in the early 1970s that employed alginates for the immobilization of insulin-secreting pancreatic islet cells (PICs), which remained morphologically and functionally intact for as long as 4 months in vitro [18], this approach subsequently was implemented in vivo when it was used to xenograft islet cells in a rat model of diabetes [149]. Importantly, the microencapsulated PICs had a superior ability to maintain normoglycemia in animal models compared to nonencapsulated cells; for example, in a study conducted in streptozotocin-induced diabetic Wistar Lewis rats, encapsulated PICs maintained normoglycemia for 3 weeks compared to only 6–8 days with nonencapsulated cells [149]. More recently, Luca et al. [150] immobilized Sertoli cells in alginate-based microcapsules and transplanted the newly formed functional islets β-cells into nonobese diabetic mice thereby reversing the disease.

10.4.1.2 Expanding the Types of Cells Available for Treating Diabetes

Intensified insulin treatment, in particular, the delivery of insulin-secreting PICs as discussed above has proven promising for the treatment of type 1 diabetes [151]. Unfortunately, transplantation of whole human pancreases, isolated islets, or PICs into patients has been severely hampered by the scarcity of cadaveric human donor organs, which has mandated a search for insulin producing cells/tissue source alternatives. Taking one direction toward overcoming this hurdle, Abalovich et al. [152] performed a preclinical study investigating the potential of encapsulated pig islet cells for xenotransplantation. Type 1 diabetic dogs transplanted with encapsulated PICs experienced a significant reduction (20–80%) in insulin necessity after transplantation. Elliott et al. [144] showed that porcine islet cells survived and insulin production was maintained in a human patient 10 years after transplant of pig islet cells. In a second direction, recent progress in stem cell biology has provided promising leads for the conversion of progenitor cells into functionally competent, insulin-secreting cells. Currently, several studies have reported that MSCs [153], ESCs [55], and recently induced pluripotent stem cells (iPSCs) [154] can differentiate into insulin producing cells potentially useful for the treatment of diabetes. Tuch et al. [155] have proposed replacing pancreatic β-cells of diabetic recipients with pancreatic progenitors derived from human ESCs (hESCs).

10.4.1.3 Diabetes Leads the Way in Clinical Translation of Cell Encapsulation Therapies

Translation of microencapsulated cell systems, as evidenced by preclinical trials, now is under way (listed in full in Table 10.2) and has often focused on diabetes. Encapsulated islet transplantation was first performed in 1999 in a 38-year-old male patient who had severe diabetes [20]. In 2000, Elliott et al. [19] transplanted encapsulated porcine islets into diabetic patients and confirmed no PERV infection in patients. Calafiore et al. [21] subsequently performed human islets microencapsulation using alginate and transplanted them into 10 T1DM patients in Italy in an approach approved by the Italian Ministry of Health [21] and an Australia group transplanted barium-alginate encapsulated human islets into four T1DM patients without immunosuppression and followed them for over 2 years in a phase I study. Similarly, Tuch et al. [22] investigated the transplantation of barium-alginate microcapsules containing human islet cells in four type 1 diabetic patients and demonstrated the safety of this method, with little C-peptide—which reflects the level of blood sugars—detected, normal renal function, negligible cytokine release, and no major infection detected during the trial. Although insulin independence was not achieved, glycemic control was improved, with a reduction in insulin daily requirement. To quickly summarize clinical testing, a high level of safety has been demonstrated with an absence of adverse events in four phase I trials showing that intraperitoneally infused microencapsulated human islets can be considered safe for up to 3 years [22,156]; conversely, efficacy remains marginal.

10.4.1.4 Status of Clinic Translation

Important lessons have been learned and precedents set from efforts to clinically translate cell encapsulation therapies to treat diabetes; in particular, this approach provides protection from immunorejection and facilitates the long-term health and function of the transplanted cells. Although the (yet to be achieved) Holy Grail of specific induction of transplant tolerance would be ideal, in vivo testing has demonstrated cell encapsulation is a promising technology for achieving sufficient immunoprotection to allow β-cell replacement [157,158]. The early studies to date have confirmed that an encapsulation approach limits undesirable immunological responses to the nonhost islet cells and increases the access of cell replacement therapy to a broader patient base by minimizing the need for chronic administration of immunosuppressive drugs. Furthermore, the use of a macroencapsulation device will enable retrievability of the entire graft, which may be a desirable feature for products using hESC-derived cells as a potential safety benefit, allowing removal of the cells if complications do arise [159]. Overall, as the examples of clinical testing of cell encapsulation therapies outlined in the previous paragraph illustrates, hurdles remain to make this therapeutic approach reproducible and effective. On the other hand, a high degree of safety has been achieved, making this approach attractive to treat an expanded set of diseases; examples include cancer, CNS disorders, and liver disease as discussed next.

10.4.3.1 Parkinson’s Disease

Initial proof of concept data for treating Parkinson’s disease (PD) with encapsulated cells came from studies in rodents and primates that demonstrated that encapsulated PC12 cells could be implanted into the striatum and deliver L-3,4-dihydroxyphenylalanine (L-DOPA) and dopamine in sufficient quantities to induce behavioral improvements [171,172]. Encapsulated PC12 cells survived in vivo for up to 6 months, while maintaining a typical morphology and production of catecholamines [172]. Other cell types, including encapsulated bovine adrenal chromaffin cells, survived for 1.5 years with continued production of catecholamines and met-enkephalin [173]. The most widely employed cell-based approach for the treatment of PD has entailed the administration of encapsulated choroid plexus (CP) cells in conjunction with glial-cell-derived neurotrophic factor (GDNF) or vascular endothelial growth factor (VEGF) secreting cells [174–176]. In one study using this strategy, encapsulated GDNF-secreting fibroblasts were implanted into the striatum of 6-hydroxydopamine (6-OHDA) injured rats. Long-term expression of GDNF, up to 6 months, was detected along with behavioral improvement and tissue biocompatibility of the capsules [175]. In another study, porcine CP cells were encapsulated in alginate–PLO–alginate capsules and implanted into the striatum of 6-OHDA lesioned rats. Results showed that the presumed neurorestorative proteins secreted by the encapsulated cells improved motor behavior and nigrostriatal dopaminergic activity [177].

10.4.3.2 Alzheimer’s Disease

Alzheimer’s disease (AD) is a progressive disorder with a successive decline in cognitive functions. One early symptom is episodic memory loss. The progressive memory loss in AD patients is associated with neurodegeneration and dysfunction of the central cholinergic [34]. Positive effects of nerve growth factor (NGF) on cell survival and neurological functionality have been demonstrated in several animal studies. Initial studies demonstrated that implants of encapsulated NGF-secreting BHK cells prevented cholinergic neuron loss following aspiration of the fimbria/fornix. Control-implanted animals had an extensive loss (88%) of ChAT-positive cholinergic neurons ipsilateral to the lesion that was prevented by NGF cell implants (14% loss) [178]. Nagahara’s [179] study indicated that sustained local delivery of NGF to the cholinergic basal forebrain can arrest and even reverse the degeneration of the basal forebrain cholinergic neurons involved in the cognitive decline in AD. The immobilization of genetically engineered cells to produce VEGF, ciliary neurotrophic factor (CNTF) or glucagon-like peptide-1 (GLP-1) have also been employed for the treatment of AD. Spuch et al. [170] implanted microencapsulated fibroblast cells releasing VEGF into a transgenic (APP/PS1) mouse model of AD with degenerative alterations in the microvasculature. The study reported increased angiogenesis, decreased presence of Aβ and tau protein, less apoptosis, and protection of the cognitive behavior in the APP/PS1 mice after the implantation of the microcapsules.

When encapsulated cells secreting CNTF were implanted intra-cerebroventricularly into mice expressing the mutant amyloid precursor protein, a significant improvement in cognitive function occurred [180]. In other studies, encapsulated MSCs expressing GLP-1 were evaluated in an AD double transgenic murine model for their ability to GLP-1 decrease amyloid deposition and A-β-induced toxicity [181]. Results suggested a reduction in A-β-induced toxicity in vitro along with anti-inflammatory and neuroprotective properties. Most recently, the authors describe the first clinical trial with encapsulated cell biodelivery implants that deliver NGF to the cholinergic basal forebrain with the intention of halting the degeneration of cholinergic neurons and the associated cognitive decline in patients with AD [182]. Follow-up MRI at 3 and 12 months postimplantation showed no evidence of inflammation or device displacement. At 12 months, implants were successfully retrieved and low but persistent NGF secretion was detected in half of the patients.