Multifunctional hydrogel coatings on the surface of neural cuff electrode for improving electrode-nerve tissue interfaces

Accepted Manuscript Multifunctional hydrogel coatings on the surface of implantable cuff electrode for improving electrode-peripheral nerve tissue interfaces Dong Nyoung Heo, Su-Jin Song, Han-Jun Kim, Yi Jae Lee, Wan-Kyu Ko, Sang Jin Lee, Donghyun Lee, Sung Jin Park, Lijie Grace Zhang, Ji Yoon Kang, Sun Hee Do, Soo Hyun Lee, Il Keun Kwon PII: DOI: Reference:

S1742-7061(16)30218-5 http://dx.doi.org/10.1016/j.actbio.2016.05.009 ACTBIO 4241

To appear in:

Acta Biomaterialia

Received Date: Revised Date: Accepted Date:

9 December 2015 18 March 2016 3 May 2016

Please cite this article as: Heo, D.N., Song, S-J., Kim, H-J., Lee, Y.J., Ko, W-K., Lee, S.J., Lee, D., Park, S.J., Zhang, L.G., Kang, J.Y., Do, S.H., Lee, S.H., Kwon, I.K., Multifunctional hydrogel coatings on the surface of implantable cuff electrode for improving electrode-peripheral nerve tissue interfaces, Acta Biomaterialia (2016), doi: http:// dx.doi.org/10.1016/j.actbio.2016.05.009

This is a PDF file of an unedited manuscript that has been accepted for publication. As a service to our customers we are providing this early version of the manuscript. The manuscript will undergo copyediting, typesetting, and review of the resulting proof before it is published in its final form. Please note that during the production process errors may be discovered which could affect the content, and all legal disclaimers that apply to the journal pertain.

Multifunctional hydrogel coatings on the surface of implantable cuff electrode for improving electrodeperipheral nerve tissue interfaces Dong Nyoung Heo d

a,b,1

, Su-Jin Song

a,1

, Han-Jun Kim c , Yi Jae Lee

, Wan-Kyu Ko a , Sang Jin Lee e , Donghyun Lee e , Sung Jin Park d ,

Lijie Grace Zhang b , Ji Yoon Kang d , Sun Hee Do c , Soo Hyun Lee d,

a

*, Il Keun Kwon a, *

Department of Dental Materials, School of Dentistry, Kyung Hee

University, Seoul 02447, Republic of Korea b

Department of Mechanical and Aerospace Engineering, The George

Washington University, DC 20052, USA c

Department of Clinical Pathology, College of Veterinary Medicine, Konkuk

University, Seoul 05029, Republic of Korea d

Center for BioMicroSystems, Korea Institute of Science and Technology,

Seoul 02455, Republic of Korea e

Department of Dental Materials, Graduate School, Kyung Hee University,

Seoul 02447, Republic of Korea

1

AUTHOR INFORMATION Corresponding Author *Tel.: +82 2 958 6440; E-mail: [email protected] (S.H. Lee).

*Tel.: +82 2 961 0771; E-mail: [email protected] (I.K. Kwon).

Author Contributions 1

D.N. Heo and S.J. Song both authors contributed equally to this work.

*

S.H. Lee and I.K. Kwon both corresponding authors contributed equally to

this work.

2

ABSTRACT Recent ly, implantable neural electrodes have been developed for recording and st imulat ion of the nervous system. However, when the electrode is implanted onto the nerve trunk, the rigid polyimide has a risk of damaging the nerve and can also cause inflammat ion due to a mechanical mismatch between the st iff polyimide and the soft biological t issue. These processes can interrupt the transmission of nerve signaling. In this paper, we have developed a nerve electrode coated with PEG hydrogel that contains poly(lact ic-co-glycolic) acid (PLGA) microspheres (MS) loaded wit h ant iinflammatory cyclosporin A (CsA). Micro-wells were introduced onto the electrode in order to increase their surface area. This allows for loading a high-dose of the drug. Addit ionally, chemically treat ing the surface wit h aminopropylmethacrylamide can improve the adhesive interface between the electrode and the hydrogel. The surface of the micro-well cuff electrode (MCE) coated with polyet hylene glycol (PEG) hydrogel and drug loaded PLGA microparticle microscopy.

(MP)

were

characterized

Addit ionally,

the

conductive

by

SEM

polymers,

and

optical

poly(3,4-

ethylenedioxyt hiophene)-poly(st yrenesulfonate) (PEDOT/PSS), were formed on the hydrogel layer for improving the nerve signal qualit y, and then

3

characterized for their electrochemical properties. The loading efficiencies and release profiles were invest igated by High Performance Liquid Chromatography (HPLC). The drug loaded electrode resulted in a sustained release of CsA. Moreover, the surface coated electrode with PEG hydrogel and CsA loaded MP showed a significant ly decreased fibrous tissue deposit ion and increased axonal densit y in animal tests. We expect that the developed

nerve electrode will

minimize the tissue damage during

regenerat ion of the nervous system.

Keywords: Cuff electrode · Cyclosporine A · Sciat ic nerve · Hydrogel coating · Drug delivery · Neural signal recording

4

1. Introduction Implantable neural prosthet ics have been developed for use in treat ing injured nervous systems. These nerve electrodes are used for interfacing with the central nervous system (CNS) or with the peripheral nervous system

(PNS).

The

interface

electrodes

should

facilitate

a

closed

interconnect ion with the nerve t issue. These should provide for select ive st imulat ion and recording from mult iple, independent, neurons of the neural system. [1-5] These nerve electrodes can be divided into two main t ypes based on their locat ion after implantat ion. There are intraneural t ypes which are located on the inside of a nerve trunk and extraneural t ypes which wrap around a nerve trunk. Amongst these several designs, extraneural electrodes such as cuff and perineural electrodes are widely invest igated because they can completely cover the nerve trunk and provide for a wide interface area. Also, these can be easily implanted wit hout nerve damage as compared wit h the intraneural t ype, which require addit ional trauma for their implantation. [6, 7] Although the current extraneural electrodes are more stable and bio co mpat ible than intraneural electrodes, they also have some limit at ions and side effects such as inflammat ion, fibrosis, and long-term stabilit y. [8, 9] When the electrodes are implanted, fibrous tissue layers are formed and 5

chronic

inflammatory

responses

occur

due

to

micro-motion

during

implantat ion and a mismatch in the mechanical properties between the nerve tissue

and

the

electrodes.

These

side

effects

reduce

nerve

signal

transduct ions and limit the long-term stabilit y and funct ionalit y of nerve electrodes. [10, 11] Therefore, there is a need to generate electrodes with enhanced ant i-inflammatory effects, reduced apoptosis, and prevent io n of fibrous tissue deposit ion in order to reduce the side effects and provide for improved chronic detection of neural signals. Many researchers have focused on surface modificat ions and coat ings of electrodes with polymers, drugs, or ant i-inflammatory biomolecules. Many different methods have been developed and tested to improve electrode bio co mpat ibilit y and prevent loss of signal qualit y over t ime. These include (1) surfaces coated with poly(ethylene glycol) (PEG)-based polymer brushes for prevent ing non-specific protein adsorption and cell adhesion by steric repulsion

[12-14],

(2)

coating

of the

neural

electrodes

with

ant i-

inflammatory drugs to prevent inflammat ion [15-17], and (3) coating the electrode with a conduct ive polymer to improve the communicat ion between the neural t issue and electrode and to obtain a high qualit y nerve signal [1820]. Particularly, conduct ive polymer-hydrogel composites have been widely 6

used for compensat ing electrical properties loss. For example, Abidian et al. reported the use of PEDOT nanotubes wit h PLGA and ant i-inflammatories (dexamethasone) for controlled release by electrical st imulat ion [21]. Hassarat i

et

al.

developed

a

conductive

polymer/hydrogel

hybrid

(PEDOT/pTS with PVA) as a cochlear implant in order to improve the electrical performance of the neural int erface [22]. For the neural probe, Kim et al. reported the use of polypyrrole/PSS electrochemically and vert ically grown through alginate hydrogel scaffolds for promoting signal transport [23]. In order to increase conductivit y, Ouyang et al. invest igated conduct ivit y of PEDOT:PSS with various addit ives such as ethylene glycol, dimet hyl sulfoxide and sorbitol [24]. However, in spite of these efforts, improvements in long-term funct ionalit y are still necessary. In view of the importance of the above studies, we have designed and prepared a funct ionalized nerve cuff electrode coated with PEG hydrogel containing Poly lact ic-co-glycol acid (PLGA) microspheres (MS) loaded with cyclosporine A (CsA). Before surface coating, the electrode substrates were altered to possess micro-wells and chemically treated in order to improve hydrogel adhesion and increase the total CsA load. Addit ionally, the

conduct ive

po lymers,

poly(3,4-ethylenedioxyt hiophene)7

poly(st yrenesulfonate) (PEDOT/PSS), were formed on the hydrogel layer for improving the nerve signal qualit y. The overall schemat ic illustration is shown in Figure 1. The surface chemical characterizat ion and distribut ion o f CsA-loaded MS on the electrode surface were determined by X-ray photoelectron spectroscopy (XPS), scanning electron microscopy (SEM), and optical microscopy. Electrochemical impedance spectroscopy (EIS) and cyclic voltammetry (CV) of funct ionalized nerve cuff electrodes were evaluated to determine the electrochemical effect iveness. Addit ionally, the effect iveness of the funct ionalized nerve cuff electrode for improvement of bio co mpat ibilit y and reduct ion of inflammatory responses and format ion o f fibrous tissue was evaluated under in vivo condit ions.

8

2. Materials and methods 2.1. Materials Polyimide (PI, VTEC PI-1338) was purchased from Richard Blaine Internat ional, Inc. (Pennsylvania, USA). Poly lact ic-co-glycolic acid (PLGA, Reso mer RG 757S) was purchased from Evonik (Essen, Germany). Cyclo sporin A (CsA), N-(3-Aminopropyl) methacrylamide hydrochloride (APMA), tribut ylamine, and poly(vinyl alcohol) (PVA) were purchased fro m Sigma-Aldrich (St. Louis, MO). Poly(ethylene glycol) diacrylate (PEGDA, average

Mw:

(Philadelphia,

1,000 PA,

g/mol)

USA).

was

The

purchased

from

photoinitiator,

Polyscience,

Inc.

1-[4-(2-hydroxyet hoxy)-

phenyl]-2-hydroxy-2-met hyl-1-propanone (Irgacure 2959) was purchased fro m Ciba Specialt y Chemicals (Basel, Switzerland). All chemical solvents were used as received without further purificat ion.

2.2. Fabrication of micro-wells on the cuff electrode PI, dissolved in n-met hyl-2-pyrrolidone, was spin-coated to a thickness o f 20 µ m on a silicon wafer. Then, the PI coated silicon wafer was cured at 90 °C for 10 min, 110 °C for 10 min, and 220 °C for 60 min in a convect ion oven (JEIO Tech Co., Ltd., Daejeon, Korea). After coat ing the negat ive 9

photoresist (DNR-L300-30, Dongjin, Seoul, Korea), Ti/Pt (50/300 nm) layers were patterned using a lift-off process. A second PI layer was spincoated to a thickness of 5 µm for insulat ion and cured in acetone. Posit ive photoresist (AZ 9260, AZ Electronic Materials, NJ, USA) was coated on the second PI layer to open the electrode site and make the micro-well. The micro-well-t ype PI patterns were etched by react ive ion etching (RIE) (Plasma Therm, St. Petersburg, FL, USA). The micro-well-t ype cuff electrode was cut by a laser dicing machine (M-2000, Exitech, Oxford, UK).

2.3. Preparation of Surface-modified PI based cuff electrode The PI based cuff electrode was immersed into 2 mL of methanol in the presence of 20 mg APMA and 0.5 mL tributylamine protected from light at room temperature for 24 h. The methacrylated PI substrates were taken out and then successively washed wit h methanol and double dist illed water (DW). Following this, they were dried in a N 2 stream. The chemical composit io n of the met hacrylated PI was determined by Xray photoelectron spectroscopy (XPS). XPS spectra were collected using a K-Alpha (Thermo Scient ific, UK) configured with a monochromat ic A1 Kα X-ray source, a spot size of 400 µ m, a flood gun to counter charging effects, 10

and an ultra-high vacuum. The chemical elements present on the samples were ident ified from the survey spectra, and expressed as atomic %.

2.4. Preparation of CsA loaded PLGA microspheres (MS) CsA loaded PLGA MS were prepared

by an oil-in-water (O/W)

emulsion/solvent evaporatio n method. Briefly, 200 mg of PLGA and 100 mg of CsA were dissolved in 2 ml of dichloromethane at room temperature. This oil phase was then emulsified in 10 ml of 1% (w/v) PVA solut ion and ho mogenized at 13,500 rpm for 2 min. The resultant emulsion was poured into 125 ml of 0.1% (w/v) PVA solut ion and stirred at 250 rpm at room temperature for 6 h to allow solvent evaporation. The MS was collected b y centrifugat ion, washed three t imes wit h DW, and freeze dried. The surface morphologies, size, and distribut ion of CsA loaded MS was characterized by scanning electron microscopy (SEM, S2300, Hitachi, Japan). Image analysis software (EyeviewAnalyzer ver. 5.0, Digiplus, Korea) was used to analyze the size and distribut ion of CsA loaded MS.

2.5. Surface coating with a hybrid hydrogel composed of PEG and CsA loaded MS 11

The PEG and CsA loaded MS composite hydrogel were coated on the surface-modified

PI

based

cuff

electrode

by

UV-induced

chemical

crosslinking. Briefly, CsA loaded MS (final concentration: 3 mg in 90 µ l) were mixed wit h 10 % (w/v) PEG and 0.1 % (w/v) photo-init iator. Subsequent ly, 90 µ l of these mixed solut ions were pipetted on the surfacemodified PI substrates and exposed to UV light (320–500 nm, 7.0 mW cm -2 , EXFO Omnicure S2000) for 60 s. A customized jig was used to expose the designated areas over the electrode to UV irradiat ion and control the thickness of a hydrogel. Addit ionally, a conductive polymer layer o f poly(3,4-ethylenedioxyt hiophene)-poly(st yrenesulfonate) (PEDOT/PSS) was polymerized electrochemically over four line electrode sites on the hybrid hydrogel-coated cuff electrode. A convent ional three-electrode configurat ion was used for the polymerizat ion of PEDOT/PSS in galvanostatic mode. The working electrode was connected to the nerve cuff electrode through an external connector. An Ag/AgCl electrode was used as a reference electrode, and Pt wire was used as a counter electrode. A current densit y of 8 µA/mm 2 was applied to each line electrode site for 300 sec. [16]

2.6. Characterization of electrochemical properties 12

Electrochemical impedance spectroscopy (EIS) and cyclic voltammetr y (CV) were used to characterize the electrochemical properties of the funct ionalized cuff electrode. An Autolab PGSTAT 302N was used to record EIS and CV of the electrode sites using a frequency response analyzer (FRA) (NOVA software, Ecochemie, Utrecht, Netherlands). A solut ion of 0.1M phosphate-buffered saline (PBS) buffer solution (pH = 7) was used as an electrolyte in all electrochemical measurements. The working electrode was connected to the funct ionalized cuff electrode, and the reference electrode was connected to the Ag/AgCl reference electrode. An AC sinusoidal signal with an amplitude of 10 mV was applied to record the EIS over a frequency range of 1 to 10 5 Hz. The CV graph was recorded after 30 repetit ive potent ial cycles to ensure that the films had reached a stable status. A scan rate of 100 mV s -1 was used, and the potent ial between the working electrode and the reference electrode was kept in the range of -0.65 to 0.8 V to avoid electrolyzing the water.

2.7. Characterization of Drug loading The release behavior of CsA from the funct ional nerve cuff electrode coated with hybrid hydrogel was determined using reverse phase high 13

performance liquid chromatography (RP-HPLC, Shimatzu, Japan). The HPLC system consisted of a solvent delivery unit (LC-20AT), an autoinjector (SIL-20A), and a Photodiode Array (PDA) detector (SPD-M20A). Briefly, the funct ional nerve cuff electrode deposited with nanofibers was soaked in 2 ml of phosphate buffered saline (PBS) solut ion and then incubated wit h cont inuous shaking (100 rpm) at 37 °C for 28 days. At predetermined periods, the incubat ion solution was collected and replaced with the same volume of fresh PBS solut ion. The collected solut ions were then filtered using a 0.5µm filter prior to inject ion. The mobile phase (1 ml/min) consisted of acetonitrile and water (65:35). The 40 µl samples were injected into an ACE 5 C18 column (250 mm x 4.6 mm, ACE®) wit h a particle size of 5 µm and monitored at 205 nm using the UV detector.

2.8. In vivo biocompatibility Animal experiments were carried out using 10 week-old, male SpragueDawley (SD) rats, weighing an average of 300–350 g (Young Bio, Seongnam, Korea). All animal experiments were approved by the Inst itut ional Animal Care and Use Committee of Konkuk Universit y (KU15082). The animals were divided into 3 groups: (1) bare cuff electrode (control), (2) cuff 14

electrode coated with PEG hydrogel and PEDOT/PSS, and (3) cuff electrode coated with PEG hydrogel, CsA loaded MSs, and PEDOT/PSS. After surface coating, all electrodes were rinsed wit h PBS and put in UV-sterilizing laminar flow hood prior to the further studies. UV sterilizat ion was performed using UV-C germicidal lamp (UV output: 19.8W, Sankyo Denki, Japan) about 2 hours of exposure to the surface of electrodes. Immediately prior

to

the

implantat ion,

each

electrode

was

rinsed

wit h

0.05%

chlorhexidine solut ion. For each group the cuff electrodes were implanted around the sciat ic nerve o f the rats. After 5 weeks of implantat ion, the sciat ic nerve of each animal was removed and stained wit h hematoxylin and eosin (H&E) to analyze nerve structure and inflammat ion. Inflammat ion was scored as 0 to 3, where 0 represents none 1- slight, 2-moderate and 3 represents severe inflammat ion. Scoring was performed by an observer blinded to sample ident it y. Fibrous t issue volume and axonal area were measured using a Leica applicat ion suite software (Leica Microsystems, Wetzlar, Germany). Cross sect ions of sciat ic nerves were stained Luxol Fast Blue-Cresyl Etch Violet (LFB-CEV) for analysis of myelin content. The total number of myelinated axons in the t issue were quant ified using ImageJ software (US Nat ional Inst itutes of Health, MD) by mult iplying the 15

est imated axonal densit y (counted myelinated axon number/measured vo lume) wit h the axonal area.

2.8. Statistical analysis All result s are expressed as means ± standard deviat ions (SD). Statist ical analysis was performed using PASW Statist ics 21 software (SPSS, Inc., Chicago, IL). A value of p < 0.05 was considered to be statist ically significant.

16

3. Results and Discussion 3.1. Characterization of the micro-well-type nerve cuff electrode To improve the interact io n between PI and biomaterials, we attempted to fabricate a micro-well-t ype PI electrode and then immobilize a UV crosslinked hydrogel onto its surface. The micro-well-t ype nerve electrodes were fabricated by the lift-off process and RIE treatment. As illustrated in Figure 1, the nerve electrode consisted of four line electrode sites and a lot of micro-wells. In this work, we used the electrode which had micro-wells with a diameter of 63.7 µ m and a width of 5.9 µ m (Figure 2A). In order to surface coat the photo-curable PEG hydrogel onto the PI electrode surface, chemical treatments wit h APMA were performed to create a methacrylated PI surface. This was confirmed by XPS analysis including survey scans and high-resolut ion N 1s spectra (Figure 2B and C). Elemental composit ions are summarized in Table 1. The bare PI electrode is mainly composed of C1s (75.57 ± 1.55 %), N1s (5.08 ± 0.36 %), and O1s (19.35 ± 1.88 %). Addit ionally, high resolut ion scans of N1s showed one main peak (Figure 2B). After chemical treatment, it was found that the N1s of methacrylated PI surface increased from 5.08 % to 7.78 % due to the presence of t he nitrogen in APMA. Also, the successful immobilization of APMA was clearly shown 17

fro m increased N/C ratio and the presence of three different peaks in high reso lut ion scans of nitrogen peaks (Figure 2C).

3.2. Surface coating with a hybrid hydrogel composed of PEG and CsA loaded MS In this study, ant i-inflammatory CsA was incorporated in PLGA MS to improve the biocompat ibilit y and reduce the inflammat ion on the implanted electrode. CsA is one of the immunosuppressive drugs that is widely used in bio medical research for treatment of inflammatory responses. [25, 26] However, CsA, like other ant i-inflammat ory drugs, induces serious side effects at high doses. Overdoses of CsA can cause transient neurological symptoms, such as ataxia, impaired locomotion, and hunched posture. [27] Thus, localized administrat ion of CsA is required to reduce the serious side effects and achieve therapeut ic effects in the range of appropriate treatment concentrations. To accomplish this mission, micro-well-t ype cuff electrodes were surfacecoated with a hybrid hydrogel composed of PEG and CsA loaded MS. CsAloaded PLGA MS was prepared using an oil-in-water emulsion method. Their morphology and size distribut ion was characterized by SEM. As 18

shown in Figure 3, the fabricated CsA loaded MS was almost spherical and well-dispersed wit h an average diameter of 13.1 ± 3.9 µ m. To increase the CsA loading amount in t he PLGA MS, a sufficient amount of CsA (100 mg in 200 mg of PLGA) was used in the fabricat ion process. The loading amount of CsA was 18.6 wt% in the PLGA MS, which was confirmed by HPLC. Aft er format ion, CsA loaded MS were mixed wit h PEG solut ion and photo-cured by UV-induced chemical crosslinking. The mixed solut ion was easily coated on the surface of methacrylated PI to form a composit e hydrogel. As shown in the opt ical and SEM images (Figure 4) CsA loaded MS was well distributed in the hydrogel layer and the composite hydrogel was evenly covered over the ent ire treated side of the electrode. The thickness of the result ing co mposite hydrogel layer was 208 ± 11 µ m. Also, the PLGA MS was well-dispersed inside the photo-cured PEG hydrogel layer. The composite hydrogel layer had improved adhesion to the electrode due to the micro-well format ion and surface methacrylat ion.

3.3. Electrochemical properties of surface-coated cuff electrodes In this study, PEG hydrogel embedded wit h CsA loaded MS was deposited on the micro-well-t ype cuff electrode for sustained delivery of ant i19

inflammatory agents. Alt hough the PEG hydrogel layer provides for mechanical dampening between neural t issue and cuff electrode, it can lead to decreased electrical signal for nerve signal recording. When the thickness of the hydrogel coating was increased, the impedances of hydrogel-coated electrodes were slight ly increased (Figure S1, Supporting Informat ion). To combat this disadvantage, PEDOT/PSS was polymerized in the PEG hydrogel layer to improve its electrochemical properties. Figure 5 shows the comparison of EIS and CV wit h the different types of surface-treated electrode. The five different types of samples were prepared as shown below: (1) Control : bare cuff electrode (2) Hydrogel : coat ing wit h PEG hydrogel (3) Hydrogel + PEDOT : PEDOT deposit ion inside the PEG hydrogel (4) Hydrogel + MS : coating wit h PEG hydrogel and CsA loaded MS (5) Hydrogel + MS + PEDOT : PEDOT deposit ion inside the hybrid hydrogel layer wit h PEG and CsA loaded MS EIS were recorded wit h amplit ude of 10 mV in the frequency range of 1 to 10 5 Hz. The signals were compared wit h the value of each other at 1 kHz of impedance. This is in concordance wit h the unique frequency of neuronal act ion potentials. [28] As shown in Figure 5A, the bare PI electrode had a 20

high impedance of 851.4 ± 27.5 Ω as measured at 1 kHz. The impedances o f hydrogel-coated groups were slight ly increased to 904.7 ± 12.1 Ω (Hydrogel) and 915.3 ± 127.9 Ω (Hydrogel + MS) by the int erference of the nonconduct ive layer. In order to achieve reduced impedance value, the PEDOT was deposited inside the hydrogel layer. Consequent ly, the measured impedances at 1 kHz were decreased to 586.8 ± 33.8 Ω (Hydrogel + PEDOT) and 580.2 ± 40.1 Ω (Hydrogel + MS + PEDOT), respectively. The high impedance value of bare PI electrode could be reduced after PEDOT deposit ion. This was not affected by the addit ion of CsA loaded MSs inside the hydrogel layer. Fo llowing surface coating, the 5 different kinds of electrodes (Control, Hydrogel, Hydrogel + PEDOT, Hydrogel + MS, and Hydrogel + MS + PEDOT) were scanned simultaneously by CV. The CV values were used to calculate the charge delivery capacit ies (CDC, charge capacit y per unit area). Figure 5B displays the CV curves from -0.65 to 0.8 V with a scanning rate of 100 mV/s. The calculated CDCs of Control, Hydrogel, Hydrogel + PEDOT, Hydrogel + MS, and Hydrogel + MS + PEDOT were 1.87 ± 0.14, 1.83 ± 0.21, 2.66 ± 0.34, 1.99 ± 0.28, 2.67 ± 0.37 µC/mm2 , respect ively. The electrodes containing PEDOT exhibited substant ially larger current densit y 21

as co mpared wit h control. This indicates that PEDOT deposit ion has potent ial applicat ion in charge storage. Sekine et al. reported the micro-patterning of PEDOT on a hydrogel to develop a flexible recording electrode. [29] Conductive PEDOT was deposited on a hydrogel-based electrode by electro-polymerizat ion followed by electrochemical-actuation-assisted peeling. The fabricated hydrogelbased electrode was able to supply electrical st imulat ion through the PEDOT. Also, Guo et al. reported the development of a flexible microelectrode array for in vitro and in vivo bio-signal recording. [30] The electrode arrays were composed of poly(dimethylsiloxane) (PDMS) as a plast ic carrier and track insulat ion material. These used PEDOT as an organic conductor. Their electrical

and

recording

funct ionalit ies

were

confirmed

by in

vivo

epicortical and epidural recording. Their results showed that the qualit y o f electrochemical properties was improved by introduction of PEDOT. Likewise, our results showed that deposit ing the PEDOT inside the hydrogel layer causes dramat ic changes in electrochemical performance and provides higher qualit y signals.

3.4. Release profile of CsA from the surface-coated cuff electrode 22

Figure 6 shows the release profiles of CsA from surface-coated cuff electrodes. For loading of CsA, CsA loaded PLGA polymer was incorporated in the hydrogel as MS or as nanofibers. CsA-loaded PLGA nanofibers were deposited on the nerve cuff electrode as previously described. [16] The total load amounts of CsA were approximately 458.4 µ g/electrode (MS) and 48.2 µ g/electrode (nanofiber), respect ively. As shown in Figure 6, the cumulat ive release profiles of CsA from the surface-coated cuff electrode exhibited two kinds of release pattern wit h an init ial burst release and a late sustained release. During an init ial release stage, the burst release of CsA was 55.9 µ g fro m MS and 16.3 µ g from nanofibers, respect ively. Such a release pattern happened by penetration of water into the polymer coated CsA, which diffused out init ially through surface erosion. Subsequent ly, the sustained release pattern happened by CsA-loaded polymer degradat ion and bulk erosion. [31] As a result, the total amounts of released CsA after 960 h, were 106.8 µ g from MS and 37.9 ug from nanofibers. The cumulat ive release amounts of CsA were dependent on its total loading amount on the electrode, diffusion and degradat ion o f the PLGA polymers. [32] Also, these release patterns have been frequent ly reported. [33, 34, 35]

23

3.5. In vivo Biological Evaluation In our previous research, we developed a funct ional nerve cuff electrode coated with dexamet hasone (DEX)-loaded PLGA nanofiber and PEG hydrogel. We confirmed their funct ion for stable recording of sciat ic nerve signals and controlled drug release. [16]DEX-loaded PLGA nanofibers were deposited on the nerve cuff electrode using an electrospinning method. This showed a long-term period of sustained drug release. However, our previous study only confirmed the funct ionalit y in terms of effect ive neural signal recording. Alt hough the cuff electrode coated with DEX-loaded PLGA nano fiber had a well-controlled drug release rate, the inflammat ion suppression effects under in vivo condit ions were not confirmed. Therefore, in vivo studies are necessary to ensure successful reduct ion of inflammat ion. The in vivo bioact ivit y of CsA was investigated as compared with DEX. Both ant i-inflammatory drugs, CsA and DEX, were deposited on the nerve cuff electrode using PLGA nanofibers. These were implanted into the rat sciat ic nerves (Figure S1 and S2, Supporting Informat ion). Based on these supplementary results, we found that both CsA and DEX are effect ive to prevent fibrous tissue infiltrat ion and apoptotic cell death. We also found 24

that CsA is more effect ive t han DEX and that an addit ional loading amount of CsA should be necessary for full therapeutic effect. Nafe el al. found that the effect ive drug concentrations to provide for 50% inhibit ion of T-cell proliferat ion is 1.4 ± 0.7 mg/mL for DEX and 15.8 ± 2.3 ng/mL for CsA, respect ively. [36] Therefore, the CsA loaded PLGA nanofiber group had a lower inflammat ion score, fibrous thickness, and apoptotic neuronal cell count

as

compared

wit h the

DEX

loaded

PLGA nanofiber

group.

Furthermore, an addit ional inject ion of CsA (0.1 mg/kg) along wit h the CsA loaded PLGA nanofiber group showed enhanced ant i-inflammatory, ant ifibrotic, and ant i-apoptotic effects as compared to other tested groups. These results indicate that the total loading amount of CsA in the PLGA nano fiber was not enough to fully prevent the inflammatory response at the interfaces of the nerve cuff electrode. For this reason, we developed CsA loaded PLGA MS to increase the total loading amount of CsA in order to enhance the local biocompat ibilit y as compared with the CsA loaded PLGA nano fiber. As shown in t he kinet ic release profile, the total amount of released CsA from MS was more than 2.8 times that from the nanofiber (Figure 6). After fabricat ion of the CsA loaded MS, they were mixed wit h PEG solut ion and deposited on the micro-well-t ype cuff electrode. These 25

surface coated electrodes were used to confirm the effects of inflammat ion suppression under in vivo condit ion. To analyze the effect iveness of funct ionalized nerve cuff electrode, three different kinds of electrodes (Control, Hydrogel + PEDOT, and Hydrogel + MS + PEDOT) were implanted to the rat sciat ic nerves. After 5 weeks, histo-morphological changes of the nerve tissues were examined. These changes included thickness of epineural fibrous tissue infiltrat ion and nerve integrit y as determined by H & E, and Luxol fast blue / Cresyl Etch Violet staining, respect ively (Figure 7 and 8). In general, damage to the nerve tissue in the implantat ion site may occur due to stress during the implantat ion procedure, physical movement, and a mismatch in the mechanical properties between the st iff electrode and soft tissue. [37, 38] As a result of these mechanical stresses, fibrous t issue deposit ion occurs between the surface of the electrode and the epineurium. However, after surface coat ing wit h PEG hydrogel and CsA loaded MS, the relat ive fibrous tissue volume (fibrous tissue volume / total tissue volume (%)) was significant ly decreased as compared wit h control. These result s may indicate that applicat ion of a CsA loaded MS/PEG hydrogel can reduce the mechanical stresses at the interface between the cuff electrode and the 26

bio logical t issue. Addit ionally, surface coating the hydrogel + PEDOT group and hydrogel + MS + PEDOT group showed an increased axonal area, densit y, and total axo n number as compared to control. The axonal densit y of the hydrogel + MS + PEDOT group was higher than the other groups. This is due to a reduct ion of the mechanical mismatch by the soft PEG hydrogel layer and ant i-inflammatory effects of CsA. These results indicate that a CsA loaded MS coating yielded increased biocompat ibilit y through decreased foreign body responses and axonal loss as compared with both the control and the hydrogel + PEDOT group.

4. Conclusion In conclusion, we have successfully designed and developed a funct ional cuff electrode for prevent ing inflammat ions around the implanted neural electrode surface and improving their electrochemical properties. This was done by surface-coating a CsA loaded MS/PEG hydrogel onto the electrode fo llowed by electro-polymerizat ion to deposit a conduct ive PEDOT inside of the hydrogel layer. Before surface coating, the physicochemical adhesio n properties between the electrode and the hydrogel layer were increased 27

through the format ion of micro-wells electrochemical

experiments

showed

and

that

chemical treat ment.

the

deposit ion

of

The

PEDOT

improved the electrochemical properties through decreased impedance and increased CDC. The release test of CsA indicated that the surface coated electrode can deliver a sufficient amount of CsA to reduce inflammat ion for a long t ime. Moreover, animal tests indicated that the CsA loaded MP/PEG surface coated electrode had a significant decrease in fibrous t issue deposit ion and increase in axonal densit y as compared to control. Our findings suggest that surface coating a soft-hydrogel along wit h an ant iinflammatory drug loaded MS can be a useful strategy for improving the lo ng-term biocompat ibilit y of electrodes.

Acknowledgement This research was supported by the Public Welfare & Safet y research program through the Nat ional Research Foundat ion of Korea (NRF) funded by the Ministry of Educat ion, Science and Technology (NRF-2010-0019346), (NRF-2012-0008610), (NRF-2012R1A5A2051388).

References 28

[1] Navarro X, Krueger TB, Lago N, Micera S, Stieglit z T, Dario P. A crit ical review of interfaces wit h the peripheral nervous system for the control of neuroprostheses and hybrid bionic systems. J Peripher Nerv Syst 2005;10:229-258. [2] Grill WM, Norman SE, Bellamkonda RV. Implanted neural interfaces: biochallenges and engineered solut ions. Annu Rev Biomed Eng 2009;11:1-24. [3] Stieglit z T, Schuetter M, Koch KP. Implantable biomedical microsystems for neural prostheses. IEEE Eng Med Biol Mag 2005;24:58-65. [4] Boretius T, Badia J, Pascual-Font A, Schuettler M, Navarro X, Yoshida K, Stieglitz T. A transverse intrafascicular mult ichannel electrode (TIME) to interface wit h the peripheral nerve. Biosens Bioelectron 2010;26:62-69. [5] Cheung KC. Implantable microscale neural interfaces. Biomed Microdevices 2007;9:923-938. [6] Chu J, Song K, Han S, Lee SH, Kim J, Kang JY, Hwang D, et al. Improvement of signal-to-interference rat io and signal-to-noise rat io in nerve cuff electrode systems. Physiol Meas 2012;33:943. [7] Micera S, Navarro X. Bidirect ional interfaces wit h the peripheral nervous system. Int Rev Neurobiol 2009;86:23-38. [8] Heiduschka P, Thanos S. Implantable bioelectronic interfaces for lost nerve funct ions. Prog Neurobiol 1998;55:433-461. 29

[9] Li Y, Peng C, Chen L, Lin W, Chu C, Chen J. Applicat ion of implantable wireless bio microsystem for monitoring nerve impedance of rat after sciat ic nerve injury. IEEE Trans Neural Syst Rehabil Eng 2013;21:121-128. [10]

Thil M, Colin IM, Delbeke J. Time course of t issue remodelling and

electrophysiology in the rat sciat ic nerve after spiral cuff electrode implantat ion. J Neuroimmunol 2007;185:103-114. [11]

Cheung KC, Renaud P, Tanila H, Djupsund K. Flexible polyimide

microelectrode array for in vivo recordings and current source densit y analysis. Biosens Bioelectron 2007;22:1783-1790. [12]

Ma H, Li D, Sheng X, Zhao B, Chilkoti A. Protein-resistant polymer

coatings on silicon oxide by surface-init iated atom transfer radical polymerizat ion. Langmuir 2006;22:3751-3756. [13]

Heo DN, Yang DH, Lee JB, Bae MS, Park HN, Kwon IK. Cell

fouling resistance of PEG-grafted polyimide film for neural implant applicat ions. Proc of SPIE 2012;84099:84091F. [14]

Bridges A, García A. Ant i-inflammatory polymeric coat ings for

implantable biomaterials and devices. J Diabetes Sci Technol 2008;2:984-994. [15]

Astolfi L, Guaran V, Marchetti N, Olivetto E, Simoni E, Cavazzini

A, Jolly C, et al. Cochlear implants and drug delivery: In vitro evaluat ion of dexamet hasone release. J Biomed Mater Res B Appl Biomater 2014;102:267-273.

30

[16]

Park SJ, Lee YJ, Heo DN, Kwon IK, Yun K, Kang JY, Lee SH.

Funct ional nerve cuff electrode with controllable ant i-inflammatory drug loading and release by biodegradable nanofibers and hydrogel deposit ion. Sensors Actuators B: Chem 2015;215:133-141. [17]

Abidian MR, Martin DC. Mult ifunct ional nanobiomaterials for neural

interfaces. Adv Func Mater 2009;19:573-585. [18]

Green RA, Lovell NH, Wallace GG, Poole-Warren LA. Conduct ing

polymers for neural interfaces: challenges in developing an effect ive lo ng-term implant. Biomaterials 2008;29:3393-3399. [19]

Khodagholy D, Doublet T, Gurfinkel M, Quilichini P, Ismailova E,

Leleux P, Herve T, et al. Highly conformable conduct ing polymer electrodes for in vivo recordings. Adv Mater 2011;23:268-272. [20]

Guo L, Ma M, Zhang N, Langer R, Anderson DG. Stretchable

Polymeric Mult ielectrode Array for Conformal Neural Interfacing. Adv Mater 2014;26:1427-1433. [21]

Abidian MR, Kim DH, Mart in DC, Conduct ing-polymer nanotube for

controlled drug release. Adv Mater 2006;18:405-409. [22]

Hassarat i RT, Dueck, WF, Tasche C, Carter PM, Poole-warren LA,

Green RA, Improving cochlear implant properties through conduct ive hydrogel coat ings. IEEE Trans Neural Syst Rehabil Eng 2014;22:411418. [23]

Kim DH, Abidian MR, Mart in DC, Conduct ing polymers grown in

hydrogel scaffolds coated on neural prosthetic devices. J Biomed Mater Res A 2004;71:577-585. 31

[24]

Quyang L, Musumeci C, Jafari MJ, Ederth T, Inganas O, Imaging the

phase separat ion between PEDOT and polyelectrolytes during processing of highly conduct ive PEDOT:PSS films. ACS Appl Mater Interfaces 2015:7:19764-19773. [25]

Holan V, Chudickova M, Trosan P, Svobodova E, Krulova M,

Kubinova S, Sykova E, et al. Cyclosporine A-loaded and stem cellseeded electrospun nano fibers for cell-based therapy and local immunosuppression. J Controlled Release 2011;156:406-412. [26]

Canudas AM, Jordà EG, Verdaguer E, Jiménez A, Sureda FX,

Rimbau V, Camins A, et al. Cyclosporin A enhances colchicine‐induced apoptosis in rat cerebellar granule neurons. Br J Pharmacol 2004;141:661-669. [27]

Cruz F, Wolf A. Effects of the novel cyclosporine derivat ive PSC833

on glucose metabolism in rat primary cultures of neuronal and glial cells. Biochem Pharmaco l 2001;62:129-139. [28]

Maher M, Pine J, Wright J, Tai Y. The neurochip: a new

mult ielectrode device for stimulat ing and recording from cultured neurons. J Neurosci Methods 1999;87:45-56. [29]

Sekine S, Ido Y, Miyake T, Nagamine K, Nishizawa M. Conduct ing

polymer electrodes printed on hydrogel. J Am Chem Soc 2010;132:13174-13175. [30]

Blau A, Murr A, Wolff S, Sernagor E, Medini P, Iurilli G, Ziegler C,

et al. Flexible, all-polymer microelectrode arrays for the capture of cardiac and neuronal signals. Biomaterials 2011;32:1778-1786.

32

[31]

Heo DN, Lee JB, Bae MS, Hwang YS, Kwon KH, Kwon IK.

Development of nanofiber coated indomethacin-elut ing stent for tracheal regenerat ion. J Nanosci Nanotechnol 2011;11:5711-5716 [32]

Heo DN, Yang DH, Lee JB, Bae MS, Kim JH, Moon SH, Chun HJ, et

al. Burn-wound healing effect of gelat in/polyurethane nanofiber scaffold containing silver-sulfadiazine. J Biomed Nanotech 2013;9:511-515. [33]

Bae MS, Yang DH, Lee JB, Heo DN, Kwon YD, Youn IC, et al.

Photo-cured hyaluronic acid-based hydrogels containing simvastat in as a bone tissue regenerat ion scaffold. Biomaterials 2011;32:8161-8171. [34]

Chew SY, Wen J, Yim EK, Leong KW. Sustained release of proteins

from electrospun biodegradable fibers. Biomacromolecules 2005;6:20172024. [35]

Freiberg S1, Zhu XX. Polymer microspheres for controlled drug

release. Int J Pharm 2004;282:1-18. [36]

Nafe LA, Dodam JR, Reinero CR. In-vitro immunosuppression of

canine T-lymphocyte-specific proliferat ion with dexamethasone, cyclosporine, and the act ive metabolites of azathioprine and leflunomide in a flow-cytometric assay. Can J Vet Res 2014;78:168-175. [37]

Moshayedi P, Ng G, Kwok JC, Yeo GS, Bryant CE, Fawcett JW,

Franze K, et al. The relat ionship between glial cell mechanosensit ivit y and foreign body react ions in the central nervous system. Biomaterials 2014;35:3919-3925.

33

[38]

Minev IR, Musienko P, Hirsch A, Barraud Q, Wenger N, Moraud

EM, Gandar J, et al. Bio materials. Electronic dura mater for long-term mult imodal neural interfaces. Science 2015;347:159-163.

34

Fig. 1 CsA loaded PLGA MS

PEG hydrogel

Microsphere & Hydrogel Coating Micro-well cuff electrode

PEDOT/PSS Deposition

Implantation

35

Fig. 2

36

Fig. 3

37

Fig. 4

38

Fig. 5

39

Fig. 6

CsA loaded PLGA microparticle CsA loaded PLGA nanofiber

Released CsA (ug) (µg) Released CsA

120

100

80

60

40

20

0 0

200

400

600

Time (h)

40

800

1000

Fig. 7

41

Fig. 8

42

Figure Captions

Fig. 1. Schematic diagram of the fabrication process for the functionalized nerve cuff

electrode. Microwell-t ype cuff el ectrode is coated with CsA-loaded PLGA, MS, and PEG

hydrogel mixture. After PEDOT deposition inside of hydrogel layer, the surface-coated

electrode is implanted into the rat sciatic nerve.

Fig. 2. (A) Surface morphology of cuff electrode by SEM. Micro-wells on the surface

had a diameter of 63.7 µ m and a width of 5.9 µ m. (B) (C) Wide-scanned XPS spectra

(left) and nitrogen peak (right) of (B) pristine and (C) methacrylated PI electrode. The

spectra were measured from 0 to 1200 eV. Experiments performed in triplicate (n = 3).

Fig. 3. (A) SEM image of CsA-l oaded PLGA MS. Their shape is almost spherical and

uniform. (B) Size distribution of CsA-l oaded PLGA MS as measured by SEM images.

About 200 of microspheres were counted, and had an average diameter of 13.1 ± 3.9 µ m

(mean ± SD).

Fig. 4. (A, B) Optical image and (C, D) cross-sect ion SEM images of the surface-coated

43

cuff electrode with CsA loaded PLGA, MS, and PEG hydrogel. CsA loaded MS was well

distributed in the hydrogel la yer. The PEG hydrogel embedded with CsA loaded MS wa s

evenl y covered across the surface of the microwell-t ype cuff electrode.

Fig. 5. Comparison of (A) EIS and (B) CV of the control, PEG hydrogel, PEG + PEDOT,

PEG + microsphere, and PEG + microsphere + PEDOT groups. An AC sinusoidal signal

with an amplitude of 10 mV was applied to measure the impedance and the CV was

measured at a scan rate of 100 mV/s. Results shown as average for six experiments (n =

6).

Fig. 6. Release profile of the CsA loaded PLGA microparticle and CsA loaded PLGA

nanofiber in PBS at 37 ℃. The release test was carried out for 40 days. Results are

shown as mean ± SD for six replicates (n = 6).

Fig. 7. Fibrous tissue deposits after implantation with different surface coated el ectrodes:

(A) Fibrous tissue area of the nerve tissue as confirmed by H&E staining and (B) their

quantification. At 5 weeks, the nerve tissue after electrode implantation was mainl y

44

composed of dense fibrous tissue and inflammatory cells. Fibrosis deposit was decreased

by PEG hydrogel and CsA loaded with MS. Scale bar = 200 µm. *p < 0.05 indicates

significant difference as compared with the control.

Fig. 8. Histol ogical analysis of sciatic nerves after 5 weeks of implantation with Control,

Hydrogel + PEDOT (HP), and Hydrogel + MS + PEDOT (HMP) coated electrodes. (A)

CsA loaded with MP coating group maintains a more highly packed nerve fi ber and

myelin sheathe as compared to control and hydrogel + PEDOT group. (B) The CsA

loaded with MP group had significantly increased axonal area, axonal density, and total

axon numbers. Scale bar = 50 µm. *p < 0.05 and **p < 0.01 indicates significant

difference as compared with control.

45

Table 1

Table Captions

Table 1. Chemical compositions of PI electrode surface following treatments. Results are

mean ± SD of triplicate experiments.

46

Graphical abstract

CsA loaded PLGA MS

PEG hydrogel

Microsphere & Hydrogel Coating Micro-well cuff electrode

PEDOT/PSS Deposition

Implantation

47