Fisher John P. e.a. (ed.) Tissue Engineering

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19-8 Tissue Engineering

on surfaces using texturized silicone stamps dipped in protein, dried, and transferred onto chemically

modiﬁed substrates [85,86]. Microfabrication techniques have also been used in the photolithographic

patterning of organosilanes to silicon wafers [75,76], quartz, and standard glass [82,87]. Several chemicals

and proteins have been employed to create regions of two-dimensional patterns of adhesive domains:

laminin [88]; nerve growth factor [89]; ﬁbronectin; collagen; albumin; and laminin paired with other

chemicals including laminin–denatured laminin, laminin–albumin, polylysine-conjugated laminin [90],

and laminin–collagen [91]. Results from cell experimentation have suggested that biochemical patterning

might play a stronger role in inducing cellular response (i.e., attachment, spreading, and alignment) than

topography. Therefore, efforts have been made to combine multiple guidance cues by chemically pattern-

ing three-dimensional substrates [48,69,70]. For more information on the topographical and biochemical

patterning of scaffolds see Flemming et al. [92]; Curtis and Wilkinson [93]; Craighead et al. [73]; and

Bhatia and Chen [84].

19.4.2.2 Matrices Within Polymer Conduits

Neural tissue-engineering applications focus on mimicking the nerve and the supporting extracellular

matrix (ECM) in order to repair or regenerate axons following damage or disease. Experimentation

with ECM molecules [94] and the tailoring of matrices within conduits has led to support of axonal

regrowth following injury. Evidence has been presented that the basement membrane protein laminin

provides pathways of adhesiveness in both peripheral and central nervous system tissues [95]. This ECM

protein is capable of initiating and supporting neurite extension on glass and biodegradable polymers

of poly (lactide-co-glycolide) and poly(dl-lactide) [68,87] as well as glial cell outgrowth on polystyrene

[48] (Figure 19.3). Agarose gels derivatized with laminin oligopeptides have enabled three-dimensional

neurite outgrowth in vitro from cells containing receptors to the laminin peptides [96]. Regeneration of

transected peripheral nerves was enhanced using agarose gels having speciﬁc laminin peptides inside the

scaffold lumen [97,98]. These gel matrices provide support, create an environment that supports growth

and incorporate materials that alter surface area. Collagen, ﬁbronectin, and ﬁbrin have also been used to

enhance cell-substrate interaction [55–57,99–103]. Gels using magnetic ﬁelds have been shown to orient

ﬁbers of collagen within the gels. Compared to randomly oriented ﬁbers, these gels promoted neurite

extension both in vitro and in vivo [56]. Magnetically aligned ﬁbrin gels (MAFGs) having different ﬁbril

diameters but similar alignment (Figure 19.4) resulted in drastic changes in the contact guidance response

of neurites. In gels formed in 1.2 mM Ca

and having a smaller ﬁbril diameter, there was no response

from chick dorsal root ganglia. However, a strong response in gels formed in 12 and 30 mM Ca

with

a larger ﬁbril diameter enhanced neurite length twofold [103]. Inosine, a purine analog that promotes

axonal extension, was loaded into PLGA conduits for controlled release during sciatic nerve regeneration.

Inosine-loaded PLGA foams were fashioned into cylindrical nerve guidance channels using a novel low-

pressure injection molding technique. After ten weeks, a higher percentage cross-sectional area composed

of neural tissue was found in the inosine-loaded conduits compared with controls [104]. Furthermore,

nerve conduits can be ﬁlled with speciﬁc bioactive molecules to elicit new axonal growth following injury.

Such molecules including axon guidance and pathﬁnding molecules (netrins, semaphorins, ephrins, Slits)

[105], cell adhesion molecules (CAMs) that promote neurite growth (NCAM, L1, N -cadherin, tenascin)

[106,107], as well as proteins involved in synaptic differentiation (agrin, laminin beta 2, and ARIA) [108]

have numerous applications for nerve regeneration in vivo.

Synthetic hydrogels have served as artiﬁcial matrices for neural tissue reconstruction, for the delivery of

cells and for the promotion of axonal regeneration required for successful neurotransplantation. Cultured

neurons were found to attach to hydrogel substrates prepared from poly (2-hydroxyethylmethacrylate)

(PHEMA) but grow few nerve ﬁbers unless ﬁbronectin, collagen, or nerve growth factor was incorpor-

ated into the hydrogel. This provides a mechanism to provide controlled growth on hydrogel surfaces

[109]. Hydrogels have been created with bioactive characteristics for neural cell adhesion and growth

[110]. Arg–Gly–Asp (RGD) peptides were synthesized and chemically coupled to the bulk of poly

(N -(2-hydroxypropyl) methacrylamide) (PHPMA) based polymer hydrogels. These RGD-grafted poly-

mers implanted into the striata of rat brains promoted and supported the growth and spread of glial tissue

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Nerve Regeneration 19-9

FIGURE 19.3 Astrocytes cultured on a laminin (LAM)-coated(0.01 mg/ml in EBSS) PS substrate. On the 10/20/3 µm

LAM-PATTERN/LAM–NO PATTERN substrate, astrocytes are aligned in the direction of the groove on the patterned

side(groovesat90

◦

; right of the arrows) while astrocytes were oriented randomly on the nonpatterned side (left of the

arrows) of the substrate. Astrocytes were stained with CFDA SE in cell suspension prior to seeding. Images were taken

from PS substrate ﬁxed 24 h after seeding. Scale bar = 30 µm. (Reproduced from Recknor, J.B. et al. Biomaterials, 25,

2753–2767, 2004. With permission.)

(a)

(b) (c)

FIGURE 19.4 Confocal ﬂuorescence images of aligned ﬁbrin gels for varied Ca

concentration. Representative

confocal images are shown along with the calculated ﬁbrin ﬁbril alignment parameter for magnetically aligned ﬁbrin

gels (MAFGs) formed for three different Ca

concentrations used in the ﬁbrin-forming solution: (a) 1.2 mM,

(b) 12 mM, and (c) 30 mM. Fibril diameter and inter-ﬁbril spacing (porosity) is seen to increase with increasing Ca

concentration. (Reproduced from Dubey, N., Letourneau, P.C., and Tranquillo, R.T., Biomaterials, 22, 1065–1075,

2001. With permission.)

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19-10 Tissue Engineering

onto and into the hydrogels [111]. Cultured Schwann cells, neonatal astrocytes or cells dissociated from

embryonic cerebral hemispheres were also dispersed within PHPMA hydrogel matrices and found to pro-

mote cellular ingrowth in vivo [112]. These polymer hydrogel matrices were found to have neuroinductive

and neuroconductive properties and the potential to repair tissue defects in the central nervous system

by promoting the formation of a tissue matrix and axonal growth by replacing lost tissue [47,113–115].

Furthermore, in the injured adult and developing rat spinal cord, these biocompatible porous hydrogels

(NeuroGels) promoted axonal growth within the hydrogels, and supraspinal axons migrated into the

reconstructed cord segment [116].

19.4.2.3 Neurotrophins

Neurotrophins are proteins in the nervous system that regulate neuronal survival and outgrowth, synaptic

connectivity and neurotransmission. These growth-promoting factors are used to functionalize guidance

conduits and create a desired response from the regenerating neural environment. Manipulation of

polymer conduits for growth factor administration may be a useful treatment for neurodegenerative

diseases, such as Alzheimer’s disease or Parkinson’s disease, which are characterized by the degeneration

of neuronal cell populations. There is also much potential for overcoming severe tissue loss using growth

factors released from nerve conduits in cases where there is a large gap as a result of axotomy. Various

neurotrophic factors, including nerve growth factor (NGF), brain- derived neurotrophic factor (BDNF),

neurotrophin-3 (NT-3), and neurotrophin-4/5 (NT-4/5) as well as other important growth factors and

cytokines, including ciliary neurotrophic factor (CNTF), glial cell line-derived growth factor (GDNF),

and acidic and basic ﬁbroblast growth factor (aFGF, bFGF), have been “trapped” inside polymer conduits

that control their release. For further information on these neurotrophic factors and their effects on neural

regeneration, refer to Blesch et al. [117]; Jones et al. [119]; Terenghi [118]; Stichel and Muller [120].

Nerve growth factor (NGF) was one of the earliest neurotrophic factors identiﬁed and is one of the

most thoroughly studied neurotrophins. In early experiments incorporating NGF into silicone chambers,

the effects of NGF on nerve regeneration were positive but limited. This was attributed to the rapid

decline in NGF concentrations in the conduit due to degradation in aqueous media and leakage from

the conduit [121]. The method of delivery of these factors is a challenge and such limitations were

overcome by providing controlled release of NGF. Controlled-release polymer delivery systems may be

an important technology in enabling the prevention of neuronal degeneration, or even the stimulation

of neuronal regeneration, by providing a sustained release of growth factors to promote the long-term

survival of endogenous or transplanted cells [122]. Polymeric implants providing controlled release of

NGF for one month were developed and found to improve neurite extension in cultured PC12 cells [123].

Continuous delivery of NGF has been shown to increase regeneration in both the PNS [124,125] and

the CNS [126,127]. Furthermore, in an effort to readily provide for prolonged, site-speciﬁc delivery of

NGF to the tissue, without adverse effects on the conduit, biodegradable polymer microspheres of poly

(l-lactide) co-glycolide containing NGF were fabricated. Biologically active NGF was released from the

microspheres, as assayed by neurite outgrowth in a dorsal root ganglion tissue culture system [128–130].

NGF co-encapsulated in PLGA microspheres along with ovalbumin was found to be bioactive for over

90 days [131]. Sustained release of NGF within nerve guide conduits has also been tested. NGF release from

biodegradable poly(phosphoester) microspheres produced using a double emulsion technique exhibited

a lower burst effect but similar protein entrapment levels and efﬁciencies when compared with those

made of PLGA [132]. These NGF-loaded poly(phosphoester) microspheres were successfully implanted

to bridge a 10 mm gap in a rat sciatic nerve model. Furthermore, the exogenous NGF had long-term

morphological regeneration effects in the sciatic nerve [133].

Basic ﬁbroblast growth factor (b-FGF) has been shown to enhance the in vitro survival and neurite

extension of various types of neurons including dorsal root ganglia. One of the earliest studies involved

controlled release of b-FGF and alpha-1 glycoprotein (α1-GP) from synthetic nerve guidance channels

fabricated using the dip molding technique. After an initial burst in the ﬁrst day, linear release was obtained

from the conduits for a period of at least 2 weeks afterward. Only the tubes releasing b-FGF or b-FGF

and alpha 1-GP displayed regenerated cables bridging both nerve stumps, which contained nerve fascicles

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Nerve Regeneration 19-11

with myelinated and unmyelinated axons [134]. Biodegradable polymer foams modiﬁed with a-FGF and

used for controlled release and the provision of a permissive environment for spinal cord regeneration

were formed using freeze-drying techniques [135]. Furthermore, evidence has been shown that a-FGF

and b-FGF promote angiogenesis and may aid in the repair of damaged nerves [136].

Other neurotrophic factors such as GDNF, BDNF, and NT-3 have been released from synthetic guidance

channels. In an effort to study facial nerve regeneration, the effects of the cytokine growth factor, GDNF,

and NT-3 on nerve regeneration were assessed after rat facial nerve axotomy. Nerve cables regenerated

in the presence of GDNF showed a large number of myelinated axons while no regenerated axons were

observed in the absence of growth factors, demonstrating that GDNF, as previously described for the

sciatic nerve, a mixed sensory and motor nerve, is also very efﬁcient in promoting regeneration of the facial

nerve, an essentially pure motor nerve [137]. Exogenous BDNF and NT-3 were delivered simultaneously

into Schwann cell-grafted semipermeable guidance channels by an Alzet minipump to test the ability of

these neurotrophins to promote axonal regeneration. This novel experiment elegantly demonstrated that

BDNF and NT-3 infusion enhanced propriospinal axonal regeneration and enhanced axonal regeneration

of speciﬁc distant populations of brain stem neurons into grafts in the adult rat spinal cord [138].

A pharmacotectonics concept, in which drug-delivery systems were arranged spatially in tissues to

shape concentration ﬁelds for potent agents, has been presented. NGF-releasing implants placed within

1 to 2 mm of the treatment site enhanced the biological function of cellular targets, whereas identical

implants placed approximately 3 mm from the target site of treatment produced no beneﬁcial effect [139].

Due to certain limitations with controlled-delivery systems, alternatives such as the encapsulation of cells

that secrete these factors are discussed in the next section.

19.4.3 Cellular Modiﬁcations

Due to certain limitations with the control of growth factor delivery systems, cells have been manipulated

for the direct delivery of certain neurotrophic factors using a variety of therapeutic strategies. Genetic

engineering has been used to modify cells for neurotrophic factor delivery [140]. (For a review of gene

therapy, refer to Tresco et al. [141] and Tinsley and Eriksson [142]) Cells that produce speciﬁc neuro-

trophic and growth factors have been encapsulated using polymeric biomaterials. Other experiments have

involved the direct seeding of cells known to secrete neurotrophic factors, such as Schwann cells, olfactory

ensheathing cells (OECs) and neural stem cells (NSCs), into nerve conduits.

19.4.3.1 Cell Encapsulation

Implanting polymer-encapsulated cells for secreting growth and neurotrophic factors has been used

for treatment for neurodegenerative disorders and to promote nerve regeneration. As an experimental

therapy for Parkinsonian patients, enhanced beneﬁt from neural transplantation can be provided through

the combination of grafting with trophic factor treatment. This strategy ultimately results in improved

survival and growth of grafted embryonic dopaminergic neurons. It has been demonstrated that the

implantation of polymer-encapsulated cells genetically engineered to continuously secrete glial cell line-

derived neurotrophic factor to the adult rat striatum improves dopaminergic graft survival and function.

This shows that polymer encapsulation of cells can be used as an effective vehicle for long-term trophic

factor supply [143]. A number of proteins have speciﬁc neuroprotective activities in vitro;however,the

local delivery of these factors into the central nervous system over the long term at therapeutic levels has

been difﬁcult to achieve. Direct administration at the target site is a logical alternative, particularly in the

central nervous system, but the limits of direct administration have not been deﬁned clearly. For instance

NGF must be delivered within several millimeters of the target to be effective in treating Alzheimer’s

disease [139]. Cells engineered to express the neuroprotective proteins, encapsulated in immunoisolation

polymeric devices and implanted at the site of lesions have the potential to alter the progression of

neurodegenerative disorders. The polymers used for encapsulation should allow transport of nutrients

and oxygen to the cells, but also afford immunoprotection. Long-term cell viability in vivo in these

constructs due to diffusional limitations has been the major drawback of this approach.

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19-12 Tissue Engineering

Ciliary neurotrophic factor (CNTF) decreases naturally occurring and axotomy-induced cell death and

has been evaluated as a treatment for neurodegenerative disorders such as amyotrophic lateral sclerosis

(ALS) and Huntington’s disease [144]. Effective administration of this protein to motoneurons has been

hampered by the exceedingly short half-life of CNTF, and the inability to deliver effective concentration

into the central nervous system after systemic administration in vivo. BHK cells stably transfected with a

plasmid construct containing the gene for human or mouse CNTF were encapsulated in polymer ﬁbers

and found to continuously release CNTF and slow down motoneuron degeneration following axotomy

[145]. Implantation of polymer-encapsulated cells genetically engineered to continuously secrete GDNF

to the adult rat striatum was found to improve dopaminergic graft survival and function [143]. Therefore

cell encapsulation is a potentially important method in nerve regeneration, and can be used alone or

in conjunction with other methods such as entubulization. For a review on the microencapsulation of

neuroactive compounds and living cells producing these substances, see Maysinger and Morinville [146].

19.4.3.2 Cell Implantation

19.4.3.2.1 Schwann Cells

Schwann cells play an important role in supporting axonal regeneration after damage or disease. These

cells clean debris from the injury site and secrete regulatory proteins that aid in neuronal survival and

axonal growth. In the PNS and CNS, Schwann cells organize the regenerating environment through the

myelination of axons, production of ECM, CAMs, and neurotrophins as well as aiding in the guidance of

regenerating axons. The use of PNS grafts in the CNS by David and Aguayo [14] in the early 1980s distin-

guished Schwann cells as essential for CNS repair [14]. Conduits incorporating these cells have enhanced

PNS regeneration [147]. CNS regeneration has also been induced after implantation of the semipermeable

polyacrylonitrile/polyvinylchloride (PAN/PVC) and Matrigel conduits seeded with Schwann cells into the

transected rat spinal cord [26,148]. Poly (α-hydroxy acids) with seeded Schwann cells or Schwann cell

grafts were also found to be effective candidates for spinal cord regeneration [22,149]. However, limited

regeneration has been demonstrated between the implant and the distal end of the host spinal cord [150]

and myelination beyond the injury site has not been observed [151]. Recent research has shown that

rolled Schwann cell monolayer grafts implanted into the transected rat sciatic nerve increased functional

regeneration compared to acellular controls [152]. Research applicable to human applications has demon-

strated that implantation of human Schwann cells in the nude (T cell deﬁcient) rat spinal cord allowed

axonal growth across the graft and re-entry into the spinal cord. For reviews of these and other Schwann

cell therapies, see Jones et al. [119] and Bunge [150].

19.4.3.2.2 Olfactory Ensheathing Cells

Olfactory bulb ensheathing cells (OECs) are the primary glial cells found in both the PNS and CNS

of the olfactory system. This system differs from other CNS tissues in that axons continue to grow

throughout adulthood. These cells share characteristics with astrocytes of the CNS and Schwann cells

from the PNS. Like astrocytes, these cells express glial ﬁbrillary acidic protein (GFAP), yet they ensheath

and myelinate axons and support axonal regrowth, which are features of Schwann cells. Most work in

this area has involved OEC transplantations performed to promote remyelination of demyelinated rat

spinal cord axons [153,154] and foster regeneration of damaged axons in the mature CNS [27,155–157].

Incorporating OECs into conduits has promotedaxonalregeneration in both the PNS [158] and CNS [156]

using Schwann cell-ﬁlled guidance channels. The results from these transplantations have demonstrated a

distinct advantage of these cells over Schwann cells in creating a regenerative environment within the CNS.

19.4.3.2.3 Neural Stem Cells

Neural stem cells (NSCs) that have the potential to produce new neurons and glia are present in the mam-

malian CNS. These cells can remain in certain regions of the adult CNS after development even though

neurogenesis no longer occurs in most areas after birth. Neural stem cells are described as generating neural

tissue or being derived from the neural system, having capacity for self-renewal, and they are multipotent

or possess the ability to adopt a variety of cellular fates. Neural progenitors with more limited capacities

in terms of growth and differentiation have been known to proliferate throughout life in a variety of

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Nerve Regeneration 19-13

Outer scaffold

Inner

scaffold

(a)

(e)

(c)

(d)

(b)

3 mm

4 mm

1.5 mm

FIGURE 19.5 (a) Schematic of the scaffold design showing the inner and outer scaffolds. (b) and (c) Inner scaffolds

seeded with NSCs. (Scale bars: 200 and 50 µm, respectively.) The outer section of the scaffold was created by means

of a solid–liquid phase separation technique that produced long, axially oriented pores for axonal guidance as well as

radial pores to allow ﬂuid transport and inhibit the ingrowth of scar tissue. (d) Scale bar, 100 µm. (e) Schematic of

surgical insertion of the implant into the spinal cord. (Reproduced from Teng, Y.D. et al., Proc. Natl Acad. Sci. USA,

99, 3024–3029, 2002. With permission.)

mammalian species, including humans [159,160]. NSCs were seeded into a dual scaffold structure made

of biodegradable polymers to address the issues of spinal cord injury. Unique biodegradable polymer scaf-

folds were fabricated where the general design of the scaffold was derived from the structure of the spinal

cord with an outer section that mimics the white matter with long axial pores to provide axonal guidance

and an inner section seeded with neural stem cells for cell replacement and mimic the general character of

the gray matter (Figure 19.5) [161]. The seeded scaffold improved functional recovery as compared with

the lesion control or cells alone following spinal cord injury. Implantation of the scaffold-neural stem

cells unit into an adult rat hemisection model of spinal cord injury promoted long-term improvement in

function that was persistent up to one year in some animals, relative to a lesion-control group [18].

Human embryonic stem (hES) cells hold promise as an unlimited source of cells for transplantation

therapies [162]. However, control of their proliferation and differentiation into complex, viable 3D tissues

is challenging. Combining physical support with chemical cues created a supportive environment for the

control of differentiation and organization of hES cells. Langer et al. developed biodegradable poly(lactic-

co-glycolic acid)/poly(l-lactic acid) polymer scaffolds to promote hES cell growth and differentiation and

formation of 3D structures. Complex structures with features of various committed embryonic tissues

were generated in vitro using early differentiating hES cells and using the supportive three-dimensional

environment to further induce differentiation. Growth factors such as retinoic acid, transforming growth

factor β activin-A, or insulin-like growth factor directed hES cell differentiation and organization within

the scaffold resulting in the formation of structures with characteristics of developing neural tissues,

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19-14 Tissue Engineering

cartilage, or liver as well as the formation of a 3D vessel-like network. These constructs were transplanted

into severe combined immunodeﬁcient mice and continued to express speciﬁc human proteins in deﬁned

differentiated structures. This recent study presents a novel mechanism for creating viable human tissue

structures for therapeutic applications [20]. For a review on stem cells in tissue engineering, refer to

Bianco and Robey [163].

19.5 Conclusions

Engineering regeneration in the nervous system presents many challenges. Many strategies that may

enhance regeneration in the PNS cannot be applied directly to the CNS due to the complexity of the envir-

onment. Novel tissue-engineering strategies and mechanisms, including the use of three dimensional

polymer constructs with or without biological components (i.e., cells) and products fabricated for the

induction of speciﬁc responses (i.e., regeneration), and the manipulation of biological cells in vitro (i.e.,

stem cells or cells for neuronal support), hold much promise for the enhancement of functional repair

and replacement of tissue function. The successful use of polymeric nerve conduits in facilitating peri-

pheral nerve regeneration has been demonstrated and polymers have shown great promise in addressing

spinal cord injuries as well. This regeneration process, with various polymers, both degradable as well as

nondegradable, has been enhanced further by promoting directed growth and by the addition of chemical

cues such as ECM molecules, nerve growth factors and neurotrophins and other agents incorporated in

the conduits to be released in a controlled fashion. Polymers have also played an important role in encap-

sulating cells and in the transplantation of neuronal support cells that release factors to promote nerve

regeneration. Multidiscliplinary tissue-engineering approaches involving such biomaterials to mimic the

native neural environment of the body have resulted in signiﬁcant progress in gaining some understanding

of the systems and signals involved in nerve regeneration. Incorporating multiple guidance cues and exper-

imental strategies will allow further opportunities to elucidate the mechanisms behind nerve regeneration

and speciﬁc considerations for the efﬁcient repair of the nervous system.

Acknowledgments

The authors would like to acknowledge NSF (BES-9983735) and USDOD for funding.

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