Biomaterials have been widely used as scaffolds for cells transplanted to the central nervous system (CNS). They have been synthetized from either natural or synthetic materials (Table 13.2). Natural sources include proteins such as Matrigel [18], collagen [49, 50], fibronectin [51], and fibrin [52], polysaccharides such as chitosan [53], agarose [54], alginate [55] and methylcellulose [56], hyaluronan (HA) [14, 57] and acellular tissue matrix [58]. Many of these molecules are physiologically found in the extracellular matrix (ECM), and therefore their use has the advantage of promoting cell signaling through their binding sites for mammalian cells. In addition, chitosan and agarose have available functional groups that facilitate chemical alterations [59]. From a translational perspective, natural molecules are clinically used as dermal fillers, lubricants, wound sealants, and surgical sponges [60] which facilitate their approval for use from safety and regulatory standpoint.

On the other hand, by assembling a list of design criteria for certain scaffold applications, one can produce synthetic polymers with defined compositions, polymerization and degradation rates, and mechanochemical properties. Synthetic molecules are produced in consistent ways that reduce data variability. Moreover, by being biologically inert, synthetic molecules decrease interaction with inflammatory cells in the brain and improve implant biocompatibility. Poly-lactic-co-glycolic acid or PLGA [61, 62], poly (acrylonitrile)/poly(vinyl chloride) or PAN/PVC [61, 62], oligo (ethylene glycol) fumarate or OPF [65], poly(N-isopropylacrylamide) or PNiPAAM, and polyethylene glycols (PEGs) [66] are common examples of synthetic molecules used as scaffolds in the CNS. It is noteworthy that polyesters such as PLGA have been used as drug delivery vehicles, orthopedic fixation devices, and absorbable sutures [67, 68], and they have established a history of clinical safety which accelerates their approval in clinical trials.

It is important to know that limited chemical and biophysical interactions of the aforementioned molecules in the context of biological complexity of NSPCs and poststroke brain tissue often necessitate a combination of natural molecules, synthetic polymers, or a conglomerate of both to compose the optimized scaffold product for the best cell survival and integration following transplantation in stroke.

Biological scaffolds can be further categorized based on their micro- and nanostructure. The earliest attempt to use scaffolds in order to improve cell viability goes back to the 1990s when scientists noticed adrenal chromaffin cells demonstrate a poor cell survival and efficacy upon transplantation to the experimental lesions of Parkinson’s disease (PD). To improve the outcome of cell therapy, they injected cells attached to collagen-coated dextran or glass beads as microcarriers [69, 70]. This form of “particulated” matrix promoted cell survival and corrected for dopamine deficits in the striatum. Another particulated matrix, gelatin beads, produced under the name of Spheramine®, was shown to promote survival and function of human retinal pigmented epithelium cells in an experimental model of PD [71]. Alginate beads were able to protect encapsulated cells against Huntington’s disease where survival of transplanted rat choroid plexus cells was improved following quinolinic acid (QA) injection [72]. Most relevant to our review, a scaffold of PLGA microspheres promoted integration of NSPCs following injection into the stroke brain cavity [36]. These studies demonstrated the importance of a viable three-dimensional (3D) matrix in transplanting cells to the brain lesions.

Besides scaffolding for cell transplantation, microspheres have been widely used to bypass the blood–brain barrier and deliver drugs, proteins, peptides, viral constructs, DNA, small interfering RNA (siRNA), and other therapeutics into the CNS. Based on their chemical properties, microspheres have a degradation rate that determines releasing pace of the compound encompassed within microparticles. This property ensures a reliable sustained release of an encapsulated molecule. Having multiple microspheres with different degradation rates gives a powerful tool to sequentially release growth factors and optimize tissue repair. For instance, to maximize proliferation of NSPCs as well as neuroprotection following stroke, it may be important to deliver sequential factors, such as epidermal growth factor (EGF) followed by release of the cytokine erythropoietin (EPO). The Shoichet group has achieved this goal by coating EPO-containing nanoparticles with a layer of poly(sebacic acid) that retards the release of EPO for about 7 days [56] and therefore optimizes tissue repair following stroke.

The slow-release capability of microspheres can be combined with cell attachment on their outer surface (Fig. 13.). In fact, microspheres could be sintered at low temperature to form a stable 3D microenvironment for cell transplantation [15]. Survival of transplanted cells attached on a 3D matrix of microspheres could be further augmented when microspheres gradually release survival and differentiation factors and thereby exert synergistic effects of adhesion and growth factor signals on cells. This strategy has been employed in treating PD where glial-derived neurotrophic factor (GDNF)-loaded PLGA microspheres promoted the efficacy of embryonic ventral mesencephalon dopaminergic cells attached on their exterior surface [73].

There is a multitude of factors contributing to low cell yield following transplantation to the stroke infarct core. We will show that biomaterial scaffolds improve cell survival by targeting mechanisms of cells death. An immediate example of technical difficulties that lead to poor transplant survival is seen in the injection of cells as a thin suspension of cells injected into the brain. The tissue retains some but inevitably restricts part of the injected volume and part of cell suspension flows back to the pial surface and may diffuse into subarachnoid space. Mixing cells with biopolymer hydrogels that have viscosity equal or above the infarcted tissue improves retainability and targets entire cell population to the focus of stroke [14, 88]. This not only improves number of transplanted cells, and therefore promotes recovery, but also avoids distribution of stem/progenitor cells across the CNS, a safety concern in stem/progenitor cell clinical trials.

The secretory profile of stem/progenitor cells and their differentiation, both mechanisms contributing to neural repair after stroke, are affected by their surrounding microenvironment. This includes the ECM composition of proteins, adjacent cells, humoral growth factors (such as brain-derived neurotrophic factor or BDNF, ciliary neurotrophic factor or CNTF, NT-3, PDGF, GDNF, and nerve growth factor or NGF), and biophysical properties of their neighboring matrix. These elements are present in the unique niche of neural stem or progenitor cells to cue them for self-renewal and differentiation. Hence, any successful application of biologic scaffolds for cell delivery should consider mimicking a similar combination of factors.

A seminal study by Discher’s group [105] revealed that focusing on biochemical factors as the sole regulators of stem cell differentiation is a simplistic view. They showed that the stiffness of the substrate matrix influences the differentiation of mesenchymal stem cells. More interestingly, the stiffness of tissues drives differentiation in resident populations of stem/progenitor cells by a very characteristic and specific manner, as when cultured on substrates mimicking bone, muscle, or brain in their stiffness, mesenchymal cells were inclined to make mature osteoblasts, myocytes, or neurons, respectively [105]. This indicates how tissue elasticity, besides biochemical factors, provides signals for renewal of specialized cells. Other groups have found the same concept in regulating NSPCs (Fig. 13.4). NSPCs’ proliferation and self-renewal increase on hydrogels below 10 kPa (compressive modulus) and maximize around 3.5 kPa [110]. Neural differentiation of NSPCs, however, requires lower elasticities below 1 kPa that in fact resembles the mechanical properties of the brain [81]. Hydrogel scaffolds with their flexibility in degree of cross-linking offer a very powerful tool to optimize self-renewal and differentiation. For instance, NSPCs transplanted in a 3.5-kPa matrix will mount a proliferative response and following a slow degradation of hydrogel biopolymer and descent of stiffness to < 1 kPa, their self-renewal slows and NSPCs start differentiating to neuron. Besides differentiation, other proregenerative functions of stem/progenitor cells such as their secretory profile are also controlled by matrix stiffness [111].

The incorporation of ECM proteins, such as collagen [112, 113], laminin [114], and fibronectin [112], has been used to mimic stem/progenitor cells niche and promote their differentiation. Stem/progenitor cells start to modify their microenvironment after transplantation and secrete their own ECM. Using cross-linkers amenable to cleavage by MMPs facilitates remodeling of hydrogel scaffolds by their encapsulated stem/progenitor cells [115] and promotes their differentiation and axon extension. Differentiation could also be promoted by other biophysical factors that affect cell–substrate interactions, such as micropatterned surfaces [116], addition of electrically charged monomer methacrylate [117], and forcing into distinctive cell morphologies [118].

Biopolymer matrices have been employed to present a range of growth factors to the population of encapsulated stem/progenitor cells and drive their proliferation and differentiation. This includes strategies to immobilize growth factors or incorporate them in a slow-releasing reservoir. Agarose-immobilized PDGF or chitosan-immobilized interferon-γ has led to differentiation of NSPCs to oligodendrocytes [119] or neurons [120], respectively. Alternatively, a sustained provision of NT-3 [121] or differentiation factor dibutyryl cyclic-adenosine monophosphate [122] from PLGA microspheres have induced neural differentiation in stem/progenitor cells. Differentiation has also been augmented by hydrogel delivery of CNTF, NT-3, PDGF, GDNF, or NGF [55, 104].

In the future, encapsulating cells within biopolymer scaffold with a selected set of factors to drive differentiation may be an alternative to transplanting pre-differentiated stem/progenitor cells. Instead, a scaffold with finely crafted elasticity, surface charge, micropattern structure, and ECM proteins composition together with timely release of appropriate proliferation and growth factors would be a viable approach to direct differentiation of encapsulated stem/progenitor cells for a prolonged time after transplantation.

Any nondegradable substrate in the brain leads to a constant activation of an inflammatory response that isolates the “foreign object” by forming an astrocytic scar [124]. This scar is a counter-integration phenomenon since it serves as a barrier for cell migration and axon growth through the scaffold-tissue interface [125]. Biopolymer hydrogels have several characteristics that improve their biocompatibility and integration in the brain.

Mechanical trauma during scaffold implantation causes an immediate inflammatory response that ultimately leads to the foreign body response against implants [126, 127]. This could be avoided in hydrogels by employing minimally invasive methods of gel injection. In addition, the high water content of hydrogels further improves their biocompatibility [61]. The size and shape of hydrogels are additional biophysical parameters that determine surface area available to inflammatory cells and therefore affects biocompatibility [127, 128]. The stiffness of the implanted matrix is another physical factor contributing to prolonged inflammation and implant isolation. Astrocytes and microglia, the main cells involved in inflammation in the brain, are primed by stiff substrates to produce proinflammatory molecules such as IL-1β and toll-like receptor (TLR)-4, respectively [123]. Implantation of hydrogels with elasticities higher than the brain’s values incites inflammation in the early phase following implantation that leads to astrocytic scar in chronic phase [123]. This indicates scaffolds similar in stiffness to the brain will optimize differentiation in encapsulated stem/progenitor cells [110] as well as promoting integration to the surrounding brain tissue. Another solution to render hydrogels biocompatible is to make them degradable in the brain. Hydrogels provide support for encapsulated cells early after implantation and following growth, differentiation, and deposition of their own matrix, hydrogel is degraded by encapsulated or host tissue cells. Examples are HA [88] and collagen [66] gels that are degraded by cellular hyaluronidase and collagenase, respectively. It is important, however, to determine that hydrogel degradation products are nontoxic and cleared away.

In the pursuit of more biological solutions to improve biocompatibility, hydrogels have been loaded with a variety of anti-inflammatory factors (for details see “Trophic Support and Protection from Inflammation”). Chondroitinase ABC, a digestive enzyme for CSPG molecules, has also been used in implantation of neural stem cell (NSC)-loaded poly-ε-caprolactone scaffold to promote migration of NSCs into the injured CNS tissue [129]. In addition, high molecular weight HA gels through their anti-inflammatory effect can attenuate scar formation and improve integration of encapsulated cells [88]. This indicates biopolymer scaffolds, through their physical, chemical, and biological properties, attenuate inflammation and scar formation, and therefore promote integration of transplanted cells to the host brain tissue.

Oxygen and nutrients diffuse up to 150–250 µm from capillaries [130]. Therefore, any poststroke tissue reconstruction beyond these dimensions necessitates establishment of an environment with accessibility to blood perfusion. This requires formation of blood vessels reaching into the transplanted matrix. In fact, there is a correlation between transplanted cell survival and extent of new vessel formation [131]. Angiogenesis involves activation of endothelial cells by proangiogenic factors, such as vascular endothelial growth factor (VEGF), PDGF, FGF, hepatic growth factor (HGF), angiopoietin-1, and transforming growth factor-β or TGF-β. Endothelial cells then branch off from existing capillaries, proliferate, and migrate into the transplanted matrix [132, 133]. For successful angiogenesis, the transplanted matrix should support survival and proliferation of endothelial cells. In addition, the matrix has to allow remodeling by endothelial cells, forming cord-like structures which will mature into capillaries, and subsequently veins and arteries.

Hydrogels as highly hydrated and porous substrates are supportive for endothelial cell migration. A network of interconnecting and communicating pores facilitate angiogenesis and obviates the need for matrix digestion by endothelial cells. Angiogenesis could be further promoted by enriching the hydrogel environment and introducing elements of ECM matrix to the hydrogels, such as fibronectin [134] or fibronectin-derived synthetic polypeptide motifs [135]. Incorporation of slow-releasing proangiogenic factors, such as VEGF or bFGF, is another means to potentiate formation of new vessels inside a hydrogel matrix [136]. Migration of endothelial cells and remodeling of matrix involves secretion of MMPs, such as MMP-2 and MMP-9. Therefore, hydrogels with motifs recognized and cleaved by MMPs will have an additional proangiogenic effect. This could be achieved by including MMP-sensitive cross-linkers. A study combined both strategies by including VEGF tethered through a MMP-sensitive linkage and showed improved angiogenesis [118]. Some biopolymer hydrogels particularly support formation and stability of new vessels. HA gels are an example: While high molecular weight HA leads to endothelial cell quiescence, oligomeric HA, produced from degradation of HA gels, promotes proliferation and migration of endothelial cells by interacting with HA cells receptors, such as CD44, hyaluronan-mediated motility receptor (RHAMM), and TLR-4 [137]. In fact, HA hydrogel has been shown to promote angiogenesis upon implantation to the brain [138].

Inclusion of NSPCs adds to a proangiogenic benefit of hydrogels. NSPCs perceive lack of oxygen and nutrients by expressing the transcription factor hypoxia-induced factor-1 that leads to increased secretion of VEGF [139]. Secretory support of angiogenesis by NSPCs could be enhanced when they are encapsulated within a hydrogel with 40 kPa (compressive modulus) [111]; this is another example of how mechanical properties contribute to regenerative potential of cell scaffolds. NSPC proangiogenic effects can be further enhanced by inclusion of endothelial cells in a co-transplantation approach. In fact, this combination of cells encapsulated in PLGA gels has improved angiogenesis at the center of traumatic CNS injury and promoted reestablishment of the blood–brain barrier [140]. The application of biopolymer matrices with optimized mechanical (i.e., porosity and elasticity), biochemical (i.e., degradability, inclusion of ECM elements and growth factors), and cellular composition will enhance formation of new vessels, which in turn contributes to survival and integration of transplanted cells.