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Journal of Biomaterials Science 22 (2011) 611–625brill.nl/jbs

Polylysine-Modified PEG-Based Hydrogels to Enhance theNeuro–Electrode Interface

Shreyas S. Rao a, Ning Han a and Jessica O. Winter a,b,∗

a William G. Lowrie Department of Chemical and Biomolecular Engineering,The Ohio State University, 140 W 19th Avenue, Columbus, OH 43210, USA

b Department of Biomedical Engineering, The Ohio State University, Columbus, OH 43210, USA

Received 18 November 2009; accepted 18 January 2010

AbstractNeural prostheses are a promising technology in the treatment of lost neural function. However, poor bio-compatibility of these devices inhibits the formation of a robust neuro–electrode interface. Several factorsincluding mechanical mismatch between the device and tissue, inflammation at the implantation site, andpossible electrical damage contribute to this response. Many researchers are investigating polymeric brainmimetic coatings as a means to improve integration with nervous tissue. Specifically, hydrogels, constructsalso employed in tissue engineering, have been explored because of their structural and mechanical similar-ity to native tissue. However, many hydrogel materials (e.g., poly(ethylene glycol) (PEG)) do not support celladhesion. In this work, we report a technique to enhance the interface between polymeric brain mimetic coat-ings and neural tissue using adhesion molecules. In particular, polylysine-modified PEG-based hydrogelswere synthesized, characterized and shown to promote neural adhesion using a PC12 cell line. In addition,we examined adhesion behavior of a PEG-co-polymer and found that these materials adhere to electrodesfor at least 4 weeks. These results suggest that polylysine–PEG hydrogel biomaterials are biocompatibleand can enhance stability of chronic neural interfaces.© Koninklijke Brill NV, Leiden, 2011

KeywordsNeural prostheses, polylysine, poly(ethylene glycol), hydrogel

1. Introduction

Neurodegenerative disorders (e.g., Alzheimer’s disease, Parkinson’s disease) causesubstantial loss of neural function, affecting over a million individuals everyyear [1]. Unfortunately, central nervous system (CNS) neurons do not regeneratefollowing injury [2]. One strategy for treatment has been the use of tissue engineer-ing to enhance nerve regeneration at the injury site; however, these approaches haveseen only limited success [3]. Alternatively, researchers have developed neural pros-

* To whom correspondence should be addressed. E-mail: [email protected]

© Koninklijke Brill NV, Leiden, 2011 DOI:10.1163/092050610X488241

612 S. S. Rao et al. / Journal of Biomaterials Science 22 (2011) 611–625

theses, which restore lost electrical signaling by converting an external signal intoan electrical pattern transmitted to cells via microelectrodes [4]. Neural prostheseshave had some success, most notably the cochlear implant [5]. However, severalchallenges, including difficulty in accessing target neuronal tissue [6] and scarringresulting from the host immune response to implantation [7], have prevented theformation of a robust neuro–electrode interface.

For the last several years, our group [6, 8, 9] and others [10–12] have examinedthe possibility of combining tissue-engineering approaches with neural prosthesesto create a composite material that promotes integration with target neural tissue.Several biomaterials have been developed as electrode coatings, including con-ducting polymers [13, 14], layer-by-layer (LbL) assemblies [15] and polymerichydrogels [6, 8]. Among these, hydrogel-based coatings have attracted consid-erable attention because their structural and mechanical properties mimic nativebrain tissue. Hydrogels are cross-linked, hydrophilic polymeric structures that swellin solution [16]. Specifically, hydrogels have been employed as drug eluting de-vices (e.g., for neurotrophin delivery) [6, 17, 18] for improving electrode–tissueintegration. Unfortunately, many of these drugs (i.e., neurotrophins) have a veryshort half-life in vivo [19]. Additionally, efficacy is compromised after the sup-ply of drugs is exhausted. Therefore, additional cues will be needed to sustain andimprove electrode–tissue integration. In the peripheral nervous system (PNS), re-generation occurs not only because of soluble factors but also because of adhesionmolecules (i.e., proteins/peptides that promote neural adhesion) [2]. Further, adhe-sion molecule-modified hydrogels have been utilized as neural tissue-engineeringconstructs [20]. Adhesion molecules, thus, are attractive candidates for improvingtissue integration with neural biomaterials [21, 22]. These molecules have been pre-viously utilized in combination with conducting polymers [23] as well as with LbLassemblies [15] as neuro-prosthetic coatings, although not with hydrogel systems.

Here, we examined the ability of adhesion molecule-modified hydrogel coatingsto promote adhesion to electrode surfaces, thereby enhancing the neuro–electrodeinterface. Specifically, we investigated poly(ethylene glycol) (PEG)-based hydro-gels modified with the non-native adhesion molecule, polylysine (PL). PEG wasselected as a hydrogel material because of its non-immunogenic properties [24]whereas PL was selected because of its ability to support neuronal adhesion [25].PL-modified PEG diacrylate (PEGDA) hydrogels were developed, characterizedand investigated for their ability to promote neuronal adhesion. Additionally, recog-nizing the fact that PEGDA hydrogels do not normally adhere to hydrophobicelectrode surfaces; we examined the adhesion of PEG–hydroxyacid co-polymers(e.g., lactide, caprolactone) to electrode surfaces. The addition of hydroxyacids alsoprovides a strategy for biodegradation of the coating, so that it is removed from theelectrode surface after the acute immune response to implantation. As proof of con-cept, poly(ethylene glycol)–poly(caprolactone) diacrylate (PEG–PCL) hydrogelswere synthesized and their adhesion to model retinal prosthesis electrodes was char-acterized. Adhesion molecule-modified hydrogel coatings offer great promise to

S. S. Rao et al. / Journal of Biomaterials Science 22 (2011) 611–625 613

enhance electrode-host integration. Furthermore, this technique is broadly applica-ble to any implant material (i.e., stimulating or recording electrodes). Hydrogelscan be tuned to mechanically resemble target tissue and to release soluble factors.The addition of adhesion molecules to this system provides a means to combinechemical and mechanical similarities to native tissue.

2. Materials and Methods

2.1. Preparation of Acryl–PEG–Polylysine (Acryl–PEG–PL)

Polylysine was conjugated to acryl–poly(ethylene glycol)–N-hydroxyl succinimide(acryl–PEG–NHS) (Laysan Bio) using standard approaches [26]. Briefly, polyly-sine (30–70 kDa, Sigma-Aldrich) was dissolved in 50 mM sodium bicarbonate(Sigma-Aldrich) and acryl–PEG–NHS (3.4 kDa, Laysan Bio) was dissolved inDMSO (Sigma-Aldrich) in a 1:2 molar ratio. These solutions were mixed andreacted for approx. 2 h in an ice bath. The resulting sample was then dialyzed(Slide-a-lyzer cassettes, Fisher) overnight and reconcentrated to the original vol-ume in phosphate-buffered saline (PBS) using a centrifugal concentrator (Fisher).For quantification, FITC-labeled poly(L-lysine) (PLL) (30–70 kDa, Sigma-Aldrich)and poly(D-lysine) (PDL) (30–70 kDa, Sigma-Aldrich) in a 1:10 (weight ratio) wereused, whereas PDL was utilized for all other experiments. Different concentrationsof acryl–PEG–PL were prepared by diluting the above stock solution in PBS.

2.2. Preparation and Characterization of Poly(ethyleneglycol)–Poly(caprolactone) (PEG–PCL) Co-polymers

Poly(ethylene glycol)–poly(caprolactone) (PEG–PCL) co-polymers were synthe-sized following Hubbell’s method [27]. Briefly, PEG (0.95–1.05 kDa, Sigma-Aldrich) was azeotropically distilled in anhydrous toluene (Sigma-Aldrich) andreacted with ε-caprolactone (Sigma-Aldrich) (1:2 molar ratio) for approx. 24 h,under an argon blanket at 130◦C in the presence of stannous octanoate (Sigma-Aldrich) catalyst. PEG–PCL co-polymer was then cooled to ambient temperature,dissolved in anhydrous dichloromethane (Sigma-Aldrich), and slowly added tocold hexane to precipitate the product. PEG–PCL co-polymer was collected us-ing Buchner funnel filtration and vacuum dried overnight. This intermediate wasthen subjected to azeotropic distillation in anhydrous toluene and dissolved in an-hydrous dichloromethane in presence of triethylamine (Sigma-Aldrich) (for HClneutralization). Acrylation of PEG–PCL was performed by reaction with acryloylchloride (Sigma-Aldrich) (1:2 molar ratio) in anhydrous dichloromethane (1:4 v/v)added at the rate of approx. 0.20 ml/min. The system temperature was maintainedat −20◦C for approx. 3 h and at ambient temperature for another 60 h. Acry-lated PEG–PCL was purified by removal of TEA-HCl salt via filtration, removalof the solvent by rotary evaporation, and subsequently dissolution in anhydroustetrahydrofuran (Sigma-Aldrich). The product was collected via filtration, dissolvedin anhydrous dichloromethane, precipitated using cold diethyl ether, filtered and


vacuum dried in the presence of P2O5 (Sigma-Aldrich). Characterization of PEG–PCL and acrylated PEG–PCL co-polymers was performed using Fourier transforminfrared spectroscopy (Nicolet 6700 FT-IR spectrometer, Thermo Scientific) andnuclear magnetic resonance spectroscopy (NMR Bruker DPX400).

2.3. Hydrogel Formation and Quantification of Polylysine Attachment

Poly(ethylene glycol)-diacrylate (PEGDA) (3.4 kDa) was procured from LaysanBio. PEGDA and PEG–PCL hydrogels were prepared following Hubbell’s method[27]. Briefly, PEGDA or PEG–PCL was dissolved in PBS at 22 wt% and sup-plemented with 1% (v/v) initiator (0.33 g Irgacure 2959 (Ciba) in 1 ml N-vinyl-2-pyrrolidone (Sigma-Aldrich)). Of this hydrogel solution, 60 µl was placed in a96-well plate and subjected to UV illumination (peak power 11.21 mW/cm2, sam-ple distance approx. 9 mm) for approx. 10 min. Acryl–PEG–PL (approx. 30 µl) withinitiator was then added to the surface of the PEG-based hydrogel, and the samplewas subjected to UV illumination (approx. 10 min) (Fig. 1), conditions which havebeen shown to be acceptable for protein conjugation [28]. Sham hydrogels, consist-ing of PEGDA hydrogels to which unconjugated PL (i.e., PL in PBS) and initiatorwere added, were also formed via photo-polymerization. For PEGDA hydrogels,conjugation efficiency was determined by monitoring the diffusion of unreactedFITC–polylysine over a period of 7 days, against a sham (unconjugated polyly-sine) and a negative control (no polylysine) using a fluorescent microplate reader(excitation wavelength = 485 nm, emission wavelength = 535 nm). The conjuga-tion efficiencies were evaluated using standard concentrations of polylysine. Datawere analyzed using ANOVA and pair-wise comparisons were performed using theTukey–Kramer HSD test (α = 0.05) via JMP statistical software (Version 7). ForPEG–PCL hydrogels, conjugation was confirmed using fluorescence microscopyand compared to PEG–PCL hydrogels with unconjugated FITC–polylysine, as wellas PEG–PCL hydrogels with no polylysine.

2.4. PC12 Cell Culture

PC12 cells (CRL-1721, ATCC) were cultured using Powdered Ham’s F12-K Me-dia (Kaighn’s modification) (Sigma-Aldrich) supplemented with 1.5 g/l NaHCO3,

Figure 1. Schematic of the conjugation procedure.


12.5% horse serum, 2.5% fetal bovine serum (all Sigma-Aldrich) and 1% peni-cillin–streptomycin (Invitrogen). Cells were cultured in an incubator at 37◦C and5% CO2. Medium was exchanged every 2–3 days, and cells were sub-cultivatedweekly before experimentation.

2.5. Quantification of Cell Adhesion Using Calcein–AM Staining

Cell adhesion to modified hydrogel surfaces was quantified using the Live–Deadassay (Invitrogen). Non-fluorescent calcein AM is converted by intracellular es-terases to intensely fluorescent calcein in live cells, whereas EthD-1 permeatesdamaged cell membranes and enhances fluorescence by binding to nucleic acidsin dead cells. Polylysine-modified hydrogels were prepared as described previouslyusing PDL. For cell experiments, all solutions were sterile filtered using a 13 mmsyringe filter, pore size 0.22 µm (Fisher) prior to gelation. Modified hydrogels wereimmersed in sterile PBS for 1 week and then subjected to UV illumination overnight(for sterilization). PC12 cells at approx. 2 × 104 cells/cm2 were seeded onto all hy-drogel surfaces and the positive control (well plate coated with polylysine). Afterincubating for approx. 24 h, all surfaces were washed with sterile PBS thoroughly(3 times). Cells were then stained with EthD-1 (1 µM) and calcein–AM (2 µM) pre-pared in sterile Dulbecco’s PBS (D-PBS) (Sigma). Following incubation for approx.45 min, excess dye was removed using sterile D-PBS (3 times) and fluorescence wasobserved. To quantify cell adhesion, calcein–AM fluorescence was used since fluo-rescence intensity is directly proportional to cell number. Background fluorescencefrom dye entrainment in the hydrogel was subtracted from all measurements byobserving hydrogels stained at the same dye concentrations, but with no cells. Im-ages were captured using an Olympus IX71 inverted optical microscope equippedwith fluorescence filters at 10× magnification. The fluorescence intensities (a.u.)were normalized to the negative control and compared using ANOVA. Pair-wisecomparison was performed using the Tukey–Kramer HSD test (α = 0.05) via JMPstatistical software (Version 7). Tests were completed twice, with replicates of 6 ineach test.

2.6. PEG–PCL Electrode Adhesion

Multielectrode arrays (3 × 5 electrodes) used to investigate adhesion of PEG–PCLhydrogels were generously provided by Dr. Stuart Cogan (EIC Laboratories). Thearrays consisted of iridium oxide-coated gold electrodes patterned on polyimidesubstrates [6]. PEG–PCL hydrogel was formed by depositing approx. 1 µl of theprecursor solution onto the electrode array surface and subjecting this bolus to UVphoto-polymerization. To evaluate the efficacy of the hydrogel coating, static adhe-sion of the coating as well as adhesion following electrode insertion was monitored.For electrode insertion, PEG–PCL hydrogel coated electrodes were inserted into a1 wt% agarose brain tissue phantom and removed immediately. Static integrity ofthe hydrogel coating over time was evaluated by placing coated electrodes under


a 1 wt% agar gel blanket, which physically constrained the hydrogel bolus on theelectrode surface, mimicking the effect of tissue. Reflected DIC images for both thetests were collected using an Olympus BX41 reflected optical microscope. Imageswere converted to grayscale using Adobe Photoshop (Version 10.0).

3. Results

3.1. PL Conjugation to PEG Hydrogels

The extent of PL conjugation to PEG hydrogels was examined using fluorescencemicroscopy to evaluate FITC–PL-conjugated PEG hydrogels (sample), PEG hydro-gels with FITC–PL added but not conjugated to the hydrogel backbone (sham) andPEG-based hydrogels with no PL (negative control). Fluorescence values were nor-malized to the negative control (no FITC–PL, fluorescence intensity = 1). PL con-jugation at different concentrations of acryl–PEG–PL was quantified and in all caseswas higher and statistically significant (P < 0.0001 for 1× and 0.01×, P = 0.0056for 0.1×, ANOVA, Tukey test) for PL-conjugated PEG hydrogels (sample) ver-sus the negative control (Fig. 2). Unconjugated PL (sham) hydrogels were alsoevaluated for comparison. Intuitively, unconjugated PL should be released from

Figure 2. Conjugation of PL with PEGDA hydrogels. (n = 4 for 1×, n = 4 for 0.1×, n = 5 for0.01×). An asterisk indicates statistical significance compared to negative control.


Table 1.Concentration of bound and entrapped PL (in µg/ml, average ± SD)

Dilution of acryl–PEG–PL

1× (n = 4) 0.1× (n = 4) 0.01× (n = 5)

Conjugated + entrapped PL 31.52 ± 5.32 6.21 ± 1.88 0.45 ± 0.11Entrapped PL 9.18 ± 6.77 3.44 ± 0.49 0.28 ± 0.05

the hydrogel over time. However, we observed that some PL becomes physicallyentrapped within the gel matrix and unconjugated PL (sham) hydrogels generallyexhibit fluorescence intensity higher and statistically significant (Tukey test) fromthe negative control. In the case of 0.1× dilution, this value, although higher thannegative control was not statistically significant (power = 0.93 < 0.95). Compar-ing PL-conjugated (sample) and unconjugated (sham) hydrogels, it was observedthat fluorescent intensity of sample hydrogels was higher and statistically signifi-cant from that of sham hydrogels at all dilutions investigated. Exact concentrationvalues (obtained from a standard curve of known concentrations) are summarizedin Table 1.

3.2. PC12 Cell Adhesion on PEG Hydrogels

Cell response to PL-modified PEG hydrogels was assessed using the PC12 cell line,a model for neuronal behavior [29–31]. PC12 cells convert to a neuronal phenotypeafter exposure to nerve growth factor (NGF) [31], including extension of neurites.Because the primary focus of this study was cell adhesion and because neurite ex-tension complicates quantification of cell surface area, NGF was excluded from thecell culture medium. Additionally, because this study focused on proof-of-conceptand demonstration of biocompatibility, additional cell types (e.g., astrocytes, mi-croglia) were not investigated. PC12 cell adhesion was quantified using calceinAM staining with fluorescence intensity serving as an indirect measure of cell num-ber.

PC12 cells adhered to PL-modified PEG hydrogels at the lower range of acryl–PEG–PL concentrations investigated. Both ‘sample’ and ‘sham’ hydrogels dis-played statistically significant adhesion (P < 0.0001, ANOVA, Tukey test, power =0.99) from the negative control (Figs 3 and 4). At high concentrations of PL (i.e.,1× dilution), substantial cell death was seen as evidenced by EthD-1 staining (datanot shown). This is not unusual, as PL is known to be cytotoxic at high concen-trations [32]. Although PC12 cell adhesion to the well plate (positive control) wasalso examined, cell adhesion was not compared with ‘sample’, ‘sham’ and negativecontrol hydrogels because it is known that differences in substrate stiffness (i.e.,soft hydrogel vs rigid tissue-culture plastic) strongly influence cell morphology andspreading [33].


Figure 3. PC12 cell response on PL-modified PEGDA hydrogels. (A) Sample, (B) sham, (C) negativecontrol (no PL) and (D) positive control (PL). This figure is published in colour in the online editionof this journal, that can be accessed via http://www.brill.nl/jbs

3.3. PEG–PCL Hydrogel Coatings

Given these promising initial results with PEG hydrogels, we next investigatedresponse to PEG–hydroxyacid (PEG–HA) hydrogel composites (i.e., PEG–PCL).PEG–HA composites have been previously used as neuro-prosthesis coatings thatprovide delivery of soluble neurotrophins [6, 8] to enhance neuronal survival, ex-tension and tissue integration. Additionally, in contrast to PEG hydrogels, PEG–HAcoatings demonstrate much greater adhesion to hydrophobic electrode array sur-faces, which are primarily composed of parylene or polyimide. PEG–HA hydrogelsare also degradable through hydrolysis of the ester linkage, thus providing a meansof controllably removing the coating after the acute immune response. Given atarget of approx. 6–8 weeks for the acute immune response, we investigated slowly-degrading PEG–poly(caprolactone) hydrogels for ability to support PL-conjugationand promote adhesion to neural prosthesis electrode arrays.

3.3.1. Synthesis and CharacterizationPEG–PCL polymers were synthesized and characterized by FT-IR and NMR spec-troscopy. In the FT-IR spectra for PEG–PCL (Fig. 5A), notable peaks were ob-served at 1735 = PCL, COO; 2870 = PEG, CH2; and 3500 cm−1 = PEG, OH.For the diacrylate PEG–PCL, peaks included 1725 = acryl, COO; 1735 = PCL,


Figure 4. Normalized fluorescence for PC12 cell adhesion on experimental PEGDA hydrogels(n = 6). An asterisk indicates statistical significance compared to negative control.

COO; 2870 = PEG, CH2; and 3550 cm−1 = PEG, OH [6, 34, 35]. The reduc-tion of the OH peak (3550, PEG) confirms the creation of the ester bond betweenPEG–PCL and acryloyl chloride following the acrylation step. In the 1H-NMRspectrum for diacrylate PEG–PCL (Fig. 5B), notable peaks were observed at3.64 = PEG, CH2, 4.08, 4.15, 4.23, 4.31 = PCL, CH2 (a, position as shown inchemical structure, Fig. 5C), 1.67 = PCL, CH2 (b), 2.33 = PCL, CH2 [35] (c),5.81, 6.12 and 6.41 ppm = acryl, CH, CH2 [6, 36]. The average number of CLsper side PEG is approx. 0.967 as calculated using yield of PEG–PCL. The yield foracrylated PEG–PCL co-polymer as calculated from NMR spectra using peak inte-gration was approx. 88.4%. PL-conjugated hydrogels were created using UV-photo-polymerization of acryl–PEG–PL segments to the PEG–PCL hydrogel backbone.To confirm conjugation, FITC–PL was used and its fluorescence was observed. Asexpected, PEG–PCL hydrogels with FITC–PL modification exhibit fluorescence,whereas unmodified PEG–PCL hydrogels do not (Fig. 6).

3.3.2. PEG–PCL Electrode AdhesionAdhesion of PEG–PCL hydrogels to an electrode array (retinal prosthesis, Centerfor Innovative Visual Rehabilitation, Fig. 7) was confirmed using insertion–removal


(A)

(B)

(C)

Figure 5. Analysis of PEG–PCL polymer. (A) FT-IR spectroscopy, (B) 1H-NMR spectroscopy and(C) chemical structure.

and agar blanket tests. The insertion–removal test showed that the hydrogel bo-lus remains intact and is minimally influenced by the shear force arising from theelectrode insertion and removal process (Fig. 8, n = 2). Further, the agar blan-ket test showed that PEG–PCL hydrogels can adhere to electrode surfaces for atleast 4 weeks (Fig. 9, n = 3). Degradation of the polymer was observed (evidenced


Figure 6. Conjugation of PL to PEG–PCL hydrogels. (A) PL conjugated/entrapped (sample), (B) PLentrapped in PEGPCL hydrogel and (C) negative control (no PL). This figure is published in colourin the online edition of this journal, that can be accessed via http://www.brill.nl/jbs

Figure 7. Retinal implant electrode.

Figure 8. Coating integrity (A) before and (B) after insertion into an agarose tissue phantom. Blackarrows indicate the edges of the polymer coating.

by decreasing surface area occupied by the hydrogel bolus on the electrode sur-face); however, the rate of degradation was slower when compared with other PEGco-polymers, e.g., poly(ethylene glycol)–poly(lactic acid) [6], permitting hydrogeladhesion over longer time periods.


Figure 9. Static PEG–PCL coating integrity under agarose tissue phantom at (A) 0, (B) 14 and (C) 28days. Black arrows indicate the boundary of the hydrogel coating. Scale bar = 500 µm.

4. Discussion

Here, we investigated the potential of adhesion molecule-modified PEG-hydrogelsto promote neuronal adhesion and thereby improve integration of prosthetic deviceswith the body. In particular, we examined the effect of polylysine (PL), a non-nativebiomolecule known to enhance neuronal adhesion. At almost all concentrations ofPL investigated, conjugated PL (sample) and sham hydrogels displayed higher andstatistically significant PL incorporation from the negative control. This indicatesthat PL can be both bound to the PEG hydrogel (sample) and physically entrappedwithin the hydrogel matrix (sample and sham). In either case, only limited diffusionof PL from the hydrogel matrix is observed, indicating the stability of the adhesionmolecule-modified neural interface. It should be noted that in the case of the 0.1×acryl–PEG–PL dilution, the PL concentration in the sham was higher, although notstatistically significant from that of the negative control. The statistical power ofthis experiment was 0.93, slightly lower than 0.95 and this difference would mostlikely have been detected by increasing the sample size, n.

PC12 cells exhibited two behaviors on these surfaces. At low PL concentrations(i.e., 0.01× dilution), PC12 cell adhesion was improved on PL-modified hydrogelsurfaces when compared with the negative control. At high PL concentrations (i.e.,1× dilution), cell death was observed (data not shown). This is not unexpected asPL at high concentrations is known to be cytotoxic [32], resulting in cell lysis. Thus,the potential for PL release to the body will be important to quantify in future invivo work. PL release from PEGDA gels is limited to desorption of surface boundand lightly entrapped molecules during an initial 5-day period. In this study, allhydrogels were immersed in solution for one week before cell testing to eliminateadsorbed PL. Release from PEGPCL hydrogels would occur as the polymer de-graded and is expected to proceed on the order of weeks [27], a sufficiently slowtime-course to prevent high concentrations of PL from being achieved.

In the low PL concentration region, PC12 adhesion on the conjugated samplewas higher than that to the unconjugated sham, although not statistically significant;indicating that physically entrapped PL was equally able to support neuronal adhe-sion. This could be explained through the mechanism of PL-mediated cell adhesion.PL promotes adhesion via the non-receptor mediated cell binding mechanism; i.e.,positively charged PL attracts the negatively charged cell membrane resulting from


the glycocalyx [21]. It also enhances deposition of serum proteins from medium,thereby indirectly improving cell adhesion. This is in contrast to other adhesionmolecules (e.g., laminin, collagen) that promote cell adhesion via ligand–receptor(LR) [21] interactions. In those cases, conjugation of the adhesion molecule to thehydrogel may be necessary to generate traction as the cell pulls on the adhesionmolecule. Molecules that are not bound, but only ‘entrapped’ within the hydrogel,often do not permit the cell to generate the necessary traction force to support ad-hesion. Additionally, adhesion molecules that employ the LR mechanism primarilypromote adhesion through specific sites (e.g., RGD in the case of collagen, IKVAVin the case of laminin [21]). Steric hindrance of an entrapped molecule can limitaccess to the adhesion site. Our results suggest that because PL supports adhe-sion through an alternative mechanism, conjugation may not be required. However,when employing PL, it is important to consider molecular weight as well as PL con-centration, as cell response may be altered to PL of varying molecular weight andconcentration [32]. It should also be noted that these results may not apply only toneural adhesion, adhesion of other cell types (e.g., astrocytes, microglia) may alsobe encouraged by PL-modified hydrogels; which will be investigated more fully infuture in vivo studies.

We also synthesized degradable PEG–PCL hydrogels and demonstrated PL at-tachment to these surfaces. The mechanism of PL attachment to this hydrogel issimilar to that of PEGDA hydrogels because both hydrogels are formed via photo-polymerization. The acryl terminus of acryl–PEG–PL serves as an ‘anchor’ for PLattachment in both cases. We demonstrated similar PL attachment for both PEGDAand PEG–PCL hydrogels. Additionally, we examined the stability of PEG–PCL hy-drogels on neural prostheses surfaces. We showed that PEG–PCL hydrogels canadhere to electrodes for at least 4 weeks, at which point degradation processesbecome significant. This is an enhancement over previous PEG–PLA hydrogelcoatings that adhered for up to 11 days [6]. Long-term adhesion of the polymerichydrogel is vital for providing sustained biochemical cues that permit neurons toform firm contacts with the electrode array. In addition to adhesion, the hydrogelcoating should withstand shear forces arising from the electrode insertion process.This was examined via the insertion test, and the coating evidenced very little dif-ference before and after insertion into an agarose tissue phantom. It is possiblethat hydrogel adhesion and degradation may be altered by electrical stimulation.Detailed electrical stimulation experiments should yield further insight into this be-havior.

5. Conclusions

PEG-based hydrogels are non-immunogenic, biocompatible and are composed ofFDA approved materials. As such, they hold tremendous promise as coatingsto enhance biocompatibility of implanted devices. Here, we show that adhesionmolecule-modified PEG hydrogels could enhance integration of neural prostheses


with target tissue. However, the PL-modified PEG hydrogel system described hereis a passive coating that promotes neuron attachment by direct contact and is lim-ited by anatomical considerations and proximity to target tissue. In future work,we hope to combine adhesion molecule presenting hydrogels with soluble fac-tor elution to further enhance tissue–device integration. We have already shownthat soluble growth factors can influence neurons distant from the electrode sur-face [6] and hope to combine this approach with PL-modified hydrogels. Adhesionmolecule-modified hydrogels, especially in combination with factor elution, couldenhance the neuro–electrode interface and increase efficacy of current neural pros-theses.

Acknowledgements

The authors would like to thank Professor S.-T. Yang (Chemical & BiomolecularEngineering, Ohio State University, Columbus, OH, USA) for use of his fluores-cent microplate reader and Dr. Stuart F. Cogan (EIC Laboratories, Boston, MA,USA) and the Center for Innovative Visual Rehabilitation (Boston VA Hospital,Boston, MA, USA) for providing electrodes. Financial support from the Ohio StateUniversity and the H. C. ‘Slip’ Slider Professorship (to J. O. W.) is gratefully ac-knowledged.

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