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Advanced biomaterials for repairing the nervous system: what can hydrogels do for the brain?
Materials Today Volume 17, Number 7 September 2014
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Advanced biomaterials for repairing the nervous system: what can hydrogels do for the brain? Zin Z. Khaing1,*, Richelle C. Thomas2, Sydney A. Geissler3 and Christine E. Schmidt1,* 1
J. Crayton Pruitt Family Department of Biomedical Engineering, The University of Florida, United States McKetta Department of Chemical Engineering, The University of Texas at Austin, United States 3 Department of Biomedical Engineering, The University of Texas at Austin, United States 2
Newly developed hydrogels are likely to play significant roles in future therapeutic strategies for the nervous system. In this review, unique features of the central nervous system (i.e., the brain and spinal cord) that are important to consider in developing engineered biomaterials for therapeutic applications are discussed. This review focuses on recent findings in hydrogels as biomaterials for use as (1) drug delivery devices, specifically focusing on how the material can change the delivery rate of small molecules, (2) scaffolds that can modify the post-injury environment, including preformed and injectable scaffolds, (3) cell delivery vehicles, discussing cellular response to natural and synthetic polymers as well as structured and amorphous materials, and (4) scaffolds for tissue regeneration, describing micro- and macro-architectural constructs that have been designed for neural applications. In addition, key features in each category that are likely to contribute to the translational success of these biomaterials are highlighted. Introduction Recent advances both in our understanding of the nervous system and the availability of sophisticated biomaterials have significantly changed the landscape of potential strategies for repairing the nervous system. New imaging and staining techniques along with the development of novel materials have opened the possibility to directly design biomaterials tailored for a particular application. In this short review, recent advances in the use of hydrogels made from both natural and synthetic polymers are discussed (see Fig. 1 as reference) and their evaluations using in vitro and in vivo preclinical models are presented where applicable. This review is by no means a comprehensive review of biomaterials for nervous system repair: the readers are referred to a number of other excellent review articles for further studies [1–11]. This review introduces important obstacles to central nervous system (CNS) regeneration focusing on the unique characteristics of the CNS. Additionally, the injury environment and scarring are unique to the CNS in that a glial scar introduces a physical and chemical *Corresponding authors:. Khaing, Z.Z. ([email protected]), Schmidt, C.E. ([email protected])
barrier to regeneration. Understanding the specific requirements and taking them into consideration in designing therapeutic platforms are likely to result in the successful development of next generation biomaterials. Briefly, the blood barriers, the endogenous immune system within the CNS, and the mechanical properties of CNS tissues are discussed. The focus of this review is on the use of hydrogels as scaffolds to aid regeneration within the central nervous system (CNS) (i.e., the brain and spinal cord). In particular, the following systems and applications are highlighted: (1) hydrogels as drug delivery platforms, (2) hydrogels to modify the post-injury environment, (3) hydrogels for cell delivery and (4) hydrogels for tissue regeneration.
Features of the central nervous system to consider for biomaterial development There are a number of challenges that are unique to the CNS that must be considered for development of therapeutic biomaterials. First, the CNS is segregated from the circulating blood by the blood brain barrier (BBB) and the blood spinal cord barrier (BSCB). These barriers are made up of tight junctions between extracellular membranes of endothelial cells and astrocytes. In normal physiological 1369-7021/ß 2014 Elsevier Ltd. All rights reserved. http://dx.doi.org/10.1016/j.mattod.2014.05.011
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FIGURE 1
Synthetic and natural materials can be used to deliver small molecules or cells to the nervous system. Micro and macro architecture can also be introduced to these hydrogels to alter injury environments and direct cell behavior. Chemical structures of some common polymers are presented here. Abbreviations: PLGA – poly(lactic-co-glycolic acid), PEG – poly(ethylene glycol), PCL – polycaprolactone.
conditions, few molecules can passively move from the circulating blood to the CNS extracellular fluid. This makes it difficult, if not impossible, to deliver drugs and small molecule therapeutics to the CNS intravenously. It has been suggested that hydrogels made from both natural and synthetic polymers that can be intrathecally placed into localized areas show great promise to deliver therapeutics into the brain and spinal cord. Under normal conditions, the BBB and BSCB also isolate the CNS tissue from the circulating immune cells. Thus, the CNS is considered an ‘immune privileged’ organ. Major immune cells within the CNS include (1) microglia, which are resident immune cells that are distributed throughout the brain and the spinal cord, and (2) perivascular macrophages that are located in the capillaries. After injury or in response to disease, microglia, astroctyes, macrophages, oligodendroctyes and even neurons to an extent, can all respond and release inflammatory cytokines [12]. Therefore, for biomaterials to modify the injury environment, the materials must interact with the resident cells’ immune response in a positive manner [13]. The brain and spinal cord are some of the softest tissues in the body with compressive moduli around 2000 Pa [14–16]. Matching the mechanical properties of an implanted biomaterial to that of the host tissue can significantly affect the success of the implanted material in vivo [2,17]. In vitro studies of neural progenitor cell (NPC) differentiation showed that hydrogels that best matched the stiffness of the brain provided the most optimal results for neuronal differentiation [16,18]. Hence, determining the appropriate mechanical property is a key factor to consider when using biomaterials in the CNS.
Injury environment of the nervous system A major difference between the peripheral nervous system (PNS) and CNS is the capacity for peripheral nerves to regenerate; CNS axons do not regenerate appreciably in their native environment.
After injury in the CNS, macrophages infiltrate the site of injury much more slowly than they do in the PNS. This delays the removal of inhibitory myelin associated proteins from the injury site. Macrophage recruitment is also limited because cell adhesion molecules in the distal end of the injured spinal cord are not appreciably up-regulated. Additionally, astrocytes in the CNS become ‘reactive’, and produce glial scar tissue. Astrocytic response after injury is characterized by cellular hypertrophy, astrocyte proliferation, process extension, and increased production of the intermediate filament proteins glial fibrillary acidic protein (GFAP), vimentin, and nestin [19]. This response of astrocytes to injury is known as reactive gliosis. When an injury does not involve penetrating the dura mater, the scar tissue formed is mainly composed of astrocytes; however, when the dura is broken there is infiltration of meningeal fibroblasts in addition to reactive astrocytes [20]. Most glial scar tissue is thought to include, in addition to the reactive astrocytes, NG2+ oligodendrocyte precursor cells, meningeal cells, infiltrating macrophages, and activated microglia. Schwann cells from adjacent dorsal roots have also been found within the CNS scar tissue in experimental injuries where the dura was broken. The mature glial scar includes proteoglycans such as neurocan, phosphocan, versican, and brevican along with secreted proteins including Semaphorin 3A and 3D [21]. One notable component of glial scar tissue is chondroitin sulfate proteoglycans (CSPGs). CSPGs have been shown to inhibit axonal outgrowth in many neuronal systems [22]. In sites distant from traumatic injury, astrocytes can also become larger in size and transform into a more pronounced stellate shape. These ‘activated astrocytes’ have been known to produce soluble trophic factors that enhance the survival of neurons and glial cells in the vicinity of the astrocytes [23]. Therefore, the effects of activated astrocytes after injury on axonal regeneration and neuronal plasticity is unclear. Moreover, myelin-related glycoproteins have long been implicated in creating a non-hospitable environment for the 333
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severed axons of the CNS [24–27]. Combined, these factors constitute a formidable environment for regenerating axons in the CNS. Although the idea of simply inhibiting the formation of scar after injury may be an attractive option, it is important to note that formation of scar tissue after injury to the CNS encourages tissue homeostasis. Contusion spinal cord injury (SCI) studies completed in transgenic mice that targeted the depletion of mitotic astrocytes immediately surrounding the injury site support this idea. The results show an increased infiltration of inflammatory cells, cellular degeneration and increased lesioned area compared to control mice with a similar injury [28]. Therefore it appears that there are specific beneficial aspects of the glial cell response. Specifically the production of cytokines, growth factors, and other extracellular matrix (ECM) components that participate in re-establishing the BBB after injury have been shown advantageous to the post-injury environment. Simply inhibiting the formation of scar is not a viable option. Instead, strategies that will specifically antagonize the receptors for the inhibitors may be a better way to interfere with the growth inhibitory nature of the injured CNS microenvironment.
Hydrogels as drug delivery platforms As stated above, delivering molecules to the CNS is a challenging problem due to the presence of BBB and BSCB. Direct intrathecal injections into the extracellular fluid have been successful in some clinical [29–32] and experimental/pre-clinical settings [33,34]. Current research in this field involves examining biocompatible and bio-inert polymers as platforms for controlled release of bioactive molecules within the CNS for sustained drug delivery (Fig. 2a). It is important to note that, after injury to the brain and spinal cord, there is a temporary leak in the BBB and BSCB that allows more molecules and cells to enter the CNS from the circulating blood compared to normal physiological conditions [35]. However, the extent to which this ‘leakiness’ can be taken advantage of for cell and drug delivery is unclear. Hydrogels with matching mechanical properties to that of the nervous tissue (compressive modulus 2000 Pa) can be placed in the intrathecal space to maintain homogeneous mechanical landscape with the surrounding soft CNS tissue [16]. In light of this feature, hydrogels made from a number of natural and synthetic polymers have been examined for their ability to deliver therapeutics directly into the brain and spinal cord. Naturally derived polymers including fibrin [36–39], hyaluronan–methylcellulose (HAMC) blend [40–45], hyaluronic acid [46– 48], agarose [49] and chitosan [50] have been used to successfully deliver molecules to the CNS. For example, an injectable hydrogel system developed by the Shoichet group shows promise for intrathecal delivery of growth factors into the injured spinal cord. This group developed HAMC hydrogels that are safe, effective [51], and able to deliver trophic factors [52]; modified versions of these HAMC hydrogels can also be used as vehicles to deliver cells into the injured spinal cord [53]. In one case, the delivery of erythropoietin (EPO) into the intrathecal space using HAMC hydrogels after SCI in a rodent model resulted in reduction in lesion cavity size and had a higher neuroprotective effect compared to direct injection of EPO alone into the intrathecal space [52]. In other cases, HAMC has been successfully combined with nanoparticles 334
FIGURE 2
A schematic representation of the three fundamental ways in which hydrogels are being developed for therapies to repair the nervous system is shown here. (a) Hydrogels alone or as composite materials with microparticles have been used to create highly tunable delivery platforms for small molecule therapeutics into the CNS. (b) Hydrogels can also be used to deliver cells into the CNS. (c) Hydrogels have been utilized, with and without internal architecture, as scaffolds to deliver cells and repair injured CNS tissue.
resulting in composite biomaterials with highly tunable delivery profiles [42,54]. Fibrin-based hydrogels have also been successfully used for delivery of growth factors [39,36,55,56], scar inhibiting enzyme chondroitinase ABC [57], and progenitor/stem cells in combination with growth factors [58–60] in rodent models of SCI. Likewise, injectable, in situ-gelling forms of agarose have been used in combination with lipid microtubules loaded with bioactive molecules after SCI [49,61]. Bioactive molecules have been delivered to the CNS with chitosan-based hydrogels [62–67], microparticles [68–70] or nanoparticles [71–80]. Chitosan-based biomaterials in particular have largely been explored for delivery through the nasal route. This delivery paradigm is attractive since the nasal passage has a large surface area, porous endothelial membrane, high total blood flow and is readily accessible. Indeed,
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limited information on possible biological effects of degradation products from synthetic polymers [97,98]. For example, lactic acid and glycolic acid are known byproducts of PLGA. Recently, an intriguing article by Rinholm et al. suggests that exogenous application of lactic acid to cultured slices of developing mouse brain cortex can support oligodendrocyte development and myelination [99]. Future studies examining the presence and possible effects of degradation products from commonly used polymers, such as PLGA, are warranted when applied in vivo.
Hydrogels to modify the post-injury environment Hydrogels can be particularly useful in the CNS post injury environment by potentially serving as a substrate to re-establish tissue continuity. Prefabricated scaffolds have been implanted [100], but can cause secondary injury to the surrounding tissue during delivery [101]. Minimally invasive injectable hydrogels are particularly well suited for these types of applications. The liquid polymer can fill irregularly shaped void spaces and injections via needles to deliver the liquid polymer circumvent the BBB and BSCB. The physical properties of hydrogels that are relevant to consider in this post-injury environment are the porosity, chemical composition and mechanical properties. Hydrogel porosity is essential to stabilize the post-injury environment by permitting nutrient flow into and out of the scaffold. Porosity also affects cell infiltration [2], cell distribution, as well as cell growth, proliferation, vascularization and local angiogenesis. Hydrogel interaction with host cells is also largely a result of the hydrogels’ chemical and mechanical compatibility. Mechanically, injectable hydrogels with storage moduli similar to surrounding tissue have been the most successful because they are able to more closely mimic the native environment [102–105]. In one study, the implantation of injectable fibrin hydrogels enhanced neural fiber sprouting and dural resealing following SCI [36]. Another injectable hydrogel, HAMC, was placed into the spinal cord after injury and the results showed that there was reduced inflammation in the animals with HAMC injection compared to controls [44]. Preformed HA hydrogels implanted into the transected spinal cord showed that animals with HA implants had lowered macrophage/microglia cell infiltration and astrocytic response compared to control animals [100]. In some cases, PEG administered following SCI has been shown to help repair damaged membranes and prevent paralysis in pre-clinical models [106–108]. In the clinic, there are a number of FDA-approved dura replacement membrane products made from synthetic (DuraSealTM from PEG hydrogel) and natural-based polymers (DurGenTM from collagen, SynthecelTM from cellulose) with varying success. Synthetic polymers do not provide natural cues and are more difficult for the host’s cells to remodel [109] compared to hydrogels made from native ECM molecules, and therefore may be more suitable as barriers but are less suitable for modifying the cellular response in the post-injury environment.
Hydrogels for cell delivery Source and cell type have been the focus of cell transplantation research in the nervous system, with attempts to implant Schwann cells [110,111], NPCs [112], bone marrow stromal cells [113,114], adult stem cells [115], induced pluripotent stem cells [116] and embryonic stem cells [117], among others (Fig. 2b). Recently, 335
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molecules that enter through the nasal mucosa can have direct access to the brain and spinal cord, and can bypass the blood–brain barrier via the olfactory and trigeminal associated pathways [81,82]. The mucoadhesive property of chitosan increases the dwell time and concentration of the deliverables at the site which results in a higher local concentration and an increase of absorption. A variety of bioactive molecules including antibiotics [70], insulin [65,81], anti-viral [62,63,80] and antidepressants [78,79] have been successfully delivered into the CNS using chitosanbased biomaterials. For example, Thoren et al. successfully used chitosan-based materials to deliver insulin-like growth factor 1 as a possible treatment of Alzheimer’s disease or stroke [81]. Others have used chitosan to deliver vaccines for cell-mediated immune response [62,63,67]. In summary, there is ample evidence that natural-based polymers are well suited for use in nasal, intrathecal and direct drug delivery applications into the CNS. Although naturally derived biomaterials are referenced more often for their innate bioactive properties, biocompatible synthetic polymers are also commonly used. Synthetic biomaterials allow more precise control of bulk properties and degradation rates because synthesis parameters can be easily modified. To this end, aligned and electrospun fibers [83] and freeze-dried macroporous scaffolds [84] have been made from poly-L-lactic acid (PLA) to deliver bioactive molecules into both brain and spinal cord. Moreover, synthetic polymers such as poly(lactic-co-glycolic acid) (PLGA) have been widely used, in experimental settings, to produce microparticles [85–88], and nanoparticles [42,54,85,89,90]. Both micro- and nano-size PLGA particles loaded with bioactive molecules have been used on their own [75,91], but these particles have also been used in combination with hydrogels to localize the particles (as mentioned above with HAMC hydrogels) and to further tune the delivery profiles of the bioactive molecules [54,85,89]. Synthetic polymers can also be used as modifiers for controlled release of small molecules into the CNS. Poly(ethylene glycol) (PEG) is one of the most commonly used synthetic biomaterials in biomedical applications and PEG hydrogels have been developed to deliver growth factors [1,18] and steroids [92] into the CNS in pre-clinical rodent models. In some studies, PEG has been used to modify polyethylenimine (PEI) as part of a gene delivery system. PEGylation, the addition of PEG onto polymers, can increase solubility and decrease protein binding in vivo. For example, researchers compared the use of either PEI alone or PEI conjugated with PEG for gene delivery to the spinal cord. They found that the addition of one linear chain of PEG onto each PEI monomer was sufficient to significantly increase the efficiency of transgene expression in a rodent spinal cord [93]. PEG has also been used to modify the growth factor for obtaining longer bioactivity by increasing the half-life of the growth factor [94,95]. Moreover, a combination of synthetic polymers (i.e., polyethylene glycol– polycaprolactone (PEG–PCL) was used to deliver silencing RNAs via a nasal delivery system in a pre-clinical rodent model [96]. As mentioned previously, bulk properties and degradation rates are more easily modified in synthetic polymers. This can be highly advantageous in developing a platform that can then be modified for tissue or disease specific applications. However, the degradation products of both synthetic and natural polymers can be bioactive and, to the best of our knowledge, currently there is
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research has also focused on mode of cell delivery. Improving widespread problems with low transplanted cell viability after SCI [118], and uncontrolled cell differentiation (transplanted and endogenous cells) in an ischemic stroke model [119] have come to the forefront of today’s research. Biomaterial scaffolds can mitigate a number of these issues by acting as delivery vehicles for cells into injured areas of the nervous system, such as after SCI or traumatic brain injury (TBI) [120,121]. Scaffolds that include natural ECM proteins can direct cell behavior by providing cues to cells during migration [122], differentiation [15], and regeneration after TBI, stroke, and in other injury models [2,123]. These attributes allow hydrogels to provide benefit to the endogenous cells surrounding the implant site as well as to the transplanted cells. Cells are often implanted directly into the lesioned cavity to repair the injured CNS tissue, therefore maintaining cell viability at the site of interest is important [124]. Xiong et al. placed collagen I scaffolds with human marrow stromal cells into a TBI model and observed a significant increase in cell viability compared to cells transplanted in culture medium only [125]. In a separate investigation, thiol crosslinked hyaluronan–heparin–collagen–polyethylene glycol diacrylate hydrogels were used as NPC carriers to the necrotic stroke cavity. Again, significant cell survival of transplanted cells was observed compared to cells transplanted in medium alone (Fig. 3a) [98]. Notably, when the cells were transplanted in combination with gels, there was a significant reduction in microglia/macrophage infiltration. This is an interesting observation because it suggests that the presence of hydrogels either modified the post-injury inflammatory environment (as discussed in the previous section) and/or reduced the inflammation in response to the presence of transplanted cells (Fig. 3b) [98]. In a separate study using a complete transection model of thoracic SCI, Lu et al. demonstrated success at increasing transplanted cell viability as well as functional recovery after SCI with the delivery of NPCs in a fibrin matrix compared to cells in medium (Fig. 3c) [120]. It is important to note that in the Lu study, the researchers included a cocktail of growth factors in the pre-hydrogel solution. Therefore, combinations of supportive scaffolds, growth factors and appropriate cells are needed for effectively repairing the injured tissue. Transplanted cells can also serve as a consistent source of growth factors when they are delivered to the CNS [126]. To achieve this, these cells must be protected from the immune system for longterm viability. Some researchers are using microspheres with cells encapsulated inside or adhered to the outside to protect the cells from the immune system upon delivery in vivo [69,127]. Specifically, alginate capsules can shield brain derived neurotrophic factor (BDNF)-producing fibroblasts from the immune system when implanted in a SCI model [128,129]. This is an intriguing option to deliver cells to the injured CNS without immune suppression; this option could provide a viable treatment without additional medication. In using biomaterials to isolate transplanted cells from the immune system, synthetic materials may provide more immune protection than natural polymers because of the absence of bioactive components. Neural stem and progenitor cells have also been transplanted with the intent that these cells will differentiate into specific cell types that can participate in the recovery process. In this case, utilizing biomaterials with natural ECM molecules, matching 336
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mechanical properties, and introducing tissue relevant factors and topographical cues are ways to direct cell differentiation. Hyaluronic acid-based hydrogels with mechanical properties similar to that of brain and spinal cord tissue have been shown to direct NPC differentiation [16]. Moreover, NPCs can also respond to structural cues and align to follow controlled microstructure created within hyaluronic acid hydrogels [130]. In 2009, Christopherson et al. examined the effects of fiber diameter using electrospun polyethersulfone nanofibers coated with laminin on NPC differentiation [131]. In this study, researchers found that under differentiation conditions, adult NPCs exhibited increased oligodendrocyte differentiation on 283 nm fibers and increased neuronal differentiation on 749 nm fibers compared to cells grown on tissue culture plates [131]. Therefore, topographical cues and mechanical cues within hydrogels are additional tools to direct the fate of NPCs. In a different investigation, Lim et al. observed that aligned electrospun polycaprolactone (PCL) fibers significantly increased NPC differentiation toward neurons and alignment when compared to NPCs grown on tissue culture treated plastic or randomly oriented fibers [132]. Effects of fiber diameter on neurons and glia were examined. The authors found that topographical features such as large fiber diameter and aligned fibers can induce apoptosis of glial cells, increasing the percentage of neuronal cells. This is an exciting finding that introduces additional features for creating more defined regions for specific cell types within a hdyrogel; for example, fiber diameter could control the cell phenotype boundary of white matter and gray matter within the spinal cord [132]. It seems clear that currently available biomaterials can increase cell viability, modify the inflammatory response and encourage differentiation of implanted cells. Next generation hydrogels for CNS repair will likely include native ECM components and similar topography to native tissue.
Hydrogels for tissue regeneration The brain and spinal cord consist of highly specific regions and intricate architecture (e.g., the hippocampus and subventricular zone in the brain, white and gray matter of the spinal cord) [133]. These geometrically complex regions have been explored as design criteria for developing biomaterials for the CNS. Many fundamental studies using cultured cells have shown that creating a complex topographical landscape [130] with submicron and nanoscale features [130] can direct neural cell behavior. As mentioned in the previous section, injectable hydrogels have the ability to conform to the shape of irregular cavities, but creating architectures and altering pore sizes within hydrogel is difficult [101]. In vitro studies of electrospun nanofibers showed embryonic stem cell differentiation into neurons, astrocytes and oligodendrocytes [134]. Moreover, the presence of parallel nanofibers enhanced the neurite extension and orientation [134]. Seidlits et al. reported increased neurite outgrowth of hippocampal cells on 3D microstructures created using 2-photon lithography. The authors created sub-micron structures of bovine serum albumin coated with laminin-derived peptides within HA scaffolds that directed neurite extension and orientation [135]. In another investigation, researchers compared hippocampal NPC affinity to polydimethylsiloxane microsctructure to chemical cues (laminin or nerve growth factor) in vitro. NPCs more preferentially extended neurites onto the microstructures than onto non-structured chemically
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FIGURE 3
Hydrogels can be used to deliver cells to an injury site. This has been shown to increase cell viability, decrease immune response, have an effect on cell differentiation, and increase functional recovery. (a) Zhong et al. transplanted neural progenitor cells (NPCs) into the stroke cavity 7 days after focal ischemia stroke induction with or without hyaluronan-heparin-collagen hydrogels [98]. Two weeks after transplantation, transplanted cells were immunostained with anti-green fluorescent protein (GFP) and quantified by stereological optical fractionator procedures. Cell viability significantly increased when cells were transplanted with a hydrogel compared to cells in saline (p = 0.035). The increase in cell viability is attributed to the delivery with a hydrogel scaffold? (b) Zhong et al. also quantified the immune response after NPC transplantation by immunostaining transplanted NPCs with anti-GFP and active macrophages and microglia with anti-Iba1. Activated macrophages and microglia were quantified around the injury site. There was a significantly decreased quantity of infiltrating activated macrophages and microglia in the injury site with composite hydrogel and cell implantation over cell implantation alone (p = 0.004), quantified by stereological optical fractionator procedures. This shows the positive effect of the hydrogel on the injury environment or on the implanted cells ability to alter the injury environment. (c) Lu et al. transplanted NPCs to fill a thoracic level complete transection site 2 weeks following injury [114]. 7 weeks post transplantation, transplanted NPCs (GFP) were observed to be evenly distributed throughout the injury area (1), to differentiate into mature neurons (2), to robustly extend axons beyond the lesion into the caudal spinal cord in gray matter and white matter (3–5), and rats were seen to have significant functional improvement over the animals with saline alone injections (6, p < 0.01). These results support the idea that the presence of hydrogel scaffold during transplantation can distribute cells throughout the transection area as well as provide support for neuronal differentiation of transplanted NPCs.
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coated substrates [136], suggesting that microstructure plays a critical role in cell behavior. In 2010, Baiguera et al. made electrospun PCL into fibrous membranes with micron-scale and submicron-scale features. Hippocampal astroctyes and endothelial cells cultured on micron-scale features allowed cell infiltration whereas submicron-scale features did not [137]. Therefore, determining the correct feature size for a specific application is an important parameter to consider. These fundamental findings from in vitro work have been incorporated in hydrogel-based biomaterials and the resulting hydrogels have been used in vivo with some success [138,141]. Architecture within hydrogels can also be accomplished by introducing a porous network (Fig. 2c). Micro- and macrostructure in hydrogel scaffolds have been created using a number of methods including porogen leaching [138], freeze drying [97,101], crystal templating [139], multiphoton lithography [135] and molding [140]. Porogen leaching and gas foaming create pore sizes relatively consistent with the size of the porogen or gas used to create the pores [102,138]. It is not entirely clear if having a consistent pore size is a critical requirement for successful biomaterial integration but optimum pore sizes have been determined for a number of applications [102]. Martinez-Ramos et al. used porogen leaching to synthesize poly(ethyl acrylate)–poly (hydroxyl ethyl acrylate) copolymer blend matrices with two different geometries for implantation near the subventricular zone (SVZ) of the brain [138]. Pre-clinical in vivo implants of hydrogels with aligned 40 mm diameter pores and with interconnected orthogonal 80 mm pores showed local angiogenesis and neurite extension within each of the scaffolds in the SVZ [138]. In in vitro experiments, NPCs readily infiltrated the scaffold and differentiated into neurons and astrocytes regardless of the scaffold geometry [138]. It appears from this study that specific scaffold geometry does not play a significant role in scaffold integration or cell penetration in the brain; it is likely that the overall porosity plays the largest role. In another study, PCL scaffolds were created with two kinds of internal architectures using a molding method: (1) longitudinal channels with microgrooves within a cylinder, or (2) orthogonally intersecting channels and axial microgrooves within a cylinder [141]. When implanted into adult rat cortex, the group with orthogonal channel implants that more closely mimic the native ECM showed the most tissue in-growth compared to the longitudinal channel implant or control scaffold implant groups. It is important to note that in this study the pore sizes were hundreds of microns. Therefore, when the pores are larger than 100 mm the internal architecture may play a more significant role in tissue ingrowth from the host. In 2005, Tian et al. implanted freeze-dried hyaluronic acid-poly-D-lysine composite gels with mean pore diameters ranging from 90 to 230 mm in a TBI model in adult rats. These porous scaffolds with internal architecture (fiber-like struts and sheet-like features) showed successful integration with host tissue, endothelialization and new ECM formation by six weeks [97]. Macro-architecture that mimics the natural tissue is thought to encourage regeneration. Repairing the spinal cord presents an interesting challenge to biomaterials scientists and engineers in that the white and gray matter architectures are dramatically different. Depending on the injury type and location, different architectures may be more relevant. The white matter contains 338
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well-aligned axon tracts, whereas the gray matter houses local motor neurons and interneuron pools and would likely require less directed growth. Friedman et al. theorized that creating architecture with specific features to mimic spinal cord tracts would encourage regeneration [142], however, the suggested designs have not been fully developed. Hydrogel scaffolds with macroarchitecture have been explored recently. Stokols et al. designed templated agarose scaffolds for use after dorsal column lesions. The extrusion method created aligned linear cylindrical structure for white matter regeneration. The authors observed host integration including Schwann cell infiltration and vascularization in implants. Further, the researchers noted a minimal increase in inflammation near the tissue-injury interface compared to untreated animals. Axon extension was observed only when Matrigel1 or fibrin matrices were included within the implants [143], suggesting that natural ECM cues are required for axonal growth within the graft. Hydrogel scaffolds with slightly more complex architecture have been explored recently to repair the spinal cord. Wong et al. cast five PCL scaffolds with different architectures and implanted them into a complete transection thoracic spinal cord. The authors compared three designs with non-specific architecture and two designs with specific architecture that more closely mimics white and gray matter of the spinal cord. At three months, non-specific architecture implants had higher inflammatory response (increased macrophages, fibroblasts, astrocytes) as well as lesion expansion compared to the structure-mimicking implants. Scaffolds with the structure-mimicking designs also encouraged more axonal infiltration, astrocyte migration and myelinated axon growth into the material compared to the other designs [140]. In this study, meaningful functional recovery was not assessed, but it provided a strong argument in support of anatomically relevant macro-architectural cues for enhancing regeneration. In summary, when it comes to incorporating structure into hydrogel scaffolds, researchers must determine the most relevant biological features, and size scale and determine how to create similar architectures within biomaterials. Although, there are new scaffolds being tested in pre-clinical models in the laboratory, currently there are none available in the clinic. Future work in this area needs to focus on fine-tuning the optimal morphological layout, length scales, and feature sizes. Molding and multiphoton lithography show the most promise as methods to create robust, relevant geometry within hydrogels because they provide precise control over macro- and microarchitecture. It is clear from fundamental studies that topography can play a significant role in directing cell behavior. Therefore, biomaterials that better mimic the native architecture will continue to be an important area of study for CNS regeneration.
Concluding remarks Looking ahead, advances being made in the field of biomaterial science will significantly impact the next generation of therapies for repair in the CNS. Hydrogels in particular are strong candidates for clinical relevance in the future. Indeed, as mentioned in this short review, there are a number of hydrogel genres that have been used successfully in the CNS in the laboratory. However, before any biomaterial can be used in the clinic, we must consider thoroughly how materials interact with surrounding cells and
host tissue in vivo. Key considerations and material features that we feel are important include: (1) a better understanding of the degradation products of parent polymers used for drug delivery applications, (2) equal consideration of the possible routes of delivery including nasal, intrathecal and direct implantation, (3) careful examination of possible biological effects of the hydrogels and their degrading products, (4) inclusion of native ECM components within the scaffolds during cell delivery for better cell survival and integration with the host tissue and (5) addition of biologically relevant topography and architectural features within scaffolds. Significant advances have been made to improve both functional recovery and cell viability following SCI. Additionally, progenitor cell differentiation and the reduction in inflammation as a result of hydrogel implantation are notable achievements in this area of research. Future efforts should focus on a more comprehensive approach to developing and studying biomaterials which include relevant in vivo models.
Acknowledgments The authors wish to acknowledge related funding from the Mission Connect (#012-113 Z.Z.K. and C.E.S.), The Craig Neilsen Foundation (#222456, C.E.S.), NSF-CBET (#1159774 C.E.S.), NSFDMR (#0805298 C.E.S.), NIH R21 (#R21 NS074162 C.E.S.), NSFGFRP (#2011112479, S.A.G.), and the National GEM Consortium (R.C.T.). References [1] [2] [3] [4] [5] [6] [7] [8] [9] [10] [11] [12] [13] [14] [15] [16] [17] [18] [19] [20] [21] [22] [23] [24] [25] [26] [27] [28] [29] [30] [31] [32] [33] [34] [35] [36] [37] [38] [39]
D.R. Nisbet, et al. J. Biomater. Appl. 24 (1) (2009) 7. J.T.S. Pettikiriarachchi, et al. Aust. J. Chem. 63 (8) (2010) 1143. D. Macaya, M. Spector, Biomed. Mater. 7 (1) (2012) 012001. D.A. McCreedy, S.E. Sakiyama-Elbert, Neurosci. Lett. 519 (2) (2012) 115. K. Straley, S.C. Heilshorn, Conf. Proc. IEEE Eng. Med. Biol. Soc. 2009 (2009) 2101. Y. Zhong, R.V. Bellamkonda, J. R. Soc. Interface 5 (26) (2008) 957. Z.Z. Khaing, C.E. Schmidt, Neurosci. Lett. 519 (2) (2012) 103. H. Kim, et al. Trends Biotechnol. 30 (1) (2012) 55. G. Orive, et al. Nat. Rev. Neurosci. 10 (9) (2009) 682. H. Cao, et al. Adv. Drug Deliv. Rev. 61 (12) (2009) 1055. R. Uibo, et al. Biochim. Biophys. Acta 1793 (5) (2009) 924. O.K. Bitzer-Quintero, I. Gonzalez-Burgos, Neural Plast. 2012 (2012) 348642. R.J. McKeon, et al. J. Neurosci. 11 (11) (1991) 3398. P.C. Georges, et al. Biophys. J. 90 (8) (2006) 3012. A.J. Engler, et al. Cell 126 (4) (2006) 677. S.K. Seidlits, et al. Biomaterials 31 (14) (2010) 3930. H. Liao, et al. Acta Biomater. 4 (5) (2008) 1161. J.A. Burdick, et al. Biomaterials 27 (3) (2006) 452. H.Y. Yang, et al. Mol. Chem. Neuropathol. 21 (2–3) (1994) 155. M.C. Shearer, J.W. Fawcett, Cell Tissue Res. 305 (2) (2001) 267. R.J. Pasterkamp, et al. Mol. Cell. Neurosci. 13 (2) (1999) 143. J. Silver, J.H. Miller, Nat. Rev. Neurosci. 5 (2) (2004) 146. C.M. Liberto, et al. J. Neurochem. 89 (5) (2004) 1092. P.A. Harvey, et al. J. Neurosci. 29 (19) (2009) 6285. W.B. Cafferty, S.M. Strittmatter, J. Neurosci. 26 (47) (2006) 12242. X. Wang, et al. J. Neurosci. 22 (13) (2002) 5505. T. GrandPre, et al. Nature 417 (6888) (2002) 547. J.R. Faulkner, et al. J. Neurosci. 24 (9) (2004) 2143. B. Gardner, et al. Paraplegia 33 (10) (1995) 551. T.M. Miller, et al. Lancet Neurol. 12 (5) (2013) 435. G. Lind, et al. Acta Neurochir. Suppl. 97 (Pt 1) (2007) 57. S. Erdine, J. De Andres, Pain Pract. 6 (1) (2006) 51. B.R. Felice, et al. Toxicol. Pathol. 39 (5) (2011) 879. C.E. Kang, et al. J. Control. Release 144 (1) (2010) 25. A.S. Lossinsky, R.R. Shivers, Histol. Histopathol. 19 (2) (2004) 535. P.J. Johnson, et al. Biotechnol. Bioeng. 104 (6) (2009) 1207. P.J. Johnson, et al. Soft Matter 6 (20) (2010) 5127. S.E. Sakiyama-Elbert, J.A. Hubbell, J. Control. Release 69 (1) (2000) 149. S.J. Taylor, et al. J. Control. Release 113 (3) (2006) 226.
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[40] [41] [42] [43] [44] [45] [46] [47] [48] [49] [50] [51] [52] [53] [54] [55] [56] [57] [58] [59] [60] [61] [62] [63] [64] [65] [66] [67] [68] [69] [70] [71] [72] [73] [74] [75] [76] [77] [78] [79] [80] [81] [82] [83] [84] [85] [86] [87] [88] [89] [90] [91] [92] [93] [94] [95] [96] [97] [98] [99] [100] [101] [102] [103] [104] [105] [106] [107] [108] [109] [110] [111]
J.W. Austin, et al. Biomaterials 33 (18) (2012) 4555. M.D. Baumann, et al. J. Control. Release 138 (3) (2009) 205. M.D. Baumann, et al. Biomaterials 31 (30) (2010) 7631. M.J. Caicco, et al. J. Biomed. Mater. Res. A 101 (5) (2013) 1472. D. Gupta, et al. Biomaterials 27 (11) (2006) 2370. M.J. Cooke, et al. Biomaterials 32 (24) (2011) 5688. W.M. Tian, et al. J. Control. Release 102 (1) (2005) 13. R.K. Mittapalli, et al. Mol. Cancer Ther. 12 (11) (2013) 2389. Y. Wang, et al. Pharm. Res. 28 (6) (2011) 1406. A. Jain, et al. PLoS ONE 6 (1) (2011) e16135. Z. Yang, et al. Biomaterials 31 (18) (2010) 4846. M.S. Shoichet, et al. Prog. Brain Res. 161 (2007) 385. C.E. Kang, et al. Tissue Eng. Part A 15 (3) (2009) 595. A.J. Mothe, et al. Biomaterials 34 (15) (2013) 3775. C.E. Kang, et al. Cells Tissues Organs 197 (1) (2013) 55. M.D. Wood, et al. J. Biomed. Mater. Res. A 89 (4) (2009) 909. S.J. Taylor, S.E. Sakiyama-Elbert, J. Control. Release 116 (2) (2006) 204. A.J. Hyatt, et al. J. Control. Release 147 (1) (2010) 24. P.J. Johnson, et al. Cell Transplant. 19 (1) (2010) 89. S.M. Willerth, et al. Stem Cells 25 (9) (2007) 2235. S.M. Willerth, et al. Stem Cell Res. 1 (3) (2008) 205. A. Jain, et al. Biomaterials 27 (3) (2006) 497. Y. Wu, et al. Eur. J. Pharm. Biopharm. 81 (3) (2012) 486. Y. Wu, et al. Biomaterials 33 (7) (2012) 2351. H. Nazar, et al. Eur. J. Pharm. Biopharm. 77 (2) (2011) 225. A.K. Agrawal, et al. Pharmazie 65 (3) (2010) 188. I.A. Alsarra, et al. Drug Dev. Ind. Pharm. 35 (3) (2009) 352. J. Wu, et al. Biomaterials 28 (13) (2007) 2220. E. Gavini, et al. J. Pharm. Sci. 100 (4) (2011) 1488. N.B. Skop, et al. Acta Biomater. 9 (6) (2013) 6834. E. Gavini, et al. J. Pharm. Sci. 100 (2011) 1488–1502. J. Malmo, et al. PLoS ONE 8 (1) (2013) e54182. M. Yemisci, et al. Methods Enzymol. 508 (2012) 253. A. Mahapatro, D.K. Singh, J. Nanobiotechnol. 9 (2011) 55. A. Hafner, et al. J. Microencapsul. 28 (8) (2011) 807. Y. Cho, R.B. Borgens, Exp. Neurol. 233 (1) (2012) 126. A. Mistry, et al. J. Drug Target. 17 (7) (2009) 543. K. Nagpal, et al. Drug Deliv. 20 (3–4) (2013) 112. K. Nagpal, et al. Int. J. Biol. Macromol. 59 (2013) 20. K. Nagpal, et al. Drug Deliv. 19 (8) (2012) 378. K. Nagpal, et al. Chem. Pharm. Bull. (Tokyo) 58 (11) (2010) 1423. R.G. Thorne, et al. Neuroscience 127 (2) (2004) 481. U. Westin, et al. Eur. J. Pharm. Sci. 24 (5) (2005) 565. N.J. Schaub, R.J. Gilbert, J. Neural Eng. 8 (4) (2011) 046026. C.M. Patist, et al. Biomaterials 25 (9) (2004) 1569. S.A. Chvatal, et al. Biomaterials 29 (12) (2008) 1967. A. Goraltchouk, et al. J. Control. Release 110 (2) (2006) 400. G.E. Rooney, et al. Tissue Eng. Part A 17 (9–10) (2011) 1287. E.Y. Tan, et al. Pharm. Res. 24 (12) (2007) 2297. J.C. Stanwick, et al. J. Control. Release 160 (3) (2012) 666. Y.C. Wang, et al. Biomaterials 29 (34) (2008) 4546. Y.T. Kim, et al. Biomaterials 30 (13) (2009) 2582. P.J. Gaillard, et al. J. Control. Release 164 (3) (2012) 364. G.P. Tang, et al. Biomaterials 24 (13) (2003) 2351. D.P. Ankeny, et al. Exp. Neurol. 170 (1) (2001) 85. R.G. Soderquist, et al. J. Biomed. Mater. Res. A 91 (3) (2009) 719. T. Kanazawa, et al. Biomaterials 34 (36) (2013) 9220. W.M. Tian, et al. Tissue Eng. 11 (3–4) (2005) 513. J. Zhong, et al. Neurorehabil. Neural Repair 24 (7) (2010) 636. J.E. Rinholm, et al. J. Neurosci. 31 (2) (2011) 538. Z.Z. Khaing, et al. J. Neural Eng. 8 (4) (2011) 046033. P.Z. Elias, M. Spector, J. Tissue Eng. Regen. Med. (2012), http://dx.doi.org/ 10.1002/term.1621. N. Annabi, et al. Tissue Eng. Part B: Rev. 16 (4) (2010) 371. K. Whang, et al. Tissue Eng. 5 (1) (1999) 35. S. Woerly, Biomaterials 14 (14) (1993) 1056. R.J. Wade, J.A. Burdick, Mater. Today 15 (10) (2012) 454. S. Kouhzaei, et al. Neurol. Res. 35 (4) (2013) 415. A. Nehrt, et al. J. Neurotrauma 27 (1) (2010) 151. R.B. Borgens, et al. J. Exp. Biol. 205 (Pt 1) (2002) 1. J. Patterson, et al. Mater. Today 13 (1–2) (2010) 14. T. Hadlock, et al. Tissue Eng. 6 (2) (2000) 119. H.E. Olson, et al. Tissue Eng. Part A 15 (7) (2009) 1797.
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[112] [113] [114] [115] [116] [117] [118] [119] [120] [121] [122] [123] [124] [125] [126] [127]
340
J.D. Kocsis, et al. J. Neurotrauma 21 (4) (2004) 441. H. Itosaka, et al. Neuropathology 29 (3) (2009) 248. D. Lu, et al. Neurosurgery 61 (3) (2007) 596. M. Tohill, et al. Neurosci. Lett. 362 (3) (2004) 200. A. Wang, et al. Biomaterials 32 (22) (2011) 5023. L. Cui, et al. Stem Cells 26 (5) (2008) 1356. A.M. Parr, et al. Neuroscience 155 (3) (2008) 760. S. Kelly, et al. Proc. Natl. Acad. Sci. U.S.A. 101 (32) (2004) 11839. P. Lu, et al. Cell 150 (6) (2012) 1264. Y.J. Ren, et al. J. Bioact. Compat. Polym. 24 (1) (2009) 56. E. Tziampazis, et al. Biomaterials 21 (5) (2000) 511. W. Daly, et al. J. R. Soc. Interface 9 (67) (2012) 202. M.J. Cooke, et al. Soft Matter 6 (20) (2010) 4988. Y. Xiong, et al. Brain Res. (2009) 1263–2183. B. Rihova, Adv. Drug Deliv. Rev. 42 (1–2) (2000) 65. R.M. Hernandez, et al. Adv. Drug Deliv. Rev. 62 (7–8) (2010) 711.
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[128] [129] [130] [131] [132] [133] [134] [135] [136] [137] [138] [139] [140] [141] [142] [143]
C.A. Tobias, et al. J. Neurotrauma 22 (1) (2005) 138. C.A. Tobias, et al. J. Neurotrauma 18 (3) (2001) 287. S.K. Seidlits, et al. Nanomedicine (Lond.) 3 (2) (2008) 183. G.T. Christopherson, et al. Biomaterials 30 (4) (2009) 556. S.H. Lim, et al. Biomaterials 31 (34) (2010) 9031. S.C. Joshi, et al. Int. J. Pattern Recognit. Artif. Intell. 11 (08) (1997) 1317. J. Xie, et al. Nanoscale 2 (1) (2010) 35. S.K. Seidlits, et al. Adv. Funct. Mater. 20 (9) (2010) 1368. N. Gomez, et al. J. R. Soc. Interface 4 (13) (2007) 223. S. Baiguera, et al. J. Mater. Sci. Mater. Med. 21 (4) (2010) 1353. C. Martinez-Ramos, et al. J. Biomed. Mater. Res. A 100 (12) (2012) 3276. S.A. Zawko, C.E. Schmidt, Acta Biomater. 6 (7) (2010) 2415. D.Y. Wong, et al. J. Neurotrauma 25 (8) (2008) 1027. D.Y. Wong, et al. J. Neurosurg. 109 (4) (2008) 715. J.A. Friedman, et al. Neurosurgery 51 (3) (2002) 742. S. Stokols, et al. Tissue Eng. 12 (10) (2006) 2777.
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Advanced biomaterials for repairing the nervous system: what can hydrogels do for the brain? Zin Z. Khaing1,*, Richelle C. Thomas2, Sydney A. Geissler3 and Christine E. Schmidt1,* 1
J. Crayton Pruitt Family Department of Biomedical Engineering, The University of Florida, United States McKetta Department of Chemical Engineering, The University of Texas at Austin, United States 3 Department of Biomedical Engineering, The University of Texas at Austin, United States 2
Newly developed hydrogels are likely to play significant roles in future therapeutic strategies for the nervous system. In this review, unique features of the central nervous system (i.e., the brain and spinal cord) that are important to consider in developing engineered biomaterials for therapeutic applications are discussed. This review focuses on recent findings in hydrogels as biomaterials for use as (1) drug delivery devices, specifically focusing on how the material can change the delivery rate of small molecules, (2) scaffolds that can modify the post-injury environment, including preformed and injectable scaffolds, (3) cell delivery vehicles, discussing cellular response to natural and synthetic polymers as well as structured and amorphous materials, and (4) scaffolds for tissue regeneration, describing micro- and macro-architectural constructs that have been designed for neural applications. In addition, key features in each category that are likely to contribute to the translational success of these biomaterials are highlighted. Introduction Recent advances both in our understanding of the nervous system and the availability of sophisticated biomaterials have significantly changed the landscape of potential strategies for repairing the nervous system. New imaging and staining techniques along with the development of novel materials have opened the possibility to directly design biomaterials tailored for a particular application. In this short review, recent advances in the use of hydrogels made from both natural and synthetic polymers are discussed (see Fig. 1 as reference) and their evaluations using in vitro and in vivo preclinical models are presented where applicable. This review is by no means a comprehensive review of biomaterials for nervous system repair: the readers are referred to a number of other excellent review articles for further studies [1–11]. This review introduces important obstacles to central nervous system (CNS) regeneration focusing on the unique characteristics of the CNS. Additionally, the injury environment and scarring are unique to the CNS in that a glial scar introduces a physical and chemical *Corresponding authors:. Khaing, Z.Z. ([email protected]), Schmidt, C.E. ([email protected])
barrier to regeneration. Understanding the specific requirements and taking them into consideration in designing therapeutic platforms are likely to result in the successful development of next generation biomaterials. Briefly, the blood barriers, the endogenous immune system within the CNS, and the mechanical properties of CNS tissues are discussed. The focus of this review is on the use of hydrogels as scaffolds to aid regeneration within the central nervous system (CNS) (i.e., the brain and spinal cord). In particular, the following systems and applications are highlighted: (1) hydrogels as drug delivery platforms, (2) hydrogels to modify the post-injury environment, (3) hydrogels for cell delivery and (4) hydrogels for tissue regeneration.
Features of the central nervous system to consider for biomaterial development There are a number of challenges that are unique to the CNS that must be considered for development of therapeutic biomaterials. First, the CNS is segregated from the circulating blood by the blood brain barrier (BBB) and the blood spinal cord barrier (BSCB). These barriers are made up of tight junctions between extracellular membranes of endothelial cells and astrocytes. In normal physiological 1369-7021/ß 2014 Elsevier Ltd. All rights reserved. http://dx.doi.org/10.1016/j.mattod.2014.05.011
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FIGURE 1
Synthetic and natural materials can be used to deliver small molecules or cells to the nervous system. Micro and macro architecture can also be introduced to these hydrogels to alter injury environments and direct cell behavior. Chemical structures of some common polymers are presented here. Abbreviations: PLGA – poly(lactic-co-glycolic acid), PEG – poly(ethylene glycol), PCL – polycaprolactone.
conditions, few molecules can passively move from the circulating blood to the CNS extracellular fluid. This makes it difficult, if not impossible, to deliver drugs and small molecule therapeutics to the CNS intravenously. It has been suggested that hydrogels made from both natural and synthetic polymers that can be intrathecally placed into localized areas show great promise to deliver therapeutics into the brain and spinal cord. Under normal conditions, the BBB and BSCB also isolate the CNS tissue from the circulating immune cells. Thus, the CNS is considered an ‘immune privileged’ organ. Major immune cells within the CNS include (1) microglia, which are resident immune cells that are distributed throughout the brain and the spinal cord, and (2) perivascular macrophages that are located in the capillaries. After injury or in response to disease, microglia, astroctyes, macrophages, oligodendroctyes and even neurons to an extent, can all respond and release inflammatory cytokines [12]. Therefore, for biomaterials to modify the injury environment, the materials must interact with the resident cells’ immune response in a positive manner [13]. The brain and spinal cord are some of the softest tissues in the body with compressive moduli around 2000 Pa [14–16]. Matching the mechanical properties of an implanted biomaterial to that of the host tissue can significantly affect the success of the implanted material in vivo [2,17]. In vitro studies of neural progenitor cell (NPC) differentiation showed that hydrogels that best matched the stiffness of the brain provided the most optimal results for neuronal differentiation [16,18]. Hence, determining the appropriate mechanical property is a key factor to consider when using biomaterials in the CNS.
Injury environment of the nervous system A major difference between the peripheral nervous system (PNS) and CNS is the capacity for peripheral nerves to regenerate; CNS axons do not regenerate appreciably in their native environment.
After injury in the CNS, macrophages infiltrate the site of injury much more slowly than they do in the PNS. This delays the removal of inhibitory myelin associated proteins from the injury site. Macrophage recruitment is also limited because cell adhesion molecules in the distal end of the injured spinal cord are not appreciably up-regulated. Additionally, astrocytes in the CNS become ‘reactive’, and produce glial scar tissue. Astrocytic response after injury is characterized by cellular hypertrophy, astrocyte proliferation, process extension, and increased production of the intermediate filament proteins glial fibrillary acidic protein (GFAP), vimentin, and nestin [19]. This response of astrocytes to injury is known as reactive gliosis. When an injury does not involve penetrating the dura mater, the scar tissue formed is mainly composed of astrocytes; however, when the dura is broken there is infiltration of meningeal fibroblasts in addition to reactive astrocytes [20]. Most glial scar tissue is thought to include, in addition to the reactive astrocytes, NG2+ oligodendrocyte precursor cells, meningeal cells, infiltrating macrophages, and activated microglia. Schwann cells from adjacent dorsal roots have also been found within the CNS scar tissue in experimental injuries where the dura was broken. The mature glial scar includes proteoglycans such as neurocan, phosphocan, versican, and brevican along with secreted proteins including Semaphorin 3A and 3D [21]. One notable component of glial scar tissue is chondroitin sulfate proteoglycans (CSPGs). CSPGs have been shown to inhibit axonal outgrowth in many neuronal systems [22]. In sites distant from traumatic injury, astrocytes can also become larger in size and transform into a more pronounced stellate shape. These ‘activated astrocytes’ have been known to produce soluble trophic factors that enhance the survival of neurons and glial cells in the vicinity of the astrocytes [23]. Therefore, the effects of activated astrocytes after injury on axonal regeneration and neuronal plasticity is unclear. Moreover, myelin-related glycoproteins have long been implicated in creating a non-hospitable environment for the 333
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severed axons of the CNS [24–27]. Combined, these factors constitute a formidable environment for regenerating axons in the CNS. Although the idea of simply inhibiting the formation of scar after injury may be an attractive option, it is important to note that formation of scar tissue after injury to the CNS encourages tissue homeostasis. Contusion spinal cord injury (SCI) studies completed in transgenic mice that targeted the depletion of mitotic astrocytes immediately surrounding the injury site support this idea. The results show an increased infiltration of inflammatory cells, cellular degeneration and increased lesioned area compared to control mice with a similar injury [28]. Therefore it appears that there are specific beneficial aspects of the glial cell response. Specifically the production of cytokines, growth factors, and other extracellular matrix (ECM) components that participate in re-establishing the BBB after injury have been shown advantageous to the post-injury environment. Simply inhibiting the formation of scar is not a viable option. Instead, strategies that will specifically antagonize the receptors for the inhibitors may be a better way to interfere with the growth inhibitory nature of the injured CNS microenvironment.
Hydrogels as drug delivery platforms As stated above, delivering molecules to the CNS is a challenging problem due to the presence of BBB and BSCB. Direct intrathecal injections into the extracellular fluid have been successful in some clinical [29–32] and experimental/pre-clinical settings [33,34]. Current research in this field involves examining biocompatible and bio-inert polymers as platforms for controlled release of bioactive molecules within the CNS for sustained drug delivery (Fig. 2a). It is important to note that, after injury to the brain and spinal cord, there is a temporary leak in the BBB and BSCB that allows more molecules and cells to enter the CNS from the circulating blood compared to normal physiological conditions [35]. However, the extent to which this ‘leakiness’ can be taken advantage of for cell and drug delivery is unclear. Hydrogels with matching mechanical properties to that of the nervous tissue (compressive modulus 2000 Pa) can be placed in the intrathecal space to maintain homogeneous mechanical landscape with the surrounding soft CNS tissue [16]. In light of this feature, hydrogels made from a number of natural and synthetic polymers have been examined for their ability to deliver therapeutics directly into the brain and spinal cord. Naturally derived polymers including fibrin [36–39], hyaluronan–methylcellulose (HAMC) blend [40–45], hyaluronic acid [46– 48], agarose [49] and chitosan [50] have been used to successfully deliver molecules to the CNS. For example, an injectable hydrogel system developed by the Shoichet group shows promise for intrathecal delivery of growth factors into the injured spinal cord. This group developed HAMC hydrogels that are safe, effective [51], and able to deliver trophic factors [52]; modified versions of these HAMC hydrogels can also be used as vehicles to deliver cells into the injured spinal cord [53]. In one case, the delivery of erythropoietin (EPO) into the intrathecal space using HAMC hydrogels after SCI in a rodent model resulted in reduction in lesion cavity size and had a higher neuroprotective effect compared to direct injection of EPO alone into the intrathecal space [52]. In other cases, HAMC has been successfully combined with nanoparticles 334
FIGURE 2
A schematic representation of the three fundamental ways in which hydrogels are being developed for therapies to repair the nervous system is shown here. (a) Hydrogels alone or as composite materials with microparticles have been used to create highly tunable delivery platforms for small molecule therapeutics into the CNS. (b) Hydrogels can also be used to deliver cells into the CNS. (c) Hydrogels have been utilized, with and without internal architecture, as scaffolds to deliver cells and repair injured CNS tissue.
resulting in composite biomaterials with highly tunable delivery profiles [42,54]. Fibrin-based hydrogels have also been successfully used for delivery of growth factors [39,36,55,56], scar inhibiting enzyme chondroitinase ABC [57], and progenitor/stem cells in combination with growth factors [58–60] in rodent models of SCI. Likewise, injectable, in situ-gelling forms of agarose have been used in combination with lipid microtubules loaded with bioactive molecules after SCI [49,61]. Bioactive molecules have been delivered to the CNS with chitosan-based hydrogels [62–67], microparticles [68–70] or nanoparticles [71–80]. Chitosan-based biomaterials in particular have largely been explored for delivery through the nasal route. This delivery paradigm is attractive since the nasal passage has a large surface area, porous endothelial membrane, high total blood flow and is readily accessible. Indeed,
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limited information on possible biological effects of degradation products from synthetic polymers [97,98]. For example, lactic acid and glycolic acid are known byproducts of PLGA. Recently, an intriguing article by Rinholm et al. suggests that exogenous application of lactic acid to cultured slices of developing mouse brain cortex can support oligodendrocyte development and myelination [99]. Future studies examining the presence and possible effects of degradation products from commonly used polymers, such as PLGA, are warranted when applied in vivo.
Hydrogels to modify the post-injury environment Hydrogels can be particularly useful in the CNS post injury environment by potentially serving as a substrate to re-establish tissue continuity. Prefabricated scaffolds have been implanted [100], but can cause secondary injury to the surrounding tissue during delivery [101]. Minimally invasive injectable hydrogels are particularly well suited for these types of applications. The liquid polymer can fill irregularly shaped void spaces and injections via needles to deliver the liquid polymer circumvent the BBB and BSCB. The physical properties of hydrogels that are relevant to consider in this post-injury environment are the porosity, chemical composition and mechanical properties. Hydrogel porosity is essential to stabilize the post-injury environment by permitting nutrient flow into and out of the scaffold. Porosity also affects cell infiltration [2], cell distribution, as well as cell growth, proliferation, vascularization and local angiogenesis. Hydrogel interaction with host cells is also largely a result of the hydrogels’ chemical and mechanical compatibility. Mechanically, injectable hydrogels with storage moduli similar to surrounding tissue have been the most successful because they are able to more closely mimic the native environment [102–105]. In one study, the implantation of injectable fibrin hydrogels enhanced neural fiber sprouting and dural resealing following SCI [36]. Another injectable hydrogel, HAMC, was placed into the spinal cord after injury and the results showed that there was reduced inflammation in the animals with HAMC injection compared to controls [44]. Preformed HA hydrogels implanted into the transected spinal cord showed that animals with HA implants had lowered macrophage/microglia cell infiltration and astrocytic response compared to control animals [100]. In some cases, PEG administered following SCI has been shown to help repair damaged membranes and prevent paralysis in pre-clinical models [106–108]. In the clinic, there are a number of FDA-approved dura replacement membrane products made from synthetic (DuraSealTM from PEG hydrogel) and natural-based polymers (DurGenTM from collagen, SynthecelTM from cellulose) with varying success. Synthetic polymers do not provide natural cues and are more difficult for the host’s cells to remodel [109] compared to hydrogels made from native ECM molecules, and therefore may be more suitable as barriers but are less suitable for modifying the cellular response in the post-injury environment.
Hydrogels for cell delivery Source and cell type have been the focus of cell transplantation research in the nervous system, with attempts to implant Schwann cells [110,111], NPCs [112], bone marrow stromal cells [113,114], adult stem cells [115], induced pluripotent stem cells [116] and embryonic stem cells [117], among others (Fig. 2b). Recently, 335
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molecules that enter through the nasal mucosa can have direct access to the brain and spinal cord, and can bypass the blood–brain barrier via the olfactory and trigeminal associated pathways [81,82]. The mucoadhesive property of chitosan increases the dwell time and concentration of the deliverables at the site which results in a higher local concentration and an increase of absorption. A variety of bioactive molecules including antibiotics [70], insulin [65,81], anti-viral [62,63,80] and antidepressants [78,79] have been successfully delivered into the CNS using chitosanbased biomaterials. For example, Thoren et al. successfully used chitosan-based materials to deliver insulin-like growth factor 1 as a possible treatment of Alzheimer’s disease or stroke [81]. Others have used chitosan to deliver vaccines for cell-mediated immune response [62,63,67]. In summary, there is ample evidence that natural-based polymers are well suited for use in nasal, intrathecal and direct drug delivery applications into the CNS. Although naturally derived biomaterials are referenced more often for their innate bioactive properties, biocompatible synthetic polymers are also commonly used. Synthetic biomaterials allow more precise control of bulk properties and degradation rates because synthesis parameters can be easily modified. To this end, aligned and electrospun fibers [83] and freeze-dried macroporous scaffolds [84] have been made from poly-L-lactic acid (PLA) to deliver bioactive molecules into both brain and spinal cord. Moreover, synthetic polymers such as poly(lactic-co-glycolic acid) (PLGA) have been widely used, in experimental settings, to produce microparticles [85–88], and nanoparticles [42,54,85,89,90]. Both micro- and nano-size PLGA particles loaded with bioactive molecules have been used on their own [75,91], but these particles have also been used in combination with hydrogels to localize the particles (as mentioned above with HAMC hydrogels) and to further tune the delivery profiles of the bioactive molecules [54,85,89]. Synthetic polymers can also be used as modifiers for controlled release of small molecules into the CNS. Poly(ethylene glycol) (PEG) is one of the most commonly used synthetic biomaterials in biomedical applications and PEG hydrogels have been developed to deliver growth factors [1,18] and steroids [92] into the CNS in pre-clinical rodent models. In some studies, PEG has been used to modify polyethylenimine (PEI) as part of a gene delivery system. PEGylation, the addition of PEG onto polymers, can increase solubility and decrease protein binding in vivo. For example, researchers compared the use of either PEI alone or PEI conjugated with PEG for gene delivery to the spinal cord. They found that the addition of one linear chain of PEG onto each PEI monomer was sufficient to significantly increase the efficiency of transgene expression in a rodent spinal cord [93]. PEG has also been used to modify the growth factor for obtaining longer bioactivity by increasing the half-life of the growth factor [94,95]. Moreover, a combination of synthetic polymers (i.e., polyethylene glycol– polycaprolactone (PEG–PCL) was used to deliver silencing RNAs via a nasal delivery system in a pre-clinical rodent model [96]. As mentioned previously, bulk properties and degradation rates are more easily modified in synthetic polymers. This can be highly advantageous in developing a platform that can then be modified for tissue or disease specific applications. However, the degradation products of both synthetic and natural polymers can be bioactive and, to the best of our knowledge, currently there is
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research has also focused on mode of cell delivery. Improving widespread problems with low transplanted cell viability after SCI [118], and uncontrolled cell differentiation (transplanted and endogenous cells) in an ischemic stroke model [119] have come to the forefront of today’s research. Biomaterial scaffolds can mitigate a number of these issues by acting as delivery vehicles for cells into injured areas of the nervous system, such as after SCI or traumatic brain injury (TBI) [120,121]. Scaffolds that include natural ECM proteins can direct cell behavior by providing cues to cells during migration [122], differentiation [15], and regeneration after TBI, stroke, and in other injury models [2,123]. These attributes allow hydrogels to provide benefit to the endogenous cells surrounding the implant site as well as to the transplanted cells. Cells are often implanted directly into the lesioned cavity to repair the injured CNS tissue, therefore maintaining cell viability at the site of interest is important [124]. Xiong et al. placed collagen I scaffolds with human marrow stromal cells into a TBI model and observed a significant increase in cell viability compared to cells transplanted in culture medium only [125]. In a separate investigation, thiol crosslinked hyaluronan–heparin–collagen–polyethylene glycol diacrylate hydrogels were used as NPC carriers to the necrotic stroke cavity. Again, significant cell survival of transplanted cells was observed compared to cells transplanted in medium alone (Fig. 3a) [98]. Notably, when the cells were transplanted in combination with gels, there was a significant reduction in microglia/macrophage infiltration. This is an interesting observation because it suggests that the presence of hydrogels either modified the post-injury inflammatory environment (as discussed in the previous section) and/or reduced the inflammation in response to the presence of transplanted cells (Fig. 3b) [98]. In a separate study using a complete transection model of thoracic SCI, Lu et al. demonstrated success at increasing transplanted cell viability as well as functional recovery after SCI with the delivery of NPCs in a fibrin matrix compared to cells in medium (Fig. 3c) [120]. It is important to note that in the Lu study, the researchers included a cocktail of growth factors in the pre-hydrogel solution. Therefore, combinations of supportive scaffolds, growth factors and appropriate cells are needed for effectively repairing the injured tissue. Transplanted cells can also serve as a consistent source of growth factors when they are delivered to the CNS [126]. To achieve this, these cells must be protected from the immune system for longterm viability. Some researchers are using microspheres with cells encapsulated inside or adhered to the outside to protect the cells from the immune system upon delivery in vivo [69,127]. Specifically, alginate capsules can shield brain derived neurotrophic factor (BDNF)-producing fibroblasts from the immune system when implanted in a SCI model [128,129]. This is an intriguing option to deliver cells to the injured CNS without immune suppression; this option could provide a viable treatment without additional medication. In using biomaterials to isolate transplanted cells from the immune system, synthetic materials may provide more immune protection than natural polymers because of the absence of bioactive components. Neural stem and progenitor cells have also been transplanted with the intent that these cells will differentiate into specific cell types that can participate in the recovery process. In this case, utilizing biomaterials with natural ECM molecules, matching 336
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mechanical properties, and introducing tissue relevant factors and topographical cues are ways to direct cell differentiation. Hyaluronic acid-based hydrogels with mechanical properties similar to that of brain and spinal cord tissue have been shown to direct NPC differentiation [16]. Moreover, NPCs can also respond to structural cues and align to follow controlled microstructure created within hyaluronic acid hydrogels [130]. In 2009, Christopherson et al. examined the effects of fiber diameter using electrospun polyethersulfone nanofibers coated with laminin on NPC differentiation [131]. In this study, researchers found that under differentiation conditions, adult NPCs exhibited increased oligodendrocyte differentiation on 283 nm fibers and increased neuronal differentiation on 749 nm fibers compared to cells grown on tissue culture plates [131]. Therefore, topographical cues and mechanical cues within hydrogels are additional tools to direct the fate of NPCs. In a different investigation, Lim et al. observed that aligned electrospun polycaprolactone (PCL) fibers significantly increased NPC differentiation toward neurons and alignment when compared to NPCs grown on tissue culture treated plastic or randomly oriented fibers [132]. Effects of fiber diameter on neurons and glia were examined. The authors found that topographical features such as large fiber diameter and aligned fibers can induce apoptosis of glial cells, increasing the percentage of neuronal cells. This is an exciting finding that introduces additional features for creating more defined regions for specific cell types within a hdyrogel; for example, fiber diameter could control the cell phenotype boundary of white matter and gray matter within the spinal cord [132]. It seems clear that currently available biomaterials can increase cell viability, modify the inflammatory response and encourage differentiation of implanted cells. Next generation hydrogels for CNS repair will likely include native ECM components and similar topography to native tissue.
Hydrogels for tissue regeneration The brain and spinal cord consist of highly specific regions and intricate architecture (e.g., the hippocampus and subventricular zone in the brain, white and gray matter of the spinal cord) [133]. These geometrically complex regions have been explored as design criteria for developing biomaterials for the CNS. Many fundamental studies using cultured cells have shown that creating a complex topographical landscape [130] with submicron and nanoscale features [130] can direct neural cell behavior. As mentioned in the previous section, injectable hydrogels have the ability to conform to the shape of irregular cavities, but creating architectures and altering pore sizes within hydrogel is difficult [101]. In vitro studies of electrospun nanofibers showed embryonic stem cell differentiation into neurons, astrocytes and oligodendrocytes [134]. Moreover, the presence of parallel nanofibers enhanced the neurite extension and orientation [134]. Seidlits et al. reported increased neurite outgrowth of hippocampal cells on 3D microstructures created using 2-photon lithography. The authors created sub-micron structures of bovine serum albumin coated with laminin-derived peptides within HA scaffolds that directed neurite extension and orientation [135]. In another investigation, researchers compared hippocampal NPC affinity to polydimethylsiloxane microsctructure to chemical cues (laminin or nerve growth factor) in vitro. NPCs more preferentially extended neurites onto the microstructures than onto non-structured chemically
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FIGURE 3
Hydrogels can be used to deliver cells to an injury site. This has been shown to increase cell viability, decrease immune response, have an effect on cell differentiation, and increase functional recovery. (a) Zhong et al. transplanted neural progenitor cells (NPCs) into the stroke cavity 7 days after focal ischemia stroke induction with or without hyaluronan-heparin-collagen hydrogels [98]. Two weeks after transplantation, transplanted cells were immunostained with anti-green fluorescent protein (GFP) and quantified by stereological optical fractionator procedures. Cell viability significantly increased when cells were transplanted with a hydrogel compared to cells in saline (p = 0.035). The increase in cell viability is attributed to the delivery with a hydrogel scaffold? (b) Zhong et al. also quantified the immune response after NPC transplantation by immunostaining transplanted NPCs with anti-GFP and active macrophages and microglia with anti-Iba1. Activated macrophages and microglia were quantified around the injury site. There was a significantly decreased quantity of infiltrating activated macrophages and microglia in the injury site with composite hydrogel and cell implantation over cell implantation alone (p = 0.004), quantified by stereological optical fractionator procedures. This shows the positive effect of the hydrogel on the injury environment or on the implanted cells ability to alter the injury environment. (c) Lu et al. transplanted NPCs to fill a thoracic level complete transection site 2 weeks following injury [114]. 7 weeks post transplantation, transplanted NPCs (GFP) were observed to be evenly distributed throughout the injury area (1), to differentiate into mature neurons (2), to robustly extend axons beyond the lesion into the caudal spinal cord in gray matter and white matter (3–5), and rats were seen to have significant functional improvement over the animals with saline alone injections (6, p < 0.01). These results support the idea that the presence of hydrogel scaffold during transplantation can distribute cells throughout the transection area as well as provide support for neuronal differentiation of transplanted NPCs.
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coated substrates [136], suggesting that microstructure plays a critical role in cell behavior. In 2010, Baiguera et al. made electrospun PCL into fibrous membranes with micron-scale and submicron-scale features. Hippocampal astroctyes and endothelial cells cultured on micron-scale features allowed cell infiltration whereas submicron-scale features did not [137]. Therefore, determining the correct feature size for a specific application is an important parameter to consider. These fundamental findings from in vitro work have been incorporated in hydrogel-based biomaterials and the resulting hydrogels have been used in vivo with some success [138,141]. Architecture within hydrogels can also be accomplished by introducing a porous network (Fig. 2c). Micro- and macrostructure in hydrogel scaffolds have been created using a number of methods including porogen leaching [138], freeze drying [97,101], crystal templating [139], multiphoton lithography [135] and molding [140]. Porogen leaching and gas foaming create pore sizes relatively consistent with the size of the porogen or gas used to create the pores [102,138]. It is not entirely clear if having a consistent pore size is a critical requirement for successful biomaterial integration but optimum pore sizes have been determined for a number of applications [102]. Martinez-Ramos et al. used porogen leaching to synthesize poly(ethyl acrylate)–poly (hydroxyl ethyl acrylate) copolymer blend matrices with two different geometries for implantation near the subventricular zone (SVZ) of the brain [138]. Pre-clinical in vivo implants of hydrogels with aligned 40 mm diameter pores and with interconnected orthogonal 80 mm pores showed local angiogenesis and neurite extension within each of the scaffolds in the SVZ [138]. In in vitro experiments, NPCs readily infiltrated the scaffold and differentiated into neurons and astrocytes regardless of the scaffold geometry [138]. It appears from this study that specific scaffold geometry does not play a significant role in scaffold integration or cell penetration in the brain; it is likely that the overall porosity plays the largest role. In another study, PCL scaffolds were created with two kinds of internal architectures using a molding method: (1) longitudinal channels with microgrooves within a cylinder, or (2) orthogonally intersecting channels and axial microgrooves within a cylinder [141]. When implanted into adult rat cortex, the group with orthogonal channel implants that more closely mimic the native ECM showed the most tissue in-growth compared to the longitudinal channel implant or control scaffold implant groups. It is important to note that in this study the pore sizes were hundreds of microns. Therefore, when the pores are larger than 100 mm the internal architecture may play a more significant role in tissue ingrowth from the host. In 2005, Tian et al. implanted freeze-dried hyaluronic acid-poly-D-lysine composite gels with mean pore diameters ranging from 90 to 230 mm in a TBI model in adult rats. These porous scaffolds with internal architecture (fiber-like struts and sheet-like features) showed successful integration with host tissue, endothelialization and new ECM formation by six weeks [97]. Macro-architecture that mimics the natural tissue is thought to encourage regeneration. Repairing the spinal cord presents an interesting challenge to biomaterials scientists and engineers in that the white and gray matter architectures are dramatically different. Depending on the injury type and location, different architectures may be more relevant. The white matter contains 338
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well-aligned axon tracts, whereas the gray matter houses local motor neurons and interneuron pools and would likely require less directed growth. Friedman et al. theorized that creating architecture with specific features to mimic spinal cord tracts would encourage regeneration [142], however, the suggested designs have not been fully developed. Hydrogel scaffolds with macroarchitecture have been explored recently. Stokols et al. designed templated agarose scaffolds for use after dorsal column lesions. The extrusion method created aligned linear cylindrical structure for white matter regeneration. The authors observed host integration including Schwann cell infiltration and vascularization in implants. Further, the researchers noted a minimal increase in inflammation near the tissue-injury interface compared to untreated animals. Axon extension was observed only when Matrigel1 or fibrin matrices were included within the implants [143], suggesting that natural ECM cues are required for axonal growth within the graft. Hydrogel scaffolds with slightly more complex architecture have been explored recently to repair the spinal cord. Wong et al. cast five PCL scaffolds with different architectures and implanted them into a complete transection thoracic spinal cord. The authors compared three designs with non-specific architecture and two designs with specific architecture that more closely mimics white and gray matter of the spinal cord. At three months, non-specific architecture implants had higher inflammatory response (increased macrophages, fibroblasts, astrocytes) as well as lesion expansion compared to the structure-mimicking implants. Scaffolds with the structure-mimicking designs also encouraged more axonal infiltration, astrocyte migration and myelinated axon growth into the material compared to the other designs [140]. In this study, meaningful functional recovery was not assessed, but it provided a strong argument in support of anatomically relevant macro-architectural cues for enhancing regeneration. In summary, when it comes to incorporating structure into hydrogel scaffolds, researchers must determine the most relevant biological features, and size scale and determine how to create similar architectures within biomaterials. Although, there are new scaffolds being tested in pre-clinical models in the laboratory, currently there are none available in the clinic. Future work in this area needs to focus on fine-tuning the optimal morphological layout, length scales, and feature sizes. Molding and multiphoton lithography show the most promise as methods to create robust, relevant geometry within hydrogels because they provide precise control over macro- and microarchitecture. It is clear from fundamental studies that topography can play a significant role in directing cell behavior. Therefore, biomaterials that better mimic the native architecture will continue to be an important area of study for CNS regeneration.
Concluding remarks Looking ahead, advances being made in the field of biomaterial science will significantly impact the next generation of therapies for repair in the CNS. Hydrogels in particular are strong candidates for clinical relevance in the future. Indeed, as mentioned in this short review, there are a number of hydrogel genres that have been used successfully in the CNS in the laboratory. However, before any biomaterial can be used in the clinic, we must consider thoroughly how materials interact with surrounding cells and
host tissue in vivo. Key considerations and material features that we feel are important include: (1) a better understanding of the degradation products of parent polymers used for drug delivery applications, (2) equal consideration of the possible routes of delivery including nasal, intrathecal and direct implantation, (3) careful examination of possible biological effects of the hydrogels and their degrading products, (4) inclusion of native ECM components within the scaffolds during cell delivery for better cell survival and integration with the host tissue and (5) addition of biologically relevant topography and architectural features within scaffolds. Significant advances have been made to improve both functional recovery and cell viability following SCI. Additionally, progenitor cell differentiation and the reduction in inflammation as a result of hydrogel implantation are notable achievements in this area of research. Future efforts should focus on a more comprehensive approach to developing and studying biomaterials which include relevant in vivo models.
Acknowledgments The authors wish to acknowledge related funding from the Mission Connect (#012-113 Z.Z.K. and C.E.S.), The Craig Neilsen Foundation (#222456, C.E.S.), NSF-CBET (#1159774 C.E.S.), NSFDMR (#0805298 C.E.S.), NIH R21 (#R21 NS074162 C.E.S.), NSFGFRP (#2011112479, S.A.G.), and the National GEM Consortium (R.C.T.). References [1] [2] [3] [4] [5] [6] [7] [8] [9] [10] [11] [12] [13] [14] [15] [16] [17] [18] [19] [20] [21] [22] [23] [24] [25] [26] [27] [28] [29] [30] [31] [32] [33] [34] [35] [36] [37] [38] [39]
D.R. Nisbet, et al. J. Biomater. Appl. 24 (1) (2009) 7. J.T.S. Pettikiriarachchi, et al. Aust. J. Chem. 63 (8) (2010) 1143. D. Macaya, M. Spector, Biomed. Mater. 7 (1) (2012) 012001. D.A. McCreedy, S.E. Sakiyama-Elbert, Neurosci. Lett. 519 (2) (2012) 115. K. Straley, S.C. Heilshorn, Conf. Proc. IEEE Eng. Med. Biol. Soc. 2009 (2009) 2101. Y. Zhong, R.V. Bellamkonda, J. R. Soc. Interface 5 (26) (2008) 957. Z.Z. Khaing, C.E. Schmidt, Neurosci. Lett. 519 (2) (2012) 103. H. Kim, et al. Trends Biotechnol. 30 (1) (2012) 55. G. Orive, et al. Nat. Rev. Neurosci. 10 (9) (2009) 682. H. Cao, et al. Adv. Drug Deliv. Rev. 61 (12) (2009) 1055. R. Uibo, et al. Biochim. Biophys. Acta 1793 (5) (2009) 924. O.K. Bitzer-Quintero, I. Gonzalez-Burgos, Neural Plast. 2012 (2012) 348642. R.J. McKeon, et al. J. Neurosci. 11 (11) (1991) 3398. P.C. Georges, et al. Biophys. J. 90 (8) (2006) 3012. A.J. Engler, et al. Cell 126 (4) (2006) 677. S.K. Seidlits, et al. Biomaterials 31 (14) (2010) 3930. H. Liao, et al. Acta Biomater. 4 (5) (2008) 1161. J.A. Burdick, et al. Biomaterials 27 (3) (2006) 452. H.Y. Yang, et al. Mol. Chem. Neuropathol. 21 (2–3) (1994) 155. M.C. Shearer, J.W. Fawcett, Cell Tissue Res. 305 (2) (2001) 267. R.J. Pasterkamp, et al. Mol. Cell. Neurosci. 13 (2) (1999) 143. J. Silver, J.H. Miller, Nat. Rev. Neurosci. 5 (2) (2004) 146. C.M. Liberto, et al. J. Neurochem. 89 (5) (2004) 1092. P.A. Harvey, et al. J. Neurosci. 29 (19) (2009) 6285. W.B. Cafferty, S.M. Strittmatter, J. Neurosci. 26 (47) (2006) 12242. X. Wang, et al. J. Neurosci. 22 (13) (2002) 5505. T. GrandPre, et al. Nature 417 (6888) (2002) 547. J.R. Faulkner, et al. J. Neurosci. 24 (9) (2004) 2143. B. Gardner, et al. Paraplegia 33 (10) (1995) 551. T.M. Miller, et al. Lancet Neurol. 12 (5) (2013) 435. G. Lind, et al. Acta Neurochir. Suppl. 97 (Pt 1) (2007) 57. S. Erdine, J. De Andres, Pain Pract. 6 (1) (2006) 51. B.R. Felice, et al. Toxicol. Pathol. 39 (5) (2011) 879. C.E. Kang, et al. J. Control. Release 144 (1) (2010) 25. A.S. Lossinsky, R.R. Shivers, Histol. Histopathol. 19 (2) (2004) 535. P.J. Johnson, et al. Biotechnol. Bioeng. 104 (6) (2009) 1207. P.J. Johnson, et al. Soft Matter 6 (20) (2010) 5127. S.E. Sakiyama-Elbert, J.A. Hubbell, J. Control. Release 69 (1) (2000) 149. S.J. Taylor, et al. J. Control. Release 113 (3) (2006) 226.
RESEARCH
[40] [41] [42] [43] [44] [45] [46] [47] [48] [49] [50] [51] [52] [53] [54] [55] [56] [57] [58] [59] [60] [61] [62] [63] [64] [65] [66] [67] [68] [69] [70] [71] [72] [73] [74] [75] [76] [77] [78] [79] [80] [81] [82] [83] [84] [85] [86] [87] [88] [89] [90] [91] [92] [93] [94] [95] [96] [97] [98] [99] [100] [101] [102] [103] [104] [105] [106] [107] [108] [109] [110] [111]
J.W. Austin, et al. Biomaterials 33 (18) (2012) 4555. M.D. Baumann, et al. J. Control. Release 138 (3) (2009) 205. M.D. Baumann, et al. Biomaterials 31 (30) (2010) 7631. M.J. Caicco, et al. J. Biomed. Mater. Res. A 101 (5) (2013) 1472. D. Gupta, et al. Biomaterials 27 (11) (2006) 2370. M.J. Cooke, et al. Biomaterials 32 (24) (2011) 5688. W.M. Tian, et al. J. Control. Release 102 (1) (2005) 13. R.K. Mittapalli, et al. Mol. Cancer Ther. 12 (11) (2013) 2389. Y. Wang, et al. Pharm. Res. 28 (6) (2011) 1406. A. Jain, et al. PLoS ONE 6 (1) (2011) e16135. Z. Yang, et al. Biomaterials 31 (18) (2010) 4846. M.S. Shoichet, et al. Prog. Brain Res. 161 (2007) 385. C.E. Kang, et al. Tissue Eng. Part A 15 (3) (2009) 595. A.J. Mothe, et al. Biomaterials 34 (15) (2013) 3775. C.E. Kang, et al. Cells Tissues Organs 197 (1) (2013) 55. M.D. Wood, et al. J. Biomed. Mater. Res. A 89 (4) (2009) 909. S.J. Taylor, S.E. Sakiyama-Elbert, J. Control. Release 116 (2) (2006) 204. A.J. Hyatt, et al. J. Control. Release 147 (1) (2010) 24. P.J. Johnson, et al. Cell Transplant. 19 (1) (2010) 89. S.M. Willerth, et al. Stem Cells 25 (9) (2007) 2235. S.M. Willerth, et al. Stem Cell Res. 1 (3) (2008) 205. A. Jain, et al. Biomaterials 27 (3) (2006) 497. Y. Wu, et al. Eur. J. Pharm. Biopharm. 81 (3) (2012) 486. Y. Wu, et al. Biomaterials 33 (7) (2012) 2351. H. Nazar, et al. Eur. J. Pharm. Biopharm. 77 (2) (2011) 225. A.K. Agrawal, et al. Pharmazie 65 (3) (2010) 188. I.A. Alsarra, et al. Drug Dev. Ind. Pharm. 35 (3) (2009) 352. J. Wu, et al. Biomaterials 28 (13) (2007) 2220. E. Gavini, et al. J. Pharm. Sci. 100 (4) (2011) 1488. N.B. Skop, et al. Acta Biomater. 9 (6) (2013) 6834. E. Gavini, et al. J. Pharm. Sci. 100 (2011) 1488–1502. J. Malmo, et al. PLoS ONE 8 (1) (2013) e54182. M. Yemisci, et al. Methods Enzymol. 508 (2012) 253. A. Mahapatro, D.K. Singh, J. Nanobiotechnol. 9 (2011) 55. A. Hafner, et al. J. Microencapsul. 28 (8) (2011) 807. Y. Cho, R.B. Borgens, Exp. Neurol. 233 (1) (2012) 126. A. Mistry, et al. J. Drug Target. 17 (7) (2009) 543. K. Nagpal, et al. Drug Deliv. 20 (3–4) (2013) 112. K. Nagpal, et al. Int. J. Biol. Macromol. 59 (2013) 20. K. Nagpal, et al. Drug Deliv. 19 (8) (2012) 378. K. Nagpal, et al. Chem. Pharm. Bull. (Tokyo) 58 (11) (2010) 1423. R.G. Thorne, et al. Neuroscience 127 (2) (2004) 481. U. Westin, et al. Eur. J. Pharm. Sci. 24 (5) (2005) 565. N.J. Schaub, R.J. Gilbert, J. Neural Eng. 8 (4) (2011) 046026. C.M. Patist, et al. Biomaterials 25 (9) (2004) 1569. S.A. Chvatal, et al. Biomaterials 29 (12) (2008) 1967. A. Goraltchouk, et al. J. Control. Release 110 (2) (2006) 400. G.E. Rooney, et al. Tissue Eng. Part A 17 (9–10) (2011) 1287. E.Y. Tan, et al. Pharm. Res. 24 (12) (2007) 2297. J.C. Stanwick, et al. J. Control. Release 160 (3) (2012) 666. Y.C. Wang, et al. Biomaterials 29 (34) (2008) 4546. Y.T. Kim, et al. Biomaterials 30 (13) (2009) 2582. P.J. Gaillard, et al. J. Control. Release 164 (3) (2012) 364. G.P. Tang, et al. Biomaterials 24 (13) (2003) 2351. D.P. Ankeny, et al. Exp. Neurol. 170 (1) (2001) 85. R.G. Soderquist, et al. J. Biomed. Mater. Res. A 91 (3) (2009) 719. T. Kanazawa, et al. Biomaterials 34 (36) (2013) 9220. W.M. Tian, et al. Tissue Eng. 11 (3–4) (2005) 513. J. Zhong, et al. Neurorehabil. Neural Repair 24 (7) (2010) 636. J.E. Rinholm, et al. J. Neurosci. 31 (2) (2011) 538. Z.Z. Khaing, et al. J. Neural Eng. 8 (4) (2011) 046033. P.Z. Elias, M. Spector, J. Tissue Eng. Regen. Med. (2012), http://dx.doi.org/ 10.1002/term.1621. N. Annabi, et al. Tissue Eng. Part B: Rev. 16 (4) (2010) 371. K. Whang, et al. Tissue Eng. 5 (1) (1999) 35. S. Woerly, Biomaterials 14 (14) (1993) 1056. R.J. Wade, J.A. Burdick, Mater. Today 15 (10) (2012) 454. S. Kouhzaei, et al. Neurol. Res. 35 (4) (2013) 415. A. Nehrt, et al. J. Neurotrauma 27 (1) (2010) 151. R.B. Borgens, et al. J. Exp. Biol. 205 (Pt 1) (2002) 1. J. Patterson, et al. Mater. Today 13 (1–2) (2010) 14. T. Hadlock, et al. Tissue Eng. 6 (2) (2000) 119. H.E. Olson, et al. Tissue Eng. Part A 15 (7) (2009) 1797.
339
RESEARCH: Review
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RESEARCH: Review
[112] [113] [114] [115] [116] [117] [118] [119] [120] [121] [122] [123] [124] [125] [126] [127]
340
J.D. Kocsis, et al. J. Neurotrauma 21 (4) (2004) 441. H. Itosaka, et al. Neuropathology 29 (3) (2009) 248. D. Lu, et al. Neurosurgery 61 (3) (2007) 596. M. Tohill, et al. Neurosci. Lett. 362 (3) (2004) 200. A. Wang, et al. Biomaterials 32 (22) (2011) 5023. L. Cui, et al. Stem Cells 26 (5) (2008) 1356. A.M. Parr, et al. Neuroscience 155 (3) (2008) 760. S. Kelly, et al. Proc. Natl. Acad. Sci. U.S.A. 101 (32) (2004) 11839. P. Lu, et al. Cell 150 (6) (2012) 1264. Y.J. Ren, et al. J. Bioact. Compat. Polym. 24 (1) (2009) 56. E. Tziampazis, et al. Biomaterials 21 (5) (2000) 511. W. Daly, et al. J. R. Soc. Interface 9 (67) (2012) 202. M.J. Cooke, et al. Soft Matter 6 (20) (2010) 4988. Y. Xiong, et al. Brain Res. (2009) 1263–2183. B. Rihova, Adv. Drug Deliv. Rev. 42 (1–2) (2000) 65. R.M. Hernandez, et al. Adv. Drug Deliv. Rev. 62 (7–8) (2010) 711.
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[128] [129] [130] [131] [132] [133] [134] [135] [136] [137] [138] [139] [140] [141] [142] [143]
C.A. Tobias, et al. J. Neurotrauma 22 (1) (2005) 138. C.A. Tobias, et al. J. Neurotrauma 18 (3) (2001) 287. S.K. Seidlits, et al. Nanomedicine (Lond.) 3 (2) (2008) 183. G.T. Christopherson, et al. Biomaterials 30 (4) (2009) 556. S.H. Lim, et al. Biomaterials 31 (34) (2010) 9031. S.C. Joshi, et al. Int. J. Pattern Recognit. Artif. Intell. 11 (08) (1997) 1317. J. Xie, et al. Nanoscale 2 (1) (2010) 35. S.K. Seidlits, et al. Adv. Funct. Mater. 20 (9) (2010) 1368. N. Gomez, et al. J. R. Soc. Interface 4 (13) (2007) 223. S. Baiguera, et al. J. Mater. Sci. Mater. Med. 21 (4) (2010) 1353. C. Martinez-Ramos, et al. J. Biomed. Mater. Res. A 100 (12) (2012) 3276. S.A. Zawko, C.E. Schmidt, Acta Biomater. 6 (7) (2010) 2415. D.Y. Wong, et al. J. Neurotrauma 25 (8) (2008) 1027. D.Y. Wong, et al. J. Neurosurg. 109 (4) (2008) 715. J.A. Friedman, et al. Neurosurgery 51 (3) (2002) 742. S. Stokols, et al. Tissue Eng. 12 (10) (2006) 2777.