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Biomaterials 67 (2015) 73e83

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Biomaterials

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Synthetic 3D diamond-based electrodes for flexible retinalneuroprostheses: Model, production and in vivo biocompatibility

Amel Bendali a, b, c, Lionel Rousseau d, Ga€elle Lissorgues d, Emmanuel Scorsone e,Milan Djilas a, b, c, Julie D�egardin a, b, c, Elisabeth Dubus a, b, c, St�ephane Fouquet a, b, c,Ryad Benosman a, b, c, Philippe Bergonzo e, Jos�e-Alain Sahel a, b, c, f, g, h, Serge Picaud a, b, c, *

a INSERM, U968, Institut de la Vision, Paris F-75012, Franceb Sorbonne Universit�es, UPMC Univ Paris 06, UMR_S968, Institut de la Vision, Paris F-75012, Francec CNRS UMR 7210, Institut de la Vision, Paris 75012, Franced ESIEE e ESYCOM Universit�e Paris Est, Cit�e Descartes, BP99, 93162 Noisy Le Grand, Francee CEA-LIST, Diamond Sensors Laboratory, Gif-sur-Yvette, Francef Fondation Ophtalmologique Adolphe de Rothschild, Paris, Franceg CHNO des Quinze-Vingts, Paris, Franceh Academie des Sciences, Paris, France

a r t i c l e i n f o

Article history:Received 6 January 2015Received in revised form9 July 2015Accepted 11 July 2015Available online 14 July 2015

Keywords:Retinal prostheses3D electrodeResolutionDiamondBipolar cellGliosis

* Corresponding author. Institut de la Vision, 17 rueE-mail address: serge.picaud@inserm.fr (S. Picaud

http://dx.doi.org/10.1016/j.biomaterials.2015.07.0180142-9612/© 2015 Published by Elsevier Ltd.

a b s t r a c t

Two retinal implants have recently received the CE mark and one has obtained FDA approval for therestoration of useful vision in blind patients. Since the spatial resolution of current vision prostheses isnot sufficient for most patients to detect faces or perform activities of daily living, more electrodes withless crosstalk are needed to transfer complex images to the retina. In this study, we modelled planar andthree-dimensional (3D) implants with a distant ground or a ground grid, to demonstrate greater spatialresolution with 3D structures. Using such flexible 3D implant prototypes, we showed that the degen-erated retina could mould itself to the inside of the wells, thereby isolating bipolar neurons for specific,independent stimulation. To investigate the in vivo biocompatibility of diamond as an electrode or anisolating material, we developed a procedure for depositing diamond onto flexible 3D retinal implants.Taking polyimide 3D implants as a reference, we compared the number of neurones integrating the 3Ddiamond structures and their ratio to the numbers of all cells, including glial cells. Bipolar neurones wereincreased whereas there was no increase even a decrease in the total cell number. SEM examinations ofimplants confirmed the stability of the diamond after its implantation in vivo. This study further dem-onstrates the potential of 3D designs for increasing the resolution of retinal implants and validates thesafety of diamond materials for retinal implants and neuroprostheses in general.

© 2015 Published by Elsevier Ltd.

1. Introduction

Visual prostheses aim to provide blind patients with useful vi-sual information for face and object recognition, as well as thereading of text and orientation in unknown environment. Despitethe degeneration of bipolar cells and retinal ganglion cells, theelectrical stimulation of retinal implants has been shown, in clinicaltrials, to be safe, to enablemost blind patients to identify contrastedobjects, to follow a line or the ground, and, in some cases, to read

Moreau, 75012 Paris, France.).

short words [1e4]. The Argus II device (2nd Sight) has obtained theCEmark and FDA approval, and the Alpha-IMS (Retinal implant AG)has received the CE mark. Preclinical studies are currently evalu-ating photovoltaic silicon materials [5] or photosensitive polymers[6,7]. Different clinical trials have also demonstrated the ability ofsuprachoroidal prostheses to activate the degenerated retina wheninserted in the space between the sclera and the choroid [8,9]. Forpatients with retinal ganglion cell degeneration in retinal diseasessuch as glaucoma or diabetic retinopathy, Brindley and his co-workers have pioneered vision prostheses for a direct activation ofthe visual cortex [10]. These cortical implants have also allowedpatients to recover partial vision, but this visual recovery appears tobe transient [11]. Finally, psychophysical experiments have

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indicated that complex visual tasks, such as text reading, orienta-tion in unknown environment or face recognition, would require atleast 600 independent pixels [12e14].

The major challenge in visual rehabilitation with neuro-prostheses is therefore to increase electrode density whilstincreasing the spatial resolution of each electrode, such that eachindividual electrode generates a pixel. Current retinal prosthesesfunction in a monopolar mode with a distant returning ground, aconfiguration, for which the spatial distribution of current wereinvestigated electrophysiologically on the chicken retina [15].However, different electrode configurations were recentlydescribed to increase the electrode resolution. For instance, currentdiffusion can be limited by local return electrodes as in bipolarstimulations using a circular electrode around the stimulatingelectrode [16]. A quasimonopolar stimulation was also reported toincrease the resolution by using a distant return electrode in aplane above hexapolar return electrodes surrounding each stimu-lating electrode [17]. If the combination of the hexapolar andmonopolar stimulations can improve the containment of the acti-vated sites, it increases the threshold level of activation due to theshunting of currents to local return electrodes [17]. In addition, thequasimonopolar stimulations requires complex current injectionsat each of the hexapolar electrodes [18]. More recently, ground gridwith a high conductivity were found to provide a greater focaliza-tion of currents [16]. The ground grid configuration should bepreferred for high-density arrays because bipolar and quasimono-polar stimulations would increase the number of connecting wires.3D implant geometries are also thought to improve electrode res-olution in the bipolar or ground grid configurations by locallymoving neurons between the stimulating and return electrodes[19e21]. The success of such 3D designs implies that the residualblind retina remains sufficiently plastic to mould itself around the3D structures. This preservation of the flexibility of the residualblind retina was suggested from studies of pillars penetrating thetissue or cavities to be filled with cells [22]. Neurons were foundaround pillars and in cavities or wells only if the opening was largerthan 20 mm across [19,20].

Increases in electrode density require a decrease in electrodesize and, thus, an increase in the charge density to be injected toachieve neuronal activity. This constraint has driven the search fornew materials with greater developed surfaces, such as blackplatinum or iridium oxide [23]. Materials with a broader electro-chemical potential window are also being tested to ensure that thesafe charge injection limit is not exceeded. One such material,diamond, is considered particularly attractive, as it displays thebroadest electrochemical window of any semiconductor provided itis doped with nitrogen or trimethyl boron [24,25]. Nanocrystallinediamond can even be deposited on 3D structures, making itpossible to synthesise materials with high aspect ratios anddeveloped structures [26,27]. Diamond electrodes have even beenshown to activate retinal neurons [28]. Finally, diamond has beenshown to display biocompatibility in vitro with embryonic corticalneurons and stem cells [29e32] and even retinal neurons [33].However, this biocompatibility of diamond in vitro does notnecessarily imply that it would be biocompatible in the long termin vivo, as other biocompatible materials have been shown toinduce retinal gliosis, or even fibrosis, in vivo [19]. Gliosis is clas-sically characterized by the multiplication of glial cells and theirconsecutive hypertrophy while fibrosis was defined in the abovestudy as the formation of a fibrous preretinal membrane, bothcellular events resulting in the spacing between retinal neuronsand the implant likely to hamper neuronal stimulation.

In this study, we first investigated retinal currents in different3D electrode configurations for image encoding. Because the re-sults of this modelling study are valid only if neurones integrate the

3D structures, we then produced 3D soft implant to examine thetissue implant interfacing. However, this first study was notintended to validate the 3D modelling by physiological measure-ments. The fabrication procedure was developed to allow thecoating of our 3D implants with diamond, for assessing the in vivobiocompatibility of this material. A specific imaging procedure wasalso used to preserve the tissue/implant interface, making itpossible to assess the biocompatibility of diamond in vivo correctly.

2. Materials & methods

2.1. Modelling

We created finite-element models of four variants of a 25 � 25stimulating electrode array within a retinal prosthetic system: (i) aplanar array with a common counter electrode in the shape of agrid surrounding the stimulating electrodes, (ii) a 3D electrodearray with the stimulation electrodes surrounded by a counterelectrode grid, again serving as a counter electrode, (iii) a planarelectrode array with a distant counter electrode, and (iv) a 3Delectrode array with a distant counter electrode. The 3D wells wereshaped as inverted pyramids with their points cut off at 30 mmheight (well depth). The well opening edge was 72 mm and the wellbottom edgewas 36 mm. The stimulation electrodewas set to be theentire bottom surface, i.e. square with 36 mm edge. The electrodedimensions and shape for the planar electrode were the same as inthe 3D array. The inter-electrode distance was 100 mm. The elec-trical conductivity of the tissue was 0.25 S/m, as in a previous study[20].

For simulation purposes, an image of Abraham Lincoln wascropped and sampled to obtain a square 25 � 25 image, the colourpalette of which was then reduced from 256 to three levels: white,grey, and black (Fig. 1A and B). The resulting image was thenmapped into a finite-element model, by assigning current densitiesto the stimulating electrodes proportional to the grey levels in thecropped image: zero for white, 1000 A/m2 for grey and 2000 A/m2

for black. If we assume a stimulation pulse width of 1 ms, theestimated charge densities would be 0.1 mC/cm2 for grey and0.2 mC/cm2 for the black intensity levels. These values are belowthe reported safety limit for platinum (0.35e0.4 mC/cm2) and wellbelow the limit for iridium-oxide (3e4 mC/cm2), and also belowthe typical levels reported in previous studies [34].

2.2. Microfabrication and SEM

Silicon moulds were prepared using KOH wet etching togenerate a 3D structure. A sacrificial oxide layer was then generatedover the silicon mould by the thermal oxidation of the wafer in afurnace at 1100 �C in the presence of oxygen and hydrogen, until a1 mm layer of silicon dioxide was achieved. Pt electrodes wereproduced by patterning sputtered Pt over the silicon moulds bystandard photolithography. The substrate was then spin-coatedwith polyimide (PI 2611) to obtain a 10 mm-thick layer of poly-mer. The polyimide was then cured at 450 �C under nitrogen flowfor 6 h, and a 500 nm-thick aluminium film was sputtered over it.We then spin-coated AZ4562 (Clariant, Muttenz, Switzerland) thickphotoresist onto the wafer to define the shape of the implant. Afterthe development step, the wafer was placed in Cl2 plasma forreactive ion etching (RIE) of the aluminium layer. The unmaskedpolymer was etched away with O2 RIE to achieve the final shape ofthe implant. The aluminium masking was then removed by wetetching. The wafer was immersed in hydrofluoric acid (HF) to etchthe sacrificial oxide layer and release the individual implants.Finally the implants were rinsed in DI water and dried.

The diamond-based implants were produced as follows.

Fig. 1. Grey-scale image of Abraham Lincoln with 256 grey-scale levels before (A) and after down sampling to 25 � 25 pixels and reducing the colour palette to 3 grey-scale levels(B). The brightness ranges covered by the 3 intensity levels are shown on the colour bar on the right. Current density profiles along the red line, with the image encoded on a25 � 25 electrode array, are shown in the next figure. The current densities above the cathode are represented in these four configurations: i) planar electrode array with a returningground grid (C), ii) 3D electrode array with a returning ground grid (D), iii) planar electrode array with a distant ground (E), iv) 3D electrode array with a distant ground (F). (Forinterpretation of the references to colour in this figure legend, the reader is referred to the web version of this article.)

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Diamond was selectively grown in silicon moulds as described byBongrain and coworkers [35]. A microwave plasma enhancedchemical vapour deposition (MPECVD) reactor (Seki AX6500) wasused to synthesize diamond in a mixture of methane (CH4) andhydrogen (H2) gases at a microwave power of 3 kW, a gas pressure25mbar, and a substrate temperature of about 800 �C. The diamondlayer obtained was about 300 nm thick. As for the platinum elec-trodes, the substrate was then covered with a polyimide film andthe process used to define the histological implants was identical tothat used for Pt-based implants. Polyimide 3D implants weregenerated with the same procedure without any previous diamondgrowth.

The implants were imaged with a ZEISS Supra-40 field emissionscanning electron microscope (SEM) operating at an accelerationvoltage of 2 kV. The implants were imaged by SEM after the im-plantation period. The retinas, fixed together with the implants inparaformaldehyde (see below), were peeled off the implant and theimplant was dehydrated in a series of alcohol concentrations (50%,70%, 90% and 100% ethanol).

2.3. In vivo studies

Homozygous P23H rats (line 1, kindly provided by Dr Lavail)were housed with a 12 h dark/12 h light cycle, with food and wateravailable ad libitum. All experiments were carried out in accordancewith European Community Council Directives (86/609/EEC) andwith the ARVO (Association for Research in Vision and Ophthal-mology) statement for the use of animals in ophthalmic and visualresearch. Animals were sacrificed by CO2 sedation and cervicaldislocation, and all efforts were made to minimize suffering. Thesurgical procedure used to implant the prototypes has beendescribed in detail elsewhere [36]. Briefly, P23H blind rats wereanesthetized by the intraperitoneal injection of a 4:1 mixture ofketamine-xylazine (ketamine 100 mg kg�1, xylazine 10 mg kg�1;Ketamine 500: Virbac, Carros, France; xylazine 2%: Rompun®, BayerPharma, Puteaux, France). A small radial sclerotomy (1.5 mm long)was performed behind the limbus with a slit knife. Viscoat®

Intraocular Viscoelastic Injection (Alcon Laboratories, Hünenberg,Switzerland) was injected into the subretinal space through the

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sclerotomy, with a 27G cannula, to obtain localised retinaldetachment. The implant was then inserted into the subretinalspace. Immediately after surgery, the correct positioning of theimplant was checked in vivo by indirect ophthalmoscopy (frost andlens). In vivo imaging was performed one week after surgery andthen again six weeks later, right before the sacrifice, for observationof the eye fundus by endoscopy. A Micron III digital endoscope(Phoenix Research Laboratories, Pleasanton, California) was usedfor imaging of the eye fundus, together with StreamPix V softwareand a rat probe.

2.4. Immunostaining, confocal imaging and quantification

After six weeks, animals were sacrificed by CO2 sedation andcervical elongation. The eyes were removed and placed inphosphate-buffered saline (PBS, 0.1 M, pH 7.4). They were dissectedso as to retain only the tissue fragment containing the implant. Thisfragment was fixed by incubation overnight at 4 �C in para-formaldehyde in PBS (4% wt/vol) and then rinsed in PBS.

For immunolabelling, retinal fragments were incubated in ablocking solution (10% bovine serum albumin (Sigma, St. Louis,Missouri), 1% Triton X-100 (Sigma), 0.5% Tween 20 (Sigma) and0.1 g/l Thimerosal (Sigma) in PBS) for 1 h at room temperature. Theywere then incubated for 2 day at room temperature with primaryantibodies in blocking solution (dilution 1:2). The antibodies usedwere a polyclonal antibody directed against rabbit PKCa (C-20)(1:1000, Santa Cruz Biotechnology, Dallas, Texas) and a monoclonalantibody directed against mouse Goa (1:200, Chemicon, Darm-stadt, Germany). The fragments were rinsed and then incubatedwith secondary antibodies: goat anti-mouse IgG and goat anti-rabbit IgG antibodies conjugated to Alexa™594 and Alexa™488,respectively (1:500, Molecular Probes, Invitrogen, Eugene, Oregon)for one day. Cell nuclei were stained with 40,6-diamidino-2-phenylindole (DAPI), which was added during the final incuba-tion period. The implant/retina ensemble was then rinsed andmounted, in permanent mounting medium (MMFrance), on a mi-croscope slide, for viewing under an upright confocal microscope.

Confocal microscopy was performed on an Olympus FV1000laser-scanning confocal microscope. DAPI counterstaining,AlexaFluor-488 and AlexaFluor-594 and AlexaFluor-647 weredetected by excitationwith a 405 nm laser diode, a 488 nmargon ionlaser, and 559 nm and 635 nm laser diode lines, respectively. Theselection of excitation and emission wavelengths was controlled byappropriate filters: a dichroic mirror (405/488/559/635), SDM490,SDM560, and SDM640 emission beamsplitters and BA430-470,BA505-540, BA575-675 and BA655-755 barrier filters. The primaryobjective used was an Olympus oil immersion UPLSAPO 20X NA0.85-WD 0.20 or UPLFLN 40X NA1.30-WD 0.20 objective. The mi-croscope and image acquisition were controlled with OlympusFluoview software version 4.1. Images were acquired at a resolutionof 1024 � 1024 pixels, with a scan rate of 10 ms pixel�1, with nozoom (20� related pixel size: 0.621 mm, 40� related pixel size:0.310 mm). Images were acquired sequentially, line-by-line, tominimise the crosstalk between excitation and emission, with a stepsize defined according to the Nyquist-Shannon sampling theorem.Exposure settings minimising the number of oversaturated pixels in

Fig. 2. Amplitude of current densities for different grey level stimulations. Current densityelectrode array, on which the Lincoln image has been encoded with 3 intensity levels: no curabove the planar electrode array with a distant counter electrode; B: 20 mm above the threewith the counter electrode surrounding the wells; D: 20 mm above the three-dimensionalelectrode array with a distant counter electrode; F: 40 mm above the three-dimensional arrelectrode surrounding the wells; H: 40 mm above the three-dimensional array with the coupixels of the Lincoln image with the different configurations at 20 mm (I) and 40 mm (J) abovelegend, the reader is referred to the web version of this article.)

the final images were used. Twelve-bit images were then processedwith ImageJ or FIJI and converted into 24-bit RGB colour mode. Theimages were then edited with Adobe Photoshop CS5 software andassembled with Adobe Illustrator CS. The presence of bipolar cellswithin the wells of 3D-structured implants was assessed by deter-mining the ratio of bipolar cells to the total number of cell nuclei perwell for each material: each Z-section was preprocessed to retainonly the staining located in each well, and the numbers of bipolarcells and of total cells (DAPI-positive counterstaining) were deter-mined. Bipolar cells were countedwith the ImageJ cell counter plug-in and total nuclei were counted semi-automatically with Imarissoftware (Bitplane AG, Zurich, Switzerland).

2.5. Statistical analysis

We present results for three polyimide implants, five diamondimplants and three platinum implants, based on the values of fourto nine wells per implant for quantification. All data are expressedas means ± SEM. The Gaussian distribution of the raw data wastested with a ShapiroeWilk normality test. One-way ANOVA wasthen carried out, followed in cases of significance by either a Bon-ferroni post-hoc test (Gaussian distribution) or Dunns post-hoc test(non-Gaussian distribution), to compare means between groups.Differences were considered significant if *p < 0.05, **p < 0.01 and***p < 0.001.

3. Results

3.1. Models of implant designs

Ground grids and 3D electrode designs have been reported toimprove the electrical stimulation of retinal tissues [16,19e21].Single-electrode models were thus generated to demonstrate theadvantage of either a ground grid on a planar substrate [16] or of a3D well with a ground grid [20]. However, these models were notused to examine the distribution of current in a 3D structure with adistant ground. They were also not used to investigate the distri-bution of current on an electrode array for image representation.Instead, we examined how a face would be encoded on suchelectrode arrays. The face of Abraham Lincoln was encoded with apalette of three grey-scale levels (Fig. 1A and B), converted intothree current intensities. Finite-element modelling was used tosimulate the current density distribution in the retinal tissue abovethe electrode arrays. Fig. 1 illustrates the current densities 40 mmabove the cathode for the four configurations considered: i) planarelectrode array with a returning ground grid (Fig. 1C), ii) 3D elec-trode array with a returning ground grid (D), iii) planar electrodearraywith a distant ground (E), iv) 3D electrode array with a distantground (F). When calculated on a line running 20 mm above thestimulating cathodes (Fig. 2AeD), the current densities presentedsquare curves above active electrodes within the 3D structures,with or without a ground grid, whereas they yielded peaks aboveactive electrodes on the flat arrays. The introduction of a groundgrid (Fig. 2C and D) suppressed the relatively high currentmeasured above inactive electrodes in conditions with a distantground (Fig. 2A and B), with this baseline current level increasing

plots along the red line from the previous figure, 20 mm and 40 mm above the 25 � 25rent injected for white pixels, 1000 A/cm2 for grey, and 2000 A/cm2 for black. A: 20 mm-dimensional array with a distant counter-electrode; C: 20 mm above the planar arrayarray with the counter electrode surrounding the wells; E: 40 mm above the planar

ay with a distant counter electrode; G: 40 mm above the planar array with the counternter electrode surrounding the wells. I,J) Quantification of current densities for all thethe electrodes (Mean ± SD). (For interpretation of the references to colour in this figure

Fig. 3. Production of flexible diamond implants. A: Diagram of the microfabricationsteps with nano-diamond seeding and selective diamond growth followed by poly-imide addition; B: Picture of the mask (Yellow: KOH 3D structures, grey shape ofimplant); C: SEM picture of the silicon mould; D: Final dummy implant for in vivoevaluation. (For interpretation of the references to colour in this figure legend, thereader is referred to the web version of this article.)

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towards the ground. Quantification of the current densities at a20 mm distance from all electrodes confirmed this large baselinecurrent in configurations with a distant ground (Fig. 2I). As aconsequence, the differences of current densities between posi-tions above active and inactive electrodes are greater in arrays witha ground grid than those generated by the corresponding arraywith a distant ground. The greatest differences are produced by the3D array with a ground grid. However, the 3D array with a distantground is in a similar range or even better than the flat array with aground grid. The worst case is the flat array with a distant ground,this configuration showing high variability in elicited currentdensities above inactive and stimulated electrodes limiting therebythe distinction between white and grey levels (Fig. 2I). At a greaterdistance from the array (40 mm above the electrodes), the resultsshowed a great reductions of current densities except for the flatarray with a distant ground. For the intermediate grey-level stim-ulations, 3D arrays still show clear peaks of current densities aboveelectrodes (Fig. 2F and H), which are less distinguishable withplanar arrays (Fig. 2E and G). However, the quantification of currentdensities indicated similar differences between gray levels exceptfor the planar array with a ground grid. Again, the planar array witha distant ground exhibits a greater variability in each group limitingthereby the distinctions between grey levels (Fig. 2J). The effect ofplacing the distant counter electrode eccentrically above the lowerright corner of the Lincoln image (rather than above its centre) canbe seen on the plots with distant ground configurations (Fig. 2AeB,EeF). In such instances, the baseline current density increases withdecreasing distance from the counter electrode, as all the returncharge from all pixels converges on the ground. No such effect isseen for configurations with a grid ground electrode, for which allplots have a constant baseline (Fig. 2CeD, GeH). These above ad-vantages of 3D implants in neuronal stimulations justify the need toassess innovative materials on such 3D structures. However, theseadvantages are expected provided neurones to be stimulatedintegrate into 3D structures. Therefore, to assess the biocompati-bility of diamond and assess neuronal integration in the 3D im-plants, we developed a fabrication process to generate diamondelectrodes on a 3D flexible implant.

3.2. Production of a diamond-coated 3D flexible foil

The ability to produce flexible substrates conforming to thecurvature of retinal tissues appears to be essential for the mainte-nance of a correct tissue interface. However, the classical synthesistechniques used to grow diamond are based on high temperaturesand microwave plasma techniques that cannot be applied tobiocompatible soft substrates. We therefore developed a new so-lution based on a peel-off process, in which the soft polyimidepolymer was deposited on top of a 3D patterned diamond layer. Wefirst generated the 3D structures by preparing silicon moulds byKOH wet etching, to generate truncated pyramids. These pyramidswere obtained by adding a structure to compensate for the etchingspeed of the 110 and 100 oriented crystalline planes. The processwas stopped when the pyramids on the silicon moulds had typi-cally attained a height of 30 mm (Fig. 3). Diamond was then grownon these 3D silicon moulds, as follows: 1) seeding of the siliconmould with nano-diamond particles (approximately 5 nm indiameter), 2) sputtering and patterning of an aluminium mask onthe silicon mould by photolithography, 3) etching away of the un-protected nano-diamond particles by reactive ion etching (RIE)under oxygen plasma, 4) removal of the aluminium by wet etching,5) growth of a diamond layer (approximately 300 mm thick) aroundthe nano-diamond particles in a microwave plasma enhancedchemical vapour deposition (MPECVD) reactor. The diamond layerwas spin-coated with a 10 mm-thick layer of polyimide (PI 2611),

whichwas then cured. Classical polymer etchingwas used to definethe shape of the implant. Finally, removal of the sacrificial oxidelayer led to the release of individual diamond-coated 3D implants.

3.3. Biocompatibility

For investigation of the biocompatibility of diamond in vivo, softpolyimide implants with or without diamond or platinum coatingswere inserted into the subretinal space of P23H rats, an animalmodel of retinitis pigmentosa at an age at which the photoreceptorshave degenerated. The correct insertion of the implants waschecked in vivo using a Micron III numerical endoscope. Images ofthe eye fundus are provided for the various implants in Fig. 4:purely polyimide, metallic, and diamond-coated. The presence ofretinal blood vessels above the devices confirmed their subretinalpositions. This examination also made it possible to visualise thedisappearance of the subretinal bleb generated for introduction ofthe subretinal implant.

In our investigations of diamond biocompatibility, we had toexamine the retinal tissue in the vicinity of the implant. Classically,such examinations are carried out on semi-thin sections or cryostatsections, on which cells can be identified by immunostaining[19e21,36]. This approach is entirely feasible with soft materialdummies (e.g. polyimide) [37], but it is very difficult to cut pro-totypes containing other hard materials, such as diamond. Implantremoval is not a viable option, because the 3D structure enhancingthe tissue/interface would complicate the operation. Instead, wedeveloped an innovative approach based on direct confocal imag-ing of the implant/tissue eye cup whole mounts. We assessedbiocompatibility in vivo by visualising cell nuclei and ON bipolar

Fig. 4. Eye fundus of P23H rats with implanted polyimide (A), diamond (B) and metallic (C) devices. The scale bar is 500 mm.

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cell neurons in the 3D wells on whole-mount preparations. Theimmunostaining protocol was adapted to preserve the implant/tissue interaction while allowing the antibody to diffuse over adistance of 100 mm within the retinal tissue (see methods). Fig. 5illustrates such confocal images of the tissue/implant interface fora diamond implant, along views corresponding to different z stacks(view “a”: top of the cavities, view “b”: bottom of the cavities),shown both with top views (AeF) and orthogonal views (GeH). Cellnuclei were labelled with DAPI (blue) and ON bipolar neurons wereimmunolabelled with Goa antibodies (green). The �40 magnifica-tion of the retina/implant whole mount makes it possible to visu-alise the DAPI-stained nuclei in all four cavities. Both theorthogonal views (xez axes) and the vertical retinal sections show

Fig. 5. Confocal imaging of stained retinae in contact with implants, 2 views. AeC: top views“b” (bottom of the cavities); GeH: orthogonal views indicating the points fromwhich viewsstained with anti-Goalpha antibody; C, F: coloured merged images of both DAPI and Goalphcolour in this figure legend, the reader is referred to the web version of this article.)

that retinal bipolar neurons fill the entire cavity, right down to thebottom (H). These data demonstrate that the residual retina isplastic enough to mould itself into the 3D implant wells.

We assessed the biocompatibility of the materials by quanti-fying cell occupancy in the cavities (Fig. 6). This quantification wasobtained by generating the 3D reconstruction of the content foreach individual well of an implant as illustrated in Fig. 6(AeC). Ourstrategy for assessing the material biocompatibility has been toquantify immunolabelled ON bipolar cells to demonstrate the sur-vival of these neurons targeted by subretinal electrical stimulations.However, to define if the material triggered reactive gliosis, we firstcalculated the total number of cell nuclei because glial cell prolif-eration would be expected to increase their number and thus to

of whole-mount retinae along view “a” (top of the cavities); DeF: top views along view“a” and “b” were taken. A, D, G: DAPI staining of all cell nuclei; B, E, H: ON bipolar cellsa staining (DAPI in blue and Goalpha in green). (For interpretation of the references to

Fig. 6. Quantification of bipolar cells within the 3D electrodes for each material. AeC: Image processing for cell counting with preprocessing (A), Imaris nucleus counting (B) andmanual bipolar cell counting (C); DeI: plots of cell numbers for each cavity; D: number of cell nuclei per volume; E: mean number of cell nuclei per volume; F: proportion of bipolarcells, expressed as a percentage of the total number of cell nuclei; G: mean value of the ratio of bipolar cells to total nuclei. H: number of bipolar cells per volume; E: mean number ofbipolar cells per volume

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decrease the ratio between neuronal numbers to all cell numbers.The quantification of all cell nuclei was achieved by defining thefluorescent spheres corresponding to DAPI nuclear staining. Thisquantification indicated that the density of cell nuclei in theimplant cavities was greater for polyimide implants(1.50 � 10�3 ± 0.023 � 10�3 cells/mm3) than for diamond-(1.10 � 10�3 ± 0.079 � 10�3 cells/mm3) or platinum-coated(1.21 � 10�3 ± 0.189 � 10�3 cells/mm3) implants. These cells caneither be neurones of the inner retina (Bipolar cells, horizontal cells,amacrines cells) or glial cells (Müller macroglial cells, microglialcells). Because subretinal implants are intended to depolarize ONbipolar cells, we quantified these neurones in the 3D wells

following their immunolabelling. The quantification demonstrateda stability of the ON bipolar cell densities for the different implantsexcept for a platinum-coated implant (Fig. 6H and I). Finally, to getan estimation of retinal gliosis, we calculated the ratio of bipolarcells to all cell nuclei. This ratio was greater with diamond-coatedimplants (39.4 ± 2.3%) than with the polyimide (28.9 ± 1.2%) orplatinum-coated implants (26.7 ± 3.5%) (Fig. 6). A lack of biocom-patibility is expected to induce neuronal degeneration and anassociated reactive gliosis with a proliferation of glial cells, whichwould thus result in a decrease in the neurone to glial cell ratio.Therefore, the higher ratio of bipolar cells to all cells in the diamondwells is consistent with a greater biocompatibility of diamond than

Fig. 7. Examination of the materials used by scanning electron microscopy, following in vivo implantation: after fabrication during which the polyimide layer is lifted off thestructure shown in Fig. 3C, the 20 mm-thick foils were surgically implanted in rats for 8 weeks, then explanted and prepared for SEM observations. The pictures display the nakedpolyimide surface (A, D), and the same covered by a thin diamond layer (B, E), or metal (C, F). Although the numerous processes that significantly altered the edges of the thinpolyimide foils, the images display that the surface qualities remained unchanged during the implantation period. For all these implants, residual cells or tissues are visible withinthe cavities on the enlarged views (DeF).

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polyimide alone or platinum. The absence of a massive inflamma-tory reaction and the presence of many bipolar neurons in thewellssuggest that the various materials used, including diamond inparticular, are not toxic to retinal neurons.

3.4. Characterization of diamond implants

The original process used here made it possible to produce softimplants with several 3D wells, which were either left uncoated orwere coated with either diamond or platinum. The diamondcoating covered the entire area of the implant visible on thephotograph in Fig. 7B and E, including the walls and the bottom ofthe cavities, whereas in the case of platinum the metal coverageappears in light grey colour in panels C and F. Following in vivoimplantation, the surfaces of the implants were observed byscanning electron microscopy (SEM), to assess the physical stability

of the implants (results for all three materials tested are shown inFig. 7). On the polyimide implant (A, D), the surface of the materialappears similar to that of the freshly produced implants, with novisible defect. Note that thewhite traces visible on Fig. 7A are due tocharge accumulation on this insulating surface during SEM imag-ing. Similarly, the diamond films (B, E) showed no discontinuitiesand the surface was correctly covered. Nevertheless, the very finecracks observed at the edges of the diamond wells (already presentbefore implantation) indicate that the deposition and growth of thematerial could be optimised further. Unlike conventional poly-crystalline diamond, the diamond surfaces appeared very smooth.This smoothness was a consequence of the process used, with theexposed side of the diamond originally in contact with the siliconsurface. Finally, on the metallic implants (C, F), the light grey areascorresponding to the platinum coating also appeared to be free ofsignificant defects and darker due to the presence of organic matter

A. Bendali et al. / Biomaterials 67 (2015) 73e8382

(residual cells). Thus, neither the diamond nor platinum surfaceswere damaged by implantation despite the long process fromsurgery to the cleaning for SEM examination including the immu-nolabelling and flat mount observation. For all these implants,some cells or tissue remained visible on the implants, particularlywithin the cavities, as observed on the enlarged views (DeF). Thisobservation confirmed the deep integration of the tissue into the3D implants, regardless of the material used.

4. Discussion and conclusions

Previous psychophysical studies have demonstrated thatretinal implants can allow face recognition, independent loco-motion and text reading if they generate at least 600 independentpixels [12e14]. This requires independent stimulation by the in-dividual electrodes of an implant. Current retinal implants arebased on classical monopolar stimulation between a stimulatingelectrode and a distant ground, but other configurations havebeen proposed, to increase the resolution of individual electrodes.These other configurations include bipolar stimulation betweentwo neighbouring or concentric electrodes, the quasimonopolaror the introduction of a local returning ground grid [16,17].Ground grids have already been introduced into some of theplanar subretinal implants currently undergoing preclinicaltesting [5]. In this study, we confirmed that local ground gridswere able to decrease current densities in areas surroundingstimulated zones. As previously described [16], we confirmed thata ground grid on a planar array can decrease the current densitiesabove non-stimulated areas. However, we show further that theground grid on a planar array also decreases the current densitiesjust above stimulated areas requiring therefore higher injectedcurrents to reach an activation threshold. This conclusion is notvalid for 3D arrays with a ground grid at short distances (20 mm)but becomes tru at greater distance. The use of 3D electrode de-signs has also been proposed as a means of increasing the reso-lution of electrical stimulation by restricting the electrical fieldwithin cavities between bipolar electrodes [20e22]. Weconfirmed that 3D configurations increased the local resolution ofstimulations with very high current densities within the 3D well.Surprisingly, we also found that, even with a distant ground, 3Dconfigurations also generated very high current densities withinthe well whilst yielding lower current densities in unstimulatedareas than for the planar configuration.

However, the use of a 3D structure is advantageous only if theneuronal tissue remains sufficiently plastic to mould itself ontothe 3D structure, to place neurons between the electrodes. Pal-anker and his group have shown that the cavities in 3D structurescan fill with cell bodies and neurons, depending on the size of thecavity opening [19]. The production of 3D implants should make itpossible to position neurons between two electrodes of opposingpolarities [21]. The retina has even been shown to mould aroundpillars [19]. However, given that it may be necessary to removeretinal implants, we preferred well shapes over pillar structures[20]. As previously described in RCS rats [19,21], we confirmed inP23H rats, another rodent model of retinitis pigmentosa [38], thatthe degenerated retina can mould around 3D structures. In apreliminary study, we reported such integration for some retinalsections, but tissue sectioning disrupted the tissue/implantinterface [20]. It was therefore not possible to characterize thetissue within the well correctly and, therefore to quantify the bi-polar cell neurons present in this volume. In this study, weshowed, by imaging retinal whole mounts, that the tissue was inintimate contact with the structures tested: polyimide, platinum,diamond. We reconstituted the contents of the well and quanti-fied the bipolar neurons present in this small volume. The

presence of many bipolar cells in such a well demonstrates thefeasibility of activating a retinal column independently of theneighbouring retinal columns in other wells. Our 3D design forsubretinal implants would therefore allow the production of in-dependent pixels for each electrode. Electrode impedance wererecently reported for chronic implantation [39]. Further studiesare required to determine whether similar chronic implantationsof 3D designs really do increase the resolution of individualstimulations. Such 3D implant arrays of electrodes could be acti-vated by an ASIC either tethered by wires as in the Argus II implant[40] or connected on the backside of the implant as in the sub-retinal electronic implant alpha-IMS [1]. However, the productionof photosensitive electrode arrays could also solve the difficultissue of tethering the 3D implant to an ASIC [5]. The use of flexiblephotosensitive polymers would be an additional advantage toenhance the implant/tissue interface [6,7].

Close proximity between tissue and electrode is required forprosthetic applications, as this decreases the diffusion of stimu-lating currents and increases their geometric localisation. Suchtight interactions are required particularly for the development ofneuroprostheses, accounting for current interest in the develop-ment of novel biocompatible materials with good electronicproperties. In addition, two types of current can be generated byelectrodes for the electrical stimulation of a neuronal structure: 1)faradic currents, involving chemical oxidation/reduction reactions;2) capacitive currents, resulting purely from charge accumulation.In neuronal prostheses, capacitive stimulation is favoured, as itlimits pH variation at the surface of the electrode [23]. In thecontext of visual implants, the need to use small electrodes with ahigh resolution introduces a supplementary risk of tissue andelectrode degradation, resulting from the injection of larger chargedensities. It is therefore essential to use materials with high chargeinjection limits, such as diamond. There has been considerable in-terest in the use of this carbon-based material for such applicationsin recent years. Once doped with boron, diamond has excellentelectronic and electrochemical properties and is chemically andmechanically inertia [41]. The potential of such diamond electrodesfor the stimulation of retinal neurons has been demonstrated inacute implantation experiments [28]. The biocompatibility of dia-mond was first demonstrated with embryonic cortical neurons andstem cells [29e32] and even retinal neurons [33]. Our resultsfurther demonstrate the in vivo biocompatibility of diamond onflexible implants.

This study confirms that 3D-structured electrodes are advan-tageous in the design of retinal implants, as they greatly increasethe resolution of stimulation. Our findings also highlight theconsiderable benefits of diamond as an attractive electrode mate-rial for neuroprostheses.

Acknowledgements

We thank Dr Matthew Lavail for providing the rat P23H line.AB received a doctoral fellowship from UPMC. This work wassupported by INSERM, UPMC (Paris VI), Foundation FightingBlindness, the F�ed�eration des Aveugles de France, Fondation de laRecherche M�edicale (grant number DBC20101021013) to SP, IRRP,the city of Paris, the Regional Council of Ile-de-France, the FrenchAgence Nationale de la Recherche (ANR) with the MEDINASproject (ANR07TECSAN014), the IMPLANTS project from ITMO-AVIESAN, the European Community's Seventh Framework Pro-gramme (FP7/2007-2013) under grant agreement no. 280433(Neurocare project), and the LabEx LIFESENSES (ANR-10-LABX-65), which was supported by French state funds managed by theANR within the Investissements d'Avenir programme (ANR-11-IDEX-0004-02).

A. Bendali et al. / Biomaterials 67 (2015) 73e83 83

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