Soft hydrogel zwitterionic coatings minimize fibroblast and

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Soft hydrogel zwitterionic coatings minimize fibroblast and macrophage adhesion on polyimide substrates Dusana Trelova, Alice Rita Salgarella, Leonardo Ricotti, Guido Giudetti, Annarita Cutrone, Petra Sramkova, Anna Zahoranova, Dusan Chorvat, Daniel Hasko, Claudio Canale, Silvestro Micera, Juraj Kronek, Arianna Menciassi, and Igor Lacik Langmuir, Just Accepted Manuscript • DOI: 10.1021/acs.langmuir.8b00765 • Publication Date (Web): 24 May 2018 Downloaded from http://pubs.acs.org on May 24, 2018

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Langmuir

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Soft hydrogel zwitterionic coatings minimize fibroblast and macrophage

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adhesion on polyimide substrates

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Dušana Treľová,a,# Alice Rita Salgarella,b,# Leonardo Ricotti,b,§ Guido Giudetti,b,§ Annarita

5

Cutrone,b,c Petra Šrámková,a Anna Zahoranová,a Dušan Chorvát Jr.,d Daniel Haško,d Claudio

6

Canale,e,f Silvestro Micera,b,g,* Juraj Kronek,a,*Arianna Menciassi,b,* Igor Lacíka,*

7 8 9 10 11

12

13

14

a

Department for Biomaterials Research, Polymer Institute of the Slovak Academy of Sciences,

Dúbravská cesta 9, 845 41 Bratislava, Slovakia b

The BioRobotics Institute, Scuola Superiore Sant’Anna, Viale R. Piaggio 34, 56025

Pontedera (PI), Italy c

SMANIA srl, via G. Volpe 12, 56121 Pisa, Italy

d

e

International Laser Centre, Ilkovičova 3, Bratislava 841 04, Slovak Republic

Department of Physics, University of Genova, Via dodecaneso 33, 16133 Genova, Italy

15

f

16

Genova (Italy)

17

g

18

and Institute of Bioengineering, Ecole Polytechnique Federale de Lausanne, Lausanne, CH

19

#,§

equally contributing authors

20

*

corresponding authors: [email protected], [email protected],

21

Department of Nanophysics, Istituto Italiano di Tecnologia (IIT), Via Morego 30, 16163

Bertarelli Foundation Chair in Translational Neuroengineering, Center for Neuroprosthetics

[email protected], [email protected]

22

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Abstract

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Minimizing the foreign body reaction to polyimide-based implanted devices has a pivotal role

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for several biomedical applications. In this work we propose materials exhibiting non-

26

biofouling properties and a Young's modulus reflecting the one of soft human tissues. We

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describe the synthesis, characterization and in vitro validation of poly(carboxybetaine)

28

hydrogel coatings covalently attached to polyimide substrates via a photolabile 4-azidophenyl

29

group, incorporated in poly(carboxybetaine) chains at two concentrations of 1.6 and 3.1

30

mol.%. The presence of coatings was confirmed by attenuated total reflectance Fourier-

31

transform infrared spectroscopy. White light interferometry was used to evaluate coating

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continuity and thickness (resulting between 3 and 6 µm in dry conditions). Confocal laser

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scanning microscopy allowed to quantify the thickness of the swollen hydrogel coatings that

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ranged between 13 and 32 µm. The different hydrogel formulations resulted in stiffness

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values ranging from 2 to 19 kPa, and led to different fibroblasts and macrophages responses,

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in vitro. Both cell types showed a minimum adhesion on the softest hydrogel type. In

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addition, both the overall macrophage activation and cytotoxicity were observed to be

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negligible for all the tested material formulations. These results are a promising starting point

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towards future advanced implantable systems. In particular, such technology paves the way to

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novel neural interfaces able to minimize the fibrotic reaction, once implanted in vivo, and to

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maximize their long-term stability and functionality.

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Keywords:

43

soft hydrogel coatings, poly(carboxybetaine), 4-azidophenyl, non-biofouling materials,

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macrophage adhesion, fibroblast adhesion, polyimide, neural interfaces.


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INTRODUCTION

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Implantable devices are gaining an increasing relevance in the biomedical engineering field.

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This is due to their high potential to restore the body functions after impairments and,

49

generally, to improve the patients’ quality of life.1 To address these clinical goals, it is

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desirable to achieve a long-term functionality of implanted systems keeping them stable

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within the body environment and avoiding undesired biological processes.

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The implant long-term functionality is particularly important in the case of neural interfaces.

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These systems are active implanted medical devices aimed at restoring a connection with the

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central and peripheral nervous system after pathological conditions, such as neuropathies,

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multiple sclerosis, amyotrophic lateral sclerosis or spinal cord injuries and limb amputations.2-

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3

Exciting results have been recently achieved, which demonstrated the potential of this

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implantable technology.4-6 Among the different options, polyimide (PI) has been identified as

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a promising candidate material for the development of advanced neural interfaces. Indeed, it

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is a biocompatible and flexible polymer featured by proper dielectric properties, thermal

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stability, resistance to solvents and strong adhesion to metals and metal oxides. Moreover, it

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can be easily patterned with lithographic and dry etching techniques and produced in the form

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of thin layers by spinning deposition.7-9 Different PI-based interface designs have been

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developed in the last years for selectively interfacing with the central nervous system for

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intracortical neural activity recordings,10-11 cortical surface field potential recordings12 as well

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as for intraneural interfacing with peripheral nerves.13-16

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Currently, a full medical and commercial exploitation of neural interfaces is hampered by a

69

limited long-term implant stability. Intraneural implants are highly demanding since they


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provide an intimate contact with axons, and trigger foreign body response and molecular

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cascades related to chronic inflammation. This leads to the formation of a scar-like fibrotic

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tissue that engulfs and electrically insulates the neural interface, thus altering the recording of

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nerve action potentials and stimulation thresholds.3, 14, 17 This process starts already with the

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injury produced by the invasive implantation procedure, which results in an acute

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inflammation and leucocytes accumulation at the site of injury. Afterwards, the inflammation

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becomes chronic, granulation tissue is produced by fibroblasts, and the foreign body reaction

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proceeds with the accumulation of giant cells and fibrosis.18-19

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Several strategies have been proposed to overcome the foreign body and inflammatory

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reactions stimulated by implanted neural electrodes. In some cases the material surface was

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functionalized with the aim of improving tissue integration within the implant. To this

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purpose, adhesion molecules,20 specific peptides,21-23 and whole engineered proteins24 have

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been proposed. Another strategy was based on the physical modification of the implant

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surface, e.g. by providing it with proper patterns.25-27 A further option was to suppress adverse

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local reactions by means of polymeric coatings exhibiting advanced features, such as

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controlled release of bioactive molecules and/or non-biofouling surface properties.28-31 The

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undesired host response to implants is due not only to their surface chemical properties, but it

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is also determined by the mechanical mismatch between the implant and the surrounding

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tissue.7, 32-33 The use of PI-based interfaces reduces this mismatch compared to silicon- and

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glass-based interfaces. Nevertheless, the PI Young's modulus of a few GPa is significantly

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higher than the Young's modulus of neural tissue, which ranges between 0.1 and 10 kPa.34

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Moreover, it is known that the substrate stiffness influences the behavior of a wide variety of

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cells, including macrophages, which are directly and primarily involved in the body reaction

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to implants. Indeed, Blakney et al. demonstrated that macrophage activation is reduced on


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softer poly(ethylene glycol)-based hydrogels exhibiting the elastic modulus in the range from

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130 to 840 kPa.35 Softer hydrogels also induced a lower expression level of genes involved in

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the production of inflammatory cytokines and showed a less severe foreign body reaction in

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vivo. Hydrogel-based coatings can thus be considered a promising solution to minimize the

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mechanical mismatch between implanted devices and tissues. These hydrogel coatings can be

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produced in a wide range of Young's moduli (even a few kPa)36 and made of polymers

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exhibiting non-biofouling properties.

102 103

Most of the polymeric layers proposed so far as neural interface coatings were based on

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poly(ethylene glycol) (PEG) and PEG-based copolymers utilizing the PEG non-biofouling

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properties.29-30,

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Spencer et al. recently reported PEG-based hydrogels with controllable elastic moduli as

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scarring-reducing coatings on neural probes.39 However, PEG is known to produce undesired

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effects in vivo, e.g. oxidative damage in biological fluids, bioaccumulation in cell lysosomes

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and generation of anti-PEG antibodies that may lead to a severe immune response.40-42

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Therefore, the exploration of other types of non-biofouling polymers for a functional

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protective coating of neural electrodes represents an ongoing challenge.

37-38

However, the role of stiffness was not investigated in these studies.

112 113

Zwitterionic

polymers,

including

poly(carboxybetaines),

114

poly(phosphobetaines), are non-toxic materials with a high hemocompatibility and ultralow

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non-specific protein and cell adhesion.43-45 The mechanism behind this behavior was ascribed

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to the formation of a hydration layer,46 which prevents hydrophobic interactions of proteins

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and lipidic membranes of cells with the polymer surface. It was recently demonstrated that

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surfaces decorated with zwitterionic molecules bind water molecules more strongly than PEG

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due to electrostatically induced hydration.47 Strategies for surface attachment of zwitterion-


poly(sulfobetaines)

and

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based compounds can be divided into “grafting from” and “grafting to” methods. In the first

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case, methacrylates and methacrylamides containing zwitterionic moieties in a side chain are

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polymerized via surface-initiated controlled radical polymerizations.48 The “grafting to”

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method typically involves chemically induced attachment of zwitterionic polymers,

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containing suitable chemical moieties, to the surface of other materials.49

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Recently, zwitterionic polymers bearing photoactivated groups were used for the preparation

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of covalently attached coatings via a photochemically controlled process. The overall benefits

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of a photochemical surface modification are (i) formation of a stable bond with hydrocarbon

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groups on various polymer substrates, (ii) simple manufacturing in dry state, and (iii)

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possibility

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phosphorylcholine and benzophenone units were used for modification of a commercial cyclic

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polyolefine.50 Similarly, photoreactive poly(sulfobetaines) and poly(carboxybetaines) bearing

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photolabile arylazide groups were applied for modification of various polymeric substrates.49,

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51-53

of

micropatterning.

Photoreactive

copolymers

bearing

zwitterionic

However, to the best of our knowledge, zwitterionic polymers have never been

135

photochemically coupled with PI substrates in view of a future use in the domain of neural

136

interfaces.

137 138

In this paper, a zwitterionic hydrogel coating of PI substrates is proposed. The hydrogel is

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based on poly(carboxybetaine methacrylamide) selected due to its favorable properties such

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as biocompatibility, non-biofouling and hydration.54-55 The coating is obtained through a

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photoactivated process enabled by a photolabile azidophenyl group. We aimed at exploiting

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the non-biofouling properties of zwitterionic hydrogels and achieving low Young’s modulus

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values in the range of the neural tissue. The hydrogel layers were characterized to verify the

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presence of hydrogel attached to the PI surface, hydrogel thickness and Young’s modulus. In


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vitro tests with fibroblasts and macrophages were carried out to evaluate the adhesion of these

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cells on the hydrogel-modified PI surfaces and to correlate these results with hydrogel

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stiffness. This work thus contributes to create a strategy for neural interface coating and

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envisions future applications of these materials in vivo.

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EXPERIMENTAL PART

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Materials

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N-[3-(dimethylamino)propyl]methacrylamide,

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hydrochloride, methacryloyl chloride, 2,2’-azobis(2-methylpropionamidine) dihydrochloride

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(AIBA), amberlite IRA 400 chloride, and Pluronic F-127 were purchased from Sigma Aldrich

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(Steinheim, Germany) and used as received. Sodium hydroxide, sodium carbonate, and

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dioxane were purchased from Lachema (Brno, Czech Republic). Acetone, ethanol, and

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dichloromethane were purchased from CentralChem (Slovakia) and distilled prior to use.

ethyl

bromoacetate,

4-azidoaniline

159 160

Preparation of 2-{dimethyl[3-(2-methylprop-2-enamido)propyl]ammonio}acetate (M1)

161

Monomer M1 was prepared by adapting the procedure described by Abraham and Unsworth56

162

(Scheme S1). N-(3-(dimethylamino)propyl)methacrylamide (5 g, 29 mmol) was diluted in 20

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mL of acetone. Then ethyl bromoacetate (3.1 mL, 28 mmol) was added dropwise into the

164

reaction mixture under an argon atmosphere. The reaction mixture was stirred overnight at

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room temperature. The white precipitate was filtered off, washed with acetone and dried

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under reduced pressure. The resulting white powder was dissolved in distilled water and

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applied onto a column filled with basic ionic exchange resin (Amberlite IRA 400) activated

168

for 30 min with 1 M aq. NaOH. After ion exchange, the solvent was evaporated and the


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product was dried under reduced pressure. The resulting product was obtained as a colorless

170

viscous compound with a yield of 7 g (~ 100%). For 1H NMR spectrum of M1, see Figure S1.

171

1

H NMR (400 MHz, D2O, δ, ppm): 1.79 (s, CH3-C=CH2, 3H), 1.88 (m, -CH2-, 2H), 3.07 (s,

172

(CH3)2-N+, 6H), 3.22 (t, NH-CH2-, 2H), 3.43-3.47 (m, -CH2-N+, 2H), 3.72 (s, N+-CH2-COO-,

173

2H), 5.33 (s, CH2=, 1H), 5.59 (s, CH2=, 1H).

174 175

Preparation of N-(4-azidophenyl)-2-methylprop-2-enamide (M2)

176

Monomer M2 was prepared according to the procedure described by Ito et al (Scheme S2).57

177

Azidoaniline hydrochloride (501 mg, 2.9 mmol) was dissolved in 60 mL distilled water

178

containing sodium carbonate (933 mg, 8.8 mmol), and cooled down to 4 °C by using an ice

179

bath. Then, methacryloyl chloride (492 µL, 5.0 mmol) in 10 mL of dioxane was added

180

dropwise and the reaction mixture was stirred for 3 h in dark. The obtained precipitate was

181

filtered off, washed with distilled water and dried under reduced pressure. The resulting

182

product was obtained as a grey powder with a yield of 0.64 g (~ 100%). For 1H NMR

183

spectrum of M2, see Figure S2.

184 185

1

H NMR (400 MHz, CDCl3, δ, ppm): 2.05 (s, 3H, CH3), 5.45 (s, 1H, =CH2), 5.77 (s, 1H,

=CH2), 6.97-6.99 (dd, 2H, NH-Ar-H), 7.53-7.56 (dd, 2H, N3-Ar-H).

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Copolymerization of M1 and M2

188

For the preparation of copolymers, two feeding molar ratios of M1/M2 were selected, namely

189

97.8/2.2 (copolymer C1) and 96.6/3.4 (copolymer C2). C1 was prepared as follows: M1 (320

190

mg, 1.4 mmol) was first dissolved in 3 mL of water/ethanol mixture (1:1 v/v), then M2 (6.2

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mg, 0.031 mmol) and AIBA (8.2 mg, 0.03 mmol) were added. Three freeze-pump-thaw

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cycles were applied and the reaction flask was filled with argon. The reaction mixture was 8 ACS Paragon Plus Environment

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stirred for 4 h at 60 °C in dark. After cooling down, the reaction mixture was dialyzed against

194

a water/ethanol mixture (1:1 v:v, 3x) and distilled water (2x) using a Spectra/Por® dialyzing

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tubing (MWCO 3.5 kDa, Spectrum Laboratories, Inc., CA, USA). After freeze-drying, 195

196

mg (yield 60%) of white powder were obtained. C2 was prepared with a similar procedure in

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a yield of 287 mg (yield 71%). The dialysis steps as well as all subsequent manipulation with

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copolymers C1 and C2 in either powder or solution forms were performed in dark.

199 200

1

H NMR (400 MHz, D2O, δ, ppm): 3.76 (s, N+-CH2-COO-, 2H), 3.50 (m, -CH2-N+, 2H), 3.10

(m, (CH3)2N+ and NH-CH2-, 8H), 1.85 (s, -CH2-, 2H), 0.97-0.84 (m, CH3, 3H).

201

202

Preparation of polyimide covered silicon wafers and round glass coverslips

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Two different types of PI plates were prepared on different substrates, namely (i) round glass

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coverslips (diameter: 6 mm), and (ii) square silicon (Si) wafers (2 x 2 cm2). Si wafers are

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typically used as substrate for the preparation of PI based neural interfaces, by lithographic

206

and dry etching techniques, while for in vitro experiments glass coverslips were identified as a

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more appropriate solution, so not to provide undesired biological response and to have rather

208

transparent substrates. Anyhow, on both substrate types a uniform PI layer was achieved. For

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(i), a layer of PI resin (PI2610, HD Microsystems, Germany) was spun at 10,000 rpm for 30 s

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onto the polished glass coverslips. Samples were then cured at 130 °C for 120 s and

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subsequently hard baked in oven with nitrogen flux at 350 °C for 1 h (Carbolite, UK). Finally,

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PI samples were washed twice with a liquid detergent (RBS 35 Concentrate, Belgium) in

213

order to eliminate contaminants and dust. For (ii), Si wafers (Si-Mat, Germany) were cut into

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square samples (2 x 2 cm2) by using a diamond tip. Samples were cleaned in acetone for 10

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min, rinsed with deionized water, cleaned with isopropanol for 10 min, and dried. Two layers

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of PI resin were subsequently spin-coated at 2,000 rpm for 30 s. Samples were cured at 130

217

°C for 120 s and hard baked in oven with nitrogen flux at 350 °C for 1 h. 9 ACS Paragon Plus Environment

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219 220

Deposition of copolymers C1 and C2 onto polyimide plates

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Both types of PI plates were coated with the C1 and C2 copolymers. PI plates were first

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washed in distilled water (3 min), ethanol (3 min), and dichloromethane (10 min) and

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subsequently dried at room temperature and atmospheric pressure. Dry plates were then

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uniformly covered with an aqueous solution of Pluronic F-127 with a concentration of 0.1 and

225

0.5 wt. %, respectively, using the solution volumes of 20 µL, in the case of (i) plates, and 283

226

µL, in the case of (ii) plates. Plates were dried overnight at room temperature and atmospheric

227

pressure. Aqueous solutions of C1 and C2 with a concentration of 0.5 and 1.0 wt. %,

228

respectively, were subsequently pipetted and homogeneously spread on the plate surface (20

229

µL in the case of (i) plates and 283 µL in the case of (ii) plates) to create a uniform layer of a

230

copolymer solution. Plates were dried overnight at room temperature and atmospheric

231

pressure in dark. Then they were irradiated by UV light for 20 min (UV Black Ray B100-A,

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100 Watt, 365 nm, distance 5 cm) with a light intensity of 20-24 mW/cm2. These irradiation

233

conditions resulted from screening experiments that provided hydrogel layers permanently

234

attached to the the PI surfaces (Table S1). After irradiation, the samples were washed in

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distilled water for 5 min and dried overnight at room temperature and atmospheric pressure

236

prior to characterization. The procedure used for formation of zwitterionic hydrogel layers

237

attached to PI surfaces is illustrated in Figure 1. The long-term storage conditions were room

238

temperature and protection from dust and light.

239 240

We selected the combinations of two copolymer types differing in the content of 4-

241

azidophenyl group, two different copolymer concentrations and two different surfactant

242

concentrations, which all were expected to affect the thickness and mechanical properties of 10 ACS Paragon Plus Environment

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resulting hydrogel layers. In summary, six different protocols for fabricating polyzwitterionic

244

hydrogel layers covalently attached to the PI plates were selected using the experimental

245

conditions listed in Table 1. The hydrogel layers formed at the surface of PI plates were

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characterized by attenuated total reflectance Fourier-transform infrared spectroscopy, white

247

light interferometry, confocal laser scanning microscopy, water contact angle measurements,

248

and nanoindentation (details given below).

249

250 251 252 253 254

Figure 1. Procedure for covalent modification of PI plates by zwitterionic hydrogel layer: (1) cleaning (in distilled water for 3 min, in ethanol for 3 min, and in dichloromethane for 10 min); (2) deposition of the surfactant solution; (3) deposition of the copolymer solution; (4) irradiation step, (5) cleaning (in distilled water for 5 min), and (6) characterization.

255 256

Nuclear magnetic resonance spectroscopy

257

Nuclear magnetic resonance spectroscopy (NMR) measurements were used to characterize

258

the structure of monomers and copolymers. 1H NMR spectra were measured at room

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temperature in deuterated solvents (D2O, CDCl3) using a Varian 400-MR spectrometer

260

(Varian, USA) and tetramethylsilane (TMS) as an internal standard. 11 ACS Paragon Plus Environment

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Attenuated total reflectance Fourier-transform infrared spectroscopy

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Attenuated total reflectance Fourier-transform infrared spectroscopy (ATR-FTIR) spectra of

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monomers, copolymers and hydrogel layers deposited on (ii) were measured by means of a

266

Nicolet 8700 FT-IR spectrometer (Thermo Scientific, USA) equipped with a Nicolet

267

Continuum Microscope and a germanium ATR crystal. Spectra were collected with a

268

resolution of 4 cm-1 using 128 scans. A software Omnic 8 (Thermo Scientific, USA) was used

269

for spectra evaluation.

270 271

UV/Vis absorbance spectroscopy

272

UV/Vis absorbance spectroscopy was used to determine the amount of M2 (N-(4-

273

azidophenyl)-2-methylprop-2-enamide) in C1 and C2 copolymers using a Shimadzu 1650 PC

274

spectrometer (Kyoto, Japan). A calibration curve was prepared from UV/Vis data for

275

monomer M2 dissolved in distilled water at the concentrations ranging from 0.01563 to

276

0.00125 g.L-1 (6.2 x 10-6 to 7.7 x 10-5 mol.L-1) and plotting the absorbance maxima at 273 nm

277

against concentration (Figure S3). Then the absorption spectra of C1 and C2 copolymers in 1

278

g.L-1 aqueous solutions were measured and molar content of M2 (cM2) in copolymers was

279

calculated from equation (1):

௡ಾమ

ಲమళయ

280

ܿெଶ (mol %) =

281

where A273 is the absorbance at 273 nm, MM1 and MM2 are the molar masses of the

282

corresponding monomers (MM1 = 228.29 g.mol-1, MM2 = 202.21 g.mol-1), nM1 and nM2 are the

283

molar fractions of monomers and ε is the extinction coefficient of monomer M2 (equivalent

284

to 4-azidophenyl group), equal to 10 680 L.mol-1.cm-1.

௡ಾమ ା ௡ಾభ

*100 =

ε ಲమళయ ಲమళయ భష ε ∗ಾಾమ ା ಾಾభ ε

*100

285


(1)

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Water contact angle

287

The hydrogel water contact angle was measured using the sessile drop technique with a

288

Surface Energy Evaluation system (SEE system with CCD camera, Advex Instruments,

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Czech Republic). Dried hydrogel surfaces were wetted with a 3 µL water drop. For each

290

sample, 6 independent measurements were performed.

291

292

White light interferometry

293

Topography, surface roughness and thickness of hydrogel layers deposited on (ii) in their dry

294

state were assessed by a white light interferometry (WLI). A GT-K1 interferometer (Bruker,

295

USA) in vertical scanning interferometry mode, using a broad-band white light source, was

296

employed. Data were acquired and evaluated using a dedicated control and analysis software

297

package Vision 64. The topography was analyzed at the sample center, in order to capture the

298

topography of a continuous hydrogel layer, and at the edge, in order to capture the topography

299

of the layer and the PI substrate simultaneously. The surface area of approximately 4.5 x 3.5

300

µm2 was assessed for seven different positions of each hydrogel. A dry hydrogel layer

301

thickness was determined from controllably scratched layers in a swollen state (in distilled

302

water) using a scalpel to remove hydrogel and uncover the PI substrate. Subsequently, the

303

samples were washed with distilled water and dried for 24 h at room temperature and

304

atmospheric pressure prior to the WLI analysis. The reported thickness values represent the

305

average of six independent measurements made for each hydrogel type.

306 307

Confocal laser scanning microscopy

308

Confocal laser scanning microscopy (CLSM) in reflection and fluorescence modes was used

309

to determine the thickness of hydrogels deposited on (ii) in a swollen state. A LSM 510

310

META scanning confocal microscope head on Axiovert 200M inverted microscope (Zeiss, 14 ACS Paragon Plus Environment

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Langmuir

311

Germany) was used with a 488 nm line of Ar:ion laser for sample excitation. Fluorescence

312

emission was detected through a 500-550 nm bandpass emission filter, the reflected laser light

313

was detected through a 435-485 nm bandpass filter and both channels were featured by a

314

pinhole setting of 1 Airy unit. Samples for CLSM measurements were prepared by careful

315

peeling off the PI layer with deposited hydrogel coating from the wafer surface. These

316

samples were placed between two microscopic coverslips separated by a glass spacer, made

317

of the coverslip, and stained with 0.5 mg/mL solution of IgG labelled with Alexa488

318

(ThermoFisher Scientific, USA) in PBS for 10 min prior to a measurement. All measurements

319

were done at room temperature. Images were obtained by scanning in the XZ plane with a 16x

320

line-averaging protocol using either PlanApochromat 20x/0.75 or C-Apochromat 40x/1.2W

321

Corr objectives. At least 5 independent images were acquired, with a total number of 15

322

measured height values determined for each hydrogel type, and averaged.

323

324

Nanoindentation via atomic force microscopy

325

Atomic force microscopy (AFM) in force spectroscopy mode58-59 was carried out to determine

326

the Young's modulus of hydrogels deposited on (ii) by using a Nanowizard III (JPK

327

Instruments, Germany) set-up. A Si nitride cantilever (DNP, Bruker, USA) with a nominal

328

spring constant of 0.24 N/m was used. The tip shape was pyramidal and the nominal radius at

329

the tip apex was 20 nm. The actual spring constant of each cantilever was determined by

330

using the in situ thermal noise method.60 All measurements were performed in a liquid

331

environment (PBS, Sigma-Aldrich, USA) at room temperature. The maximum force applied

332

on the samples was 2 nN. The velocity of the piezo-scanner was maintained constant at 3

333

µm/s. Force curves were corrected for the cantilever bending61 to calculate the tip-sample

334

separation and to build force vs indentation (F-I) curves. The Young’s modulus of the sample

335

was calculated by fitting the corresponding F-I curves with the Bilodeau model for a 15 ACS Paragon Plus Environment

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Page 16 of 44

336

pyramidal indenter.62 A proper fitting of the F-I curves was performed by means of a

337

dedicated data processing software provided by JPK Instruments. Five independent samples

338

were analyzed for each hydrogel type and each sample was indented at least in 50 different

339

points, randomly selected over the entire surface area.

340

341

Cell cultures and in vitro characterization

342

In vitro evaluation protocols were performed on the most important cell types responding

343

cooperatively to implanted biomaterials, namely fibroblasts and macrophages, as reported in

344

the sub-sections below. A qualitative assessment was also performed on a neural model (SY-

345

SY5Y cell line, CRL-2266™, ATCC®, Manassas, Virginia, USA), placed in close proximity

346

to the substrates. For this test, SY-SY5Y cells were seeded at a density of 30,000 cell/cm2,

347

cultured for 24 h in a mixture 1:1 of Dulbecco’s Modified Eagle’s Medium and Nutrient

348

Mixture F-12 (DMEM/F-12, Gibco™-Thermo Fisher Scientific, Waltham, Massachusetts,

349

USA), supplemented with Fetal Bovine Serum (FBS, Gibco™-Thermo Fisher Scientific,

350

Waltham,

351

Penicillin/Streptomycin

352

Massachusetts, USA) to a final concentration of 1 %, and then imaged in bright field to check

353

their presence and shape.

Massachusetts,

USA)

to

(PEN/STREP,

a

final

concentration

Gibco™-Thermo

Fisher

of

10%,

Scientific,

and

with

Waltham,

354 355

Fibroblast culture, adhesion and cytotoxicity tests

356

Normal human dermal fibroblasts, nHDFs (Cat. # CC-2511, Lonza, Basel, Switzerland) were

357

cultured using a complete growth medium composed of high-glucose Dulbecco’s Modified

358

Eagle’s Medium (DMEM, Gibco™ - Thermo Fisher Scientific, Waltham, Massachusetts,


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359

USA) with phenol-red supplemented with FBS to a final concentration of 10%, and with

360

PEN/STREP to a final concentration of 1 %.

361

For adhesion tests, round glass coverslips (6 mm diameter) covered with PI (i) (control

362

samples) and with PI further coated with selected zwitterionic hydrogels were used. Briefly,

363

the samples were placed in 48-well plates, washed once with PBS and then incubated for 1 h

364

in PBS, in order to let hydrogels swell. Then, samples were washed again with PBS

365

supplemented with PEN/STREP (1 %) and conditioned with the complete growth medium for

366

5 h. nHDFs were then seeded on the samples at a density of 15,000 cells/cm2 and cultured for

367

24 h. To evaluate cell adhesion on the different substrates, cells were stained for nuclei with

368

Hoechst 33342 (Molecular Probes™-Thermo Fisher Scientific, Waltham, Massachusetts,

369

USA). Hoechst 33342 was diluted to a final concentration of 10 µg/mL in Dulbecco's PBS

370

(DPBS, Gibco™-Thermo Fisher Scientific, Waltham, Massachusetts, USA), added to cells

371

and kept for 5 min at room temperature. Glass plates were then turned upside down to bypass

372

PI autofluorescence and imaged by means of a UV/Vis microscope (ECLIPSE Ti-E, Nikon

373

Corporation, Minato-ku, Tokyo, Japan). Images of the blue channel (stained nuclei, bandpass

374

filter value equal to 461 nm) were acquired at 10x magnification and the number of nuclei per

375

image was quantified through the ImageJ® software (National Institutes of Health, Bethesda,

376

Maryland, USA; available at https://imagej.nih.gov/ij/). Three independent plates were

377

analyzed for each sample type, including the control, to quantitatively evaluate cell adhesion.

378

Three images on different areas were acquired for each sample.

379

Whole cell morphology was also qualitatively assessed. To this purpose, at the same time-

380

point (24 h) cells were fixed with a paraformaldehyde (Sigma Aldrich® Corporation, St.

381

Louis, Missouri, USA) solution in PBS (4 % w/v) for 15 min, then permeabilized with 0.1 %

382

Triton X-100 (Sigma Aldrich® Corporation, St. Louis, Missouri, USA) in PBS for 15 min and

383

finally co-stained for 15 min for actin fibers with tetramethylrhodamine B isothiocyanate 17 ACS Paragon Plus Environment

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384

labeled Phalloidin (Sigma Aldrich® Corporation, St. Louis, Missouri, USA) as well as for

385

nuclei with Hoechst (as described above). Fluorescence images (blue and red channels,

386

bandpass filter values equal to 461 and 570 nm) were acquired at 10x magnification.

387

For cytotoxicity tests, a Pierce LDH Cytotoxicity Assay kit (Thermo Fisher Scientific,

388

Waltham, Massachusetts, USA) was used. Lactate dehydrogenase (LDH) is a cytosolic

389

enzyme that is released in the cell culture medium after the cell plasma membrane is

390

damaged. LDH can be quantified spectrophotometrically by means of a coupled enzymatic

391

reaction that leads to the formation of red insoluble formazan, which is directly proportional

392

to the amount of LDH released. To perform this experiment, round glass coverslips (6 mm

393

diameter) covered with PI (i) (control samples) and with PI further coated with selected

394

zwitterionic hydrogels were placed into 96-well plates, pretreated with PBS and complete

395

growth medium as described above. nHDFs were then seeded at a density of 25,000 cells/cm2

396

(200 µL of cells suspension in complete growth medium, without phenol red that would

397

interfere with the measurements, per each well) both on covered round glass coverslips (3

398

coverslips per sample type) to assess substrate-induced LDH activity, and on tissue culture

399

polystyrene wells to assess spontaneous and maximum LDH activity (3 wells per type). Cells

400

were cultured for 24 h, then treated according to the LDH kit manufacturer’s instructions

401

properly adapted to the samples. Absorbance at 490 nm (related to LDH amount) was

402

measured (3 replicates for each sample) with a plate reader (Victor X3, Perkin Elmer,

403

Waltham, Massachusetts, USA). Background was removed and the cytotoxicity was

404

computed for each hydrogel and for PI as: % Cytotoxicity = (substrate-induced LDH activity -

405

spontaneous LDH activity) / (maximum LDH activity - spontaneous LDH activity).

406 407

Macrophage culture, adhesion and activation tests


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The murine RAW 264.7 macrophage cell line (ATCC® TIB-71™, ATCC®, Manassas,

409

Virginia, USA) was cultured in the same complete growth medium used for nHDFs.

410

For adhesion tests, round glass coverslips covered with PI (i) (control samples) and with PI

411

coated with selected zwitterionic hydrogels were placed in 48-well plates and pretreated

412

similarly to fibroblasts (see above). RAW 264.7 cells were then seeded at a density of 30,000

413

cells/cm2 and cultured for 24 h. The cell nuclei were stained with Hoechst 33342 and imaged,

414

similarly to nHDFs. Three independent plates were analyzed for each sample type, including

415

the control, and three images were acquired for each sample on different areas.

416

Similarly

417

tetramethylrhodamine B isothiocyanate labeled Phalloidin and Hoechst. In order to compare

418

macrophage morphology changes due to activation, the same cell density was seeded on

419

tissue culture polystyrene wells and left to attach for 4 h. Then, lipopolysaccharide (LPS)

420

from Escherichia coli 055:B5 (Sigma Aldrich® Corporation, St. Louis, Missouri, USA) was

421

added to a concentration of 1 µg/mL and incubated for 24 h, to provide a positive control for

422

activation. Then, cells underwent fixation, permeabilization and staining as described above.

423

Macrophage activation was evaluated through the assessment of nitric oxide production.

424

Nitrite (NO2-), a stable breakdown product of NO, was measured by means of a Griess

425

Reagent system (Promega Corporation, Madison, Wisconsin, USA). Round glass coverslips

426

(6 mm diameter) covered with PI (i) (control samples) and with PI further coated with

427

selected zwitterionic hydrogels were placed in 96-well plates and pretreated as described

428

above. RAW 264.7 were seeded at a density of 80,000 cells/cm2 (150 µL of cells suspension

429

in a complete growth medium without phenol red per each well) both on samples and on bare

430

tissue culture polystyrene. The latter samples were used as negative and positive controls to

431

assess nitrite release from activated macrophages: the cells were cultured for 4 h, then 50 µL

to

nHDFs,

qualitative

cell

imaging

was


also

performed

through

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432

of complete growth medium with/without LPS (1 µg/mL) was added to induce macrophage

433

activation, while all other PI samples were only provided with complete growth medium in

434

order to compensate for volume increase. Cells were exposed to the different substrates or to

435

LPS for 24 h, then the Griess assay was performed, according to the manufacturer’s

436

instructions. The absorbance at 560 nm for each sample was measured with a plate reader

437

(Victor X3, Perkin Elmer, Waltham, Massachusetts, USA). Three independent samples were

438

analyzed for each sample type and three absorbance reading replicates were measured for

439

each sample.

440

441

Data analyses

442

Data underwent a non-parametric statistical analysis using a Kruskal-Wallis post-hoc test and

443

a Dunn's multiple comparison test in order to evaluate significant differences among the

444

samples. Results were considered statistically different for p-values ≤ 0.05.

445 446

RESULTS AND DISCUSSION

447

Preparation and photoimmobilization of zwitterionic copolymers

448

Two photoreactive zwitterionic copolymers C1 and C2, differing in the content of photolabile

449

4-azidophenyl groups, were prepared by the radical copolymerization of 2-{dimethyl[3-(2-

450

methylprop-2-enamido)propyl]ammonio}acetate (M1) with N-(4-azidophenyl)-2-methylprop-

451

2-enamide (M2) bearing photoreactive 4-azidophenyl unit (Figure 2). The chemical structure

452

of monomers and copolymers was determined by NMR and ATR-FTIR spectroscopies. The

453

presence of small broad signals in the range from 7 to 7.5 ppm in 1H NMR spectrum revealed

454

the presence of 4-azidophenyl group in the structure of both copolymers (Figures S4 and S5). 20 ACS Paragon Plus Environment

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455

In the ATR-FTIR spectra of C1 and C2, a small vibration band of the azide group was

456

observed at 2119 cm-1 (Figure S6). Asymmetric and symmetric stretching vibrations of

457

carboxylate group of zwitterionic unit were present at 1630 and 1380 cm-1. Vibration bands

458

connected to N-H bond of zwitterionic unit were also visible at 1529 and 1483 cm-1. All

459

signals were in accordance with results published recently by Sobolčiak et al.53

460 461 462

Figure 2. Scheme of the synthesis of zwitterionic copolymers C1 and C2 containing photolabile 4-azidophenyl groups.

463

464

Molar percentages of 4-azidophenyl groups in copolymers C1 and C2 of 1.6 and 3.1 mol.%,

465

respectively, were calculated from UV/Vis absorbance spectra of copolymers in water using

466

the calibration curve for M2 (Figure S3) and equation (1). These values were slightly lower

467

compared to feeding M2 comonomer content equal to 2.2 and 3.4 mol.% for C1 and C2,

468

respectively, probably due to an incomplete conversion and a minor compositional drift in

469

favor to consumption of M1 comonomer.

470

471

Characterization of PI plates modified by polyzwitterionic layers


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472

The aim of this study was to identify the conditions for a covalent attachment of soft

473

zwitterionic hydrogel layers (with a thickness in the micrometer range) to the PI surface. In

474

order to achieve this goal, we adapted principles that employ a photoactivated attachment of

475

zwitterionic polymers via a C-H insertion reaction using 4-azidophenyl groups,49,

476

schematically shown in Figure 3.

53

477 478 479 480 481 482 483 484 485 486 487 488 489 490 491

Figure 3. Schematic representation of crosslinking and grafting process of zwitterionic copolymers C1 and C2 by C-H insertion of nitrene, created from a photolabile 4-azidophenyl groups during UV-light irradiation (365 nm), with the hydrocarbon units of either zwitterionic copolymer or polyimide surface.

492 493

The experimental conditions used for hydrogel fabrication, such as distance between the UV

494

lamp and the surface, irradiation time, copolymer and surfactant concentrations, were

495

optimized to obtain hydrogel layers stably bound to the surface (Table S1). Based on these

496

data, we selected the irradiation conditions, which were kept constant for all experiments.

497

Although the studied parameters are mutually interrelated, we could conclude that the stability

498

of attached layers decreases with increased copolymer concentration (≥ 3 wt.%) and longer 22 ACS Paragon Plus Environment

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Langmuir

499

lamp distance (> 5 cm). Similar observations were reported by Roger et al.,63 where the

500

authors observed decreasing of grafting yield with increased either concentration or volume of

501

carbohydrate bearing 4-azidophenyl group. This was explained by lower penetration depth of

502

UV light leading to ineffective grafting in case of thicker layers.

503 504 505 506 507

Table 1. Preparation conditions and properties of zwitterionic hydrogel coatings covalently attached to PI substrates. Water contact angles, dry (hdry) and wet (hwet) thickness, surface roughness (Ra) in a dry state, and Young’s moduli are expressed as the mean value ± standard deviation.

Sample

M2 in Copolymer Pluronic Contact copolymer F-127 angle a) [wt.%] [wt.%] [°] [mol.%]

[µm]

[µm]

[µm]

Young’s moduluse) [kPa]

59 ± 3

5.1 ± 0.3

0.16 ± 0.06

13 ± 2

2.7 ± 0.2

1.0

0.5

66 ± 9

6.4 ± 0.4

0.10 ± 0.03

32 ± 5

2.0 ± 0.2

0.5

0.1

33 ± 6

4.6 ± 0.6

0.13 ± 0.05

28 ± 6

12 ± 3

H4

1.0

0.1

45 ± 4

5.2 ± 0.2

0.30 ± 0.10

24 ± 6

15 ± 2

H5

0.5

0.5

34 ± 6

4.8 ± 0.9

0.14 ± 0.02

14 ± 3

5.2 ± 0.8

H6

1.0

0.5

39 ±4

3.1 ± 0.4

0.12 ± 0.05

24 ± 4

19 ± 4

H3

511 512 513 514 515 516

hwetd)

0.1

H2

508 509 510

Rab)

1.0

H1

1.6

hdryb,c)

3.1

a)

the measured water contact angle for PI plate was 75° determined by white light interferometry for dried hydrogels c) hdry theoretical values are 3.5 and 7.0 µm for coatings using 0.5 and 1.0 wt.% of copolymer, respectively, assuming Pluronic F-127 does not take part in the layer formation and the copolymer density is 1 g.cm-3 d) determined by confocal laser scanning microscopy for wet hydrogels in PBS e) determined by AFM nanoindentation for wet hydrogels in PBS b)

517

The characterization of polyzwitterionic hydrogel layers on PI substrates included surface

518

chemical analysis and measurement of wettability, thickness, surface roughness and

519

mechanical properties. These data are summarized in Table 1, together with the conditions

520

used for fabricating the six representative hydrogel types (H1 - H6).

521


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522

The modification of PI substrates by hydrogel layers was characterized by ATR-FTIR, shown

523

in Figure 4, using non-modified PI substrate and C1 and C2 copolymers as controls. In

524

agreement with previously published data,64 the spectrum of non-modified PI consists of

525

peaks, which can be assigned to the asymmetric stretching vibration of C=O group at 1773

526

cm-1, the symmetric stretching vibration of C=O group at 1708 cm-1, C-N stretching vibration

527

at 1352 cm-1 and imide five-ring deformation vibration at 734 cm-1 (Figure S7a). The PI and

528

copolymer vibration peaks in the ATR-FTIR spectra of hydrogel layer-coated PI plates were

529

present at different intensities depending on the sample type (Figure 4). The layers prepared

530

of lower 0.5 wt.% copolymer concentration (H3 and H5, Figure 4b) showed a significant

531

presence of PI vibrations in the spectra compared to the layers prepared of 1.0 wt.%

532

copolymer concentration. This observation was surprising as it did not correlate with the WLI

533

analysis showing, in Figure 5, a complete coverage of PI plates, the thickness, hdry, in the

534

range from 3 to 6 µm (Table 1), and considering the penetration depth of the infrared beam

535

approximately 1 µm for Ge crystal.65 Figure 5 shows the WLI images for hydrogel layers in

536

their centers, verifying a complete PI surface coverage and a high degree of layer smoothness,

537

which was quantitatively expressed by the surface roughness values, Ra, ranged from 0.1 to

538

0.3 µm (Table 1). Figure 5 also reveals the images of the plate edges to visualize the situation

539

for imperfect PI surface coverage that is in contrast to complete coverage of the sample

540

central areas. The Ra values resulted less than 10 % of the thickness values for the dry layers,

541

determined as the height difference between the surface of hydrogel and the PI surface

542

obtained after careful removal of the hydrogel layers by a scalpel, in a swollen state (Figure

543

S8). Overall, the ATR-FTIR and WLI data show that the PI vibration peaks in the ATR-FTIR

544

spectra shown in Figure 4 could not result from a non-coated PI surface or from the extensive

545

surface roughness. We propose that the presence of the PI in the FTIR spectra was rather due

546

to using ATR mode for these very soft hydrogel layers, featured by a thickness of a few


Page 24 of 44

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Langmuir

547

micrometers. During the analysis, the hydrogel layers were likely compressed by the ATR

548

crystal towards the hard substrate (PI on silicon wafer, (ii)) resulting in the analytical

549

thickness that was within the infrared beam penetration depth.

550 551

The ATR-FTIR spectra for all hydrogel layers in Figure 4 highlighted also the position of the

552

high intensity C-O stretching vibration at 1102 cm-1 for Pluronic F-127 (the individual ATR-

553

FTIR spectra for Pluronic F-127 are shown in Figure S7b). Pluronic F-127 was used as a

554

surfactant during the hydrogel fabrication, for wetting the PI hydrophobic surface prior to

555

deposition of the copolymer solution. Its presence in the hydrogel layer was assumed since

556

the C-H insertion reaction may proceed also with the alkyleneoxy units of Pluronic F-127

557

molecule. The Pluronic F-127 concentration is in the range of copolymer concentration (0.1

558

and 0.5 wt.%, Table 1); therefore, its high intensity vibration at around 1100 cm-1 was

559

expected to be detected. However, this was not the case (Figure 4) and it is not trivial to

560

understand whether this behavior corresponds with a negligible incorporation of this

561

surfactant into the hydrogel layer followed by its gradual washing from the hydrogel, or with

562

the overlap of PI and Pluronic F-127 bands in this region of FTIR spectra.

563

564

The hydrogel layer deposited on PI substrates modified the surface wettability. The contact

565

angle determined for a non-modified PI plate was 75°, a value close to the one reported in the

566

literature (72°).66 The coating of PI plates by zwitterionic hydrogels reduced this contact angle

567

to values included in the range between 33° and 66°, as reported in Table 1. These values are

568

in the same range as previously reported for other zwitterionic polymer layers, with a

569

thickness typically in the sub-micrometer range.49,

570

hence, a slightly more hydrophobic surface, was observed for layers formed from C1

571

copolymer of lower 4-azidophenyl content compared to the layers based on C2 copolymer,

52-53, 66-68


A higher contact angle and,

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572

especially at the lower C2 content of 0.5 wt.%. During the formation of these hydrogel layers,

573

two competing C-H

574

575 576 577 578 579

Figure 4. ATR-FTIR spectra of (a) H1 and H2 hydrogel layers compared to PI surface and copolymer C1 and (b) H3-H6 hydrogel layers compared to PI surface and copolymer C2, in a wavenumber region from 2200 to 600 cm-1. Dashed lines depict the characteristic bands for PI, copolymers and Pluronic F-127. 26 ACS Paragon Plus Environment

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Langmuir

580


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581 582 583 584

Figure 5. White light interferometry visualization of hydrogel coatings on a PI substrate in the sample center and edge, respectively, for selected samples H1-H2 (made of the copolymer C1), H3-H6 (made of the copolymer C2). The main preparation conditions of hydrogels are given in Table 1.

585

586

insertion reactions of nitrene radicals proceed concurrently, i.e., with the PI surface and within

587

the copolymer chains, resulting in the attachment of the hydrogel layer to the surface and the

588

intra- and intermolecular crosslinking of C1 resp. C2 copolymer chains, respectively (Figure

589

3). The competition between these two pathways was suggested to control the water contact

590

angle for poly(carboxybetaine) layers deposited by photoactuation to various polymer

591

substrates.53 In our case, the obtained water contact angle data provide no clear reason why

592

the hydrogels made of C1 copolymer would experience a higher intra- and intermolecular

593

crosslinking compared to hydrogels made of C2 copolymer with a slightly higher content of

594

4-azidophenyl group. This issue remains open also in view of the Young’s modulus data

595

(reported in Table 1 and discussed below), showing that the hydrogel layers based on C2

596

copolymer are stiffer than those based on C1 copolymer, under identical conditions.

597 598

Hydrogel thickness values in a swollen state, hwet, ranged between 13 and 32 µm (Table 1).

599

They were determined by a CLSM technique, as shown in Figure 6. This is an important

600

feature of the layers, for their future deposition on the surface of real neural electrodes. This

601

method enabled to precisely visualize the swollen layer as a dark object well-visible in the

602

PBS solution containing the IgG labeled with Alexa 488 (Figure 6a). The labeled IgG did not

603

significantly penetrate inside the hydrogel volume within the time-scale of CLSM analysis.

604

This arrangement eliminated the influence of strong fluorescence emission of the PI layer

605

compared to Alexa 488-labeled IgG (Figure 6b). The CLSM images of hydrogels H2, H5 and

606

H6 (i.e., the samples used for the in vitro experiments discussed further below) are shown in 28 ACS Paragon Plus Environment

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Langmuir

607

Figure 6c. The hwet values did not correlate with either preparation conditions or hdry values.

608

In addition, the thickness of individual swollen hydrogels (Figure 6c) exhibited a significant

609

variation that was not expected from the low Ra values obtained for the dry layers (Table 1).

610

This suggests that the uneven swelling of hydrogel layers may be caused by a crosslinking

611

reaction that did not proceed homogeneously within the entire hydrogel volume irradiated in

612

the dry state.

613 614 615 616 617 618 619

Figure 6. Determination of the thickness of hydrogel layers swollen in PBS by CLSM. (a) Schematic illustration of experiment arrangement for CLSM measurement: 1. Glass coverslip, 2. Solution of Alexa 488-labeled IgG in PBS (c = 0.5 mg/mL), 3. PI layer, 4. Hydrogel layer, 5. Glass spacer, (b) comparison of fluorescence spectra of PI layer and solution of Alexa 488labeled IgG in PBS, and (c) CLSM images of H2, H5 and H6 in solution of Alexa 488-labeled IgG in PBS (scale bar is 20 µm).

620

621

The hydrogel Young’s moduli, determined in a PBS liquid environment by AFM

622

measurements (Figure S9) are reported in Table 1 and plotted in Figure 7. The layers made of

623

C1 copolymer with a lower 4-azidophenyl content exhibited Young’s modulus of around 2

624

kPa. A higher content of 4-azidophenyl groups in the C2 copolymer resulted in a Young’s

625

modulus increase up to 20 kPa depending on the copolymer concentration. Young’s modulus 29 ACS Paragon Plus Environment

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626

data do not correlate with the content of Pluronic F-127. Rather, they are associated with the

627

network density, which is in turn proportional to the 4-azidophenyl content and to the

628

polymer concentration. Note that the higher network density for C2 copolymer was not

629

demonstrated by the hwet with expected lower thickness in the swollen state than for C1

630

copolymer due to lower swelling. Neither the higher network density for C2 copolymer

631

resulted in a higher water contact angle compared to hydrogels made of C1 copolymer. Even

632

though some of the obtained data would require further understanding, the employed strategy

633

successfully served the primary aim of this study that was to covalently coat stiff PI plates by

634

zwitterionic hydrogels of low stiffness, thus approaching the stiffness of neural tissues.

635 636 637 638 639 640

Figure 7. Young’s moduli of zwitterionic hydrogels H1-H6 deposited on PI plates, and of the PI plate assessed by AFM nanoindenting in PBS. Box and whisker plots show the median values and interquartile ranges (whisker represent the maximum and minimum values obtained). **** = p < 0.0001.


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641

642

In vitro experiments

643

In vitro tests were performed on three hydrogel types coated on the PI plates i): H2, H5 and

644

H6. These were selected based on their Young’s modulus and chemical composition. In

645

particular, we selected the hydrogels with the lowest and the highest Young’s modulus,

646

respectively, H2 and H6, which featured the same amount of copolymer and Pluronic F-127,

647

but different content of 4-azidophenyl groups. As a third sample type, we selected the H5

648

hydrogel, which showed an intermediate value of Young’s modulus and the same Pluronic F-

649

127 content as H2 and H6 samples (see Table 1).

650 651

Fibroblast adhesion and morphology were evaluated by fluorescence imaging of living and

652

fixed nHDF cells cultured on H2, H5 and H6 hydrogels and on PI controls. Figure 8a shows

653

the fluorescence patterns of actin (stained in red) and nuclei (stained in blue) in cells attached

654

on the different substrates. In particular, almost no adhesion was found on H2 samples, while

655

cells formed few clusters in scattered areas of the H5 and H6 samples. In contrast, cells on PI

656

samples were well spread on the entire sample surface and there was no cluster formation.

657

The cell density (in terms of) cells/cm2 was evaluated by counting the stained nuclei of living

658

nHDFs attached on the different substrates; the results are reported in Figure 8b. Overall, a

659

strong reduction of cell adhesion was observed for the samples covered with zwitterionic

660

hydrogels compared to PI samples (median value reduction: H2 = 97.4%, H5 = 79.4% and

661

H6 = 72.6 %). Comparing the fibroblast adhesion on the three different hydrogels, a statistical

662

difference was found between sample H2 and H5 (p = 0.0087) and between H2 and H6 (p =

663

0.0003), which suggests that the lowest Young’s modulus of H2 hydrogel reduces the cell


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664

adhesion in a higher degree than for hydrogels H5 and H6 of somewhat higher Young’s

665

modulus.

666 667

The obtained results are in a good agreement with previously published data. Stach et al.

668

described significant reduction of RAT-2 fibroblasts adhesion on poly(sulfobetaine) modified

669

conductive surfaces using electrografting method.69 Significant reduction of RAT-2

670

fibroblasts adhesion was observed also in the case of zwitterionic hydrogels spin-coated on

671

glass coverslips.70 Similarly, the photoimmobilized sulfobetaine and carboxybetaine-based

672

polymeric films significantly inhibited the non-specific adhesion of fibroblast-like STO

673

mammalian cells in comparison with non-modified substrates.52 In the work of Chien et al.,

674

L929 cells well adhered on the non-modified poly(styrene) surface while cell adhesion was

675

inhibited almost completely on the poly(carboxybetaine)-modified surfaces.49 All these data,

676

including those obtained in this work, indicate that immobilization of either sulfo- or

677

carboxybetaine zwitterionic polymers on different surfaces enables to reduce cell adhesion in

678

a significant way. Concerning neural electrode coatings featured by tuned mechanical

679

properties, the sole recent study we can refer to is the work of Spencer et al.39 Here, the

680

authors demonstrated that PEG-based hydrogels with a Young’s modulus of ~ 10 kPa

681

significantly reduced the scarring process in vivo with respect to stiffer implants (glass

682

capillaries, with a Young’s modulus of ~ 70 kPa). The scarring process involves several

683

events, included the adhesion of fibroblasts on the implant surface. Thus, although with

684

several differences, i.e. in vitro vs in vivo data, different materials involved, different stiffness

685

values investigated, our results can be considered in line with what reported by Spencer et

686

al.39 It is worth mentioning that in our work we further clarified the role of stiffness in this

687

process, showing that values well below 10 kPa (almost one order of magnitude smaller)

688

prevent fibroblast adhesion more efficiently. 32 ACS Paragon Plus Environment

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689

690

Cytotoxicity of the different substrates was evaluated using nHDFs fibroblasts by quantifying

691

the LDH release; results are reported in Figure 7c. Almost no cytotoxicity was observed for

692

all the samples (the highest median value was 0.7%, found for the H5 sample), in agreement

693

with previous studies performed with NIH-3T3 fibroblasts on zwitterionic hydrogels.71-72 No

694

statistical difference was found between the different hydrogels and the PI samples.

695 696 697 698 699 700 701 702 703

Figure 8. In vitro evaluation of zwitterionic hydrogels on fibroblast cells. (a) Fluorescence images at 10x magnification of nHDF cells on zwitterionic hydrogels H2, H5, H6 and on PI samples, showing nuclei stained in blue and actin in red (scale bar = 100 µm). (b) Quantification of the adhesion of nHDF fibroblasts on zwitterionic hydrogels and on PI samples after 24 h culture. The box and whisker plots represent the median values and interquartile ranges of the number of adherent cells per cm2 (whiskers represent the maximum and minimum values obtained). ** = p < 0.01, *** = p < 0.001. (c) Quantification of the 33 ACS Paragon Plus Environment

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704 705 706 707 708

cytotoxicity percentage of zwitterionic hydrogels and PI by LDH cytotoxicity assay measured using nHDF fibroblasts after 24 h culture on substrates. The box and whisker plots represent the median values and interquartile ranges (whiskers represent the maximum and minimum values obtained).

709

As for fibroblasts, macrophages adhesion and morphology was evaluated by fluorescence

710

imaging of both living and fixed RAW 264.7 cells cultured on H2, H5 and H6 hydrogels and

711

on PI controls. Fluorescence images of actin (stained in red) and nuclei (stained in blue) of

712

cells attached on different substrates are shown in Figure 9a. Macrophages reflected the

713

fibroblasts behavior: almost no adhesion was observed on H2 samples and cells formed few

714

clusters in scattered areas of H5 and H6 samples, while well spread cells were found on the PI

715

samples, without cluster formation. A large reduction of cell adhesion was observed in the

716

samples covered with zwitterionic hydrogels with respect to PI samples (median value

717

reduction: H2 = 98.9%, H5 = 93.2% and H6 = 90.8 %), see Figure 9b. This cell behavior is in

718

agreement with the one observed in the study of Khandwekar et al., where sulfobetaine

719

molecules were entrapped on poly(methyl methacrylate) surfaces leading to a reduction of

720

RAW 264.7 cells adhesion.73 Comparing adhesion of macrophages on the three different

721

hydrogels, a statistical difference was found between sample H2 and H5 (p = 0.0247) and

722

between H2 and H6 (p = 0.0019). The lower adhesion of macrophages on H2 samples may be

723

again explained by its low Young's modulus value. In literature, a stiffness-based influence on

724

macrophages adhesion has been reported, although on another cell line, THP-1.74 Differently,

725

Blakney et al. indicated that substrate stiffness did not play a role in RAW 264.7 cell

726

attachment. However, in this case, the investigated Young’s moduli were in the range of

727

hundreds of kPa.35

728

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Macrophages activation was evaluated by assessing nitric oxide (NO) production,75 in

730

particular by measurement of nitrite release in the supernatant. NO is a physiological 34 ACS Paragon Plus Environment

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731

messenger and effector molecule involved in several biological processes. It is released in

732

excess during host response and it contributes to pathogenic conditions by promoting

733

oxidative stress. The nitrite concentrations are reported in Figure 9c, together with the values

734

of a positive and a negative controls for comparison, namely cells cultured on tissue culture

735

polystyrene and the same cells stimulated by LPS.76 There was no statistical difference

736

between the samples exhibiting very low absolute concentration for each hydrogel type (the

737

highest median value was found for H5 = 0.3 µM). These values were almost comparable

738

with the negative control of non-activated cell (except for samples H6 and PI for which a

739

lower reduction of nitrite production was found) and much lower than the amount of nitrite

740

released by LPS-activated macrophages used as a positive control for inflammation. This was

741

confirmed by morphological evaluation: on each sample, the cells that attached maintained a

742

round shape (Figure 98a), a morphology that is typical of non-activated cells (Figure 9d, top

743

image), whereas LPS-stimulated cells tend to spread on the surface and thus to increase their

744

adhesion area (Figure 9d, bottom image).

745

746

The four substrates were also incubated in the presence of a human bone marrow

747

neuroblastoma derived cell line, SY-S5Y5, so to qualitative prove their ability to safely

748

adhere without any abnormal morphological changes and keep a healthy shape in close

749

proximity to the coatings. Bright field images of cells 24 h after seeding are shown in Figure

750

S10 of Supporting Information. The hydrogels prevented the adhesion of neural cells,

751

similarly to fibroblasts and macrophages. However, the neural cells were able to adhere and

752

maintain the morphology and keep a healthy shape in close proximity at the interface between

753

hydrogel and glass surfacesto the hydrogels, comparably to the PI substrate. This further

754

confirmed the suitability of the coatings analyzed in this paper, in view of future in vivo

755

experiments. Neural interfaces, in fact, will be inserted into peripheral nerves and the coating 35 ACS Paragon Plus Environment

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756

will be thus in close proximity to neural tissues. The results shown in Figure S10 suggest that

757

such coatings do not induce neurotoxicity and are able to keep the surrounding neural tissue

758

healthy.

759 760 761 762 763 764 765

Figure 9. In vitro evaluation of zwitterionic hydrogels on macrophage cells. (a) Fluorescence images at 10x magnification of RAW 264.7 cells on zwitterionic hydrogels and on PI samples, showing nuclei in blue and actin in red (scale bar = 100 µm). (b) Quantification of the adhesion of RAW 264.7 macrophages on zwitterionic hydrogels H2, H5, H6 and on PI samples after 24h of culture. Box and whisker plots represent the median values and interquartile ranges of the number of adherent cells per cm2 (whiskers represent the maximum 36 ACS Paragon Plus Environment

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and minimum values obtained). * = p < 0.05, ** = p < 0.01. (c) Spontaneous nitric oxide production of cells seeded on zwitterionic hydrogels and PI samples for 24 h, provided also with reference values of non- activated cells cultured on tissue culture polystyrene (PS, negative control), and cells cultured on tissue culture polystyrene and activated by means of LPS stimulation (positive control). Box and whisker plots represent the median values and interquartile ranges of the concentration of nitric oxide produced (whiskers represent the maximum and minimum values obtained). (d) Morphological difference between nonstimulated and LPS-stimulated macrophages seeded on tissue culture polystyrene (PS). LPSstimulated macrophages show activation-dependent morphology, with higher adhesion and spreading, compared to non-stimulated ones. Nuclei are stained in blue and actin in red (scale bar = 100 µm).

778

779

CONCLUSION

780

Due to their biocompatibility, flexibility and patternability with lithographic and dry etching

781

techniques, polyimides (PI) are widely used both as implant encapsulation materials and as

782

substrates for implanted neural interfaces. Despite these promising properties, they cannot

783

prevent the formation of a fibrotic capsule. Several approaches have been reported aiming at

784

reducing the formation of a scar-like tissue around the implant. The inhibition of protein and

785

cell adhesion by means of non-biofouling coatings is one of these strategies. Moreover, it has

786

been shown that the undesired host response to implants is also determined by the mechanical

787

mismatch between the implant and the surrounding tissue. To this aim, in this work we

788

presented a strategy for covalently attaching carboxybetaine polyzwitterionic hydrogel

789

coatings to PI substrates, obtained by means of a photoactivated process. These coatings were

790

featured by low stiffness values in the range of the neural tissues, for their possible future

791

application as coatings for neural interfaces. Modified PI plates were characterized by a set of

792

physico-chemical methods to evidence the presence of coatings, to visualize the completeness

793

of the surface coverage by hydrogels, and to characterize the layer thickness and Young’s

794

modulus. The in vitro response of both fibroblasts and macrophages (the main cell types


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795

involved in the foreign body response) was assessed. The zwitterionic hydrogel coatings

796

reduced the adhesion of both cell lines respect to PI control, showing almost no cytotoxicity

797

and a clear correlation with the hydrogel stiffness. Moreover, the hydrogel layers did not

798

induce macrophages activation, evaluated through the quantification of nitric oxide

799

production. These results demonstrate the non-biofouling properties, low stiffness and UV

800

patternability of zwitterionic hydrogels and thus establish the foundation for their possible in

801

vivo investigation as coatings for PI-based neural interfaces.

802 803

SUPPORTING INFORMATION

804

The Supporting Information is available free of charge on DOI:

805

Syntheses of monomers M1 and M2 (Schemes S1 and S2), 1H NMR spectra of M1 and M2

806

(Figures S1 and S2), Calibration curve for M2 absorbance vs. concentration (Figure S3), 1H

807

NMR spectra of copolymers C1 and C2 (Figures S4 and S5), ATR-FTIR of C1 and C2

808

(Figure S6), ATR-FTIR of PI and Pluronic F-127 (Figure S7), WLI determination of hydrogel

809

thickness (Figure S8), AFM force vs indentation indentation curves for hydrogel surfaces

810

(Figure S9), ). In vitro assessment with neural model SH-SY5Y cell line (Figure S10), and

811

experimental conditions for optimizing the fabrication of hydrogel layers (Table S1).

812 813

ACKNOWLEDGEMENT

814

This work was supported by the M2Neural project (http://www.m2neural.eu), funded in the

815

FP7 M-ERA.NET Transnational framework. In addition, this work was also supported by the

816

Slovak Research and Development Agency under contract numbers APVV-14-0858 as well

817

as the VEGA Grant Agency under the contract no. 2/0059/16.

818


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Soft hydrogel zwitterionic coatings minimize fibroblast and

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