Advanced biomaterials for repairing the nervous system: what can hydrogels do for the brain?

RESEARCH:Review

Materials Today � Volume 00, Number 00 � June 2014 RESEARCH

Advanced biomaterials for repairing thenervous system: what can hydrogels dofor 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 States2McKetta Department of Chemical Engineering, The University of Texas at Austin, United States3Department of Biomedical Engineering, The University of Texas at Austin, United States

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.

IntroductionRecent advances both in our understanding of the nervous system

and the availability of sophisticated biomaterials have significant-

ly 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 pre-

clinical 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 excel-

lent review articles for further studies [1–11]. This review intro-

duces 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

Please cite this article in press as: Z.Z. Khaing, et al., Mater. Today (2014), http://dx.doi.org/

*Corresponding authors:. Khaing, Z.Z. ([email protected]),

Schmidt, C.E. ([email protected])

1369-7021/� 2014 Elsevier Ltd. All rights reserved. http://dx.doi.org/10.1016/j.mattod.2014.05.011

barrier to regeneration. Understanding the specific requirements

and taking them into consideration in designing therapeutic plat-

forms are likely to result in the successful development of next

generation biomaterials. Briefly, the blood barriers, the endoge-

nous immune system within the CNS, and the mechanical prop-

erties 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 particu-

lar, 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 forbiomaterial developmentThere 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 mem-

branes of endothelial cells and astrocytes. In normal physiological

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RESEARCH Materials Today �Volume 00, Number 00 � June 2014

MATTOD-336; No of Pages 9

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.

RESEARCH:Review

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 therapeu-

tics 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 capillar-

ies. 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]. There-

fore, 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 appro-

priate mechanical property is a key factor to consider when using

biomaterials in the CNS.

Injury environment of the nervous systemA 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.


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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 in-

creased 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+ oligoden-

drocyte 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 bre-

vican 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 plas-

ticity is unclear. Moreover, myelin-related glycoproteins have long

been implicated in creating a non-hospitable environment for the

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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 withmicroparticles 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, withand without internal architecture, as scaffolds to deliver cells and repair

injured CNS tissue.

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severed axons of the CNS [24–27]. Combined, these factors con-

stitute 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, cellu-

lar 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 micro-

environment.

Hydrogels as drug delivery platformsAs 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], hyaluro-

nan–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 intra-

thecal 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 erythro-

poietin (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


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 bioac-

tive 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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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 absorp-

tion. 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 chitosan-

based 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 synthet-

ic 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 macro-

porous 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 pro-

duce 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 degrada-

tion products of both synthetic and natural polymers can be

bioactive and, to the best of our knowledge, currently there is


4

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 appli-

cation of lactic acid to cultured slices of developing mouse brain

cortex can support oligodendrocyte development and myelina-

tion [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 environmentHydrogels can be particularly useful in the CNS post injury envi-

ronment 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 partic-

ularly well suited for these types of applications. The liquid poly-

mer 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, chemi-

cal 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, prolifera-

tion, 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 com-

pared 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 re-

sponse compared to control animals [100]. In some cases, PEG

administered following SCI has been shown to help repair dam-

aged 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 (Dura-

SealTM from PEG hydrogel) and natural-based polymers (Dur-

GenTM 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 deliverySource 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,

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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 attri-

butes 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 com-

pared to cells transplanted in culture medium only [125]. In a

separate investigation, thiol crosslinked hyaluronan–heparin–col-

lagen–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 inter-

esting observation because it suggests that the presence of hydro-

gels either modified the post-injury inflammatory environment

(as discussed in the previous section) and/or reduced the inflam-

mation 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 trans-

planted 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 scaf-

folds, growth factors and appropriate cells are needed for effec-

tively 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 long-

term 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]. Specifi-

cally, 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 sup-

pression; this option could provide a viable treatment without

additional medication. In using biomaterials to isolate trans-

planted 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


mechanical properties, and introducing tissue relevant factors

and topographical cues are ways to direct cell differentiation.

Hyaluronic acid-based hydrogels with mechanical properties sim-

ilar 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, Christo-

pherson et al. examined the effects of fiber diameter using elec-

trospun polyethersulfone nanofibers coated with laminin on NPC

differentiation [131]. In this study, researchers found that under

differentiation conditions, adult NPCs exhibited increased oligo-

dendrocyte differentiation on 283 nm fibers and increased neuro-

nal 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 significant-

ly 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 neu-

rons and glia were examined. The authors found that topographi-

cal features such as large fiber diameter and aligned fibers can

induce apoptosis of glial cells, increasing the percentage of neuro-

nal 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 bio-

materials 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 regenerationThe 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 fundamen-

tal 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 architec-

tures 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 micro-

structures created using 2-photon lithography. The authors creat-

ed 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 polydimethyl-

siloxane 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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Please cite this article in press as: Z.Z. Khaing, et al., Mater. Today (2014), http://dx.doi.org/10.1016/j.mattod.2014.05.011

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 ischemiastroke 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 implantedcells 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 hydrogelscaffold 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 electro-

spun PCL into fibrous membranes with micron-scale and submi-

cron-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 meth-

ods 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 bioma-

terial 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 (hydrox-

yl 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 orthogo-

nal 80 mm pores showed local angiogenesis and neurite extension

within each of the scaffolds in the SVZ [138]. In in vitro experi-

ments, 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 longitu-

dinal 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 in-

growth from the host. In 2005, Tian et al. implanted freeze-dried

hyaluronic acid-poly-D-lysine composite gels with mean pore dia-

meters 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


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 architec-

ture with specific features to mimic spinal cord tracts would

encourage regeneration [142], however, the suggested designs

have not been fully developed. Hydrogel scaffolds with macro-

architecture 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 integra-

tion including Schwann cell infiltration and vascularization in

implants. Further, the researchers noted a minimal increase in

inflammation near the tissue-injury interface compared to un-

treated animals. Axon extension was observed only when Matri-

gel1 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 architec-

ture 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 re-

sponse (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 anatom-

ically relevant macro-architectural cues for enhancing regenera-

tion.

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 funda-

mental 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 remarksLooking 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

10.1016/j.mattod.2014.05.011

7




RESEARCH:Review

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 hydro-

gels 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 com-

prehensive approach to developing and studying biomaterials

which include relevant in vivo models.

AcknowledgmentsThe 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.), NSF-

DMR (#0805298 C.E.S.), NIH R21 (#R21 NS074162 C.E.S.), NSF-

GFRP (#2011112479, S.A.G.), and the National GEM Consortium

(R.C.T.).

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Advanced biomaterials for repairing the nervous system: what can hydrogels do for the brain?

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