# Steerable Thin-Film Electrode Array for Cochlear Implantation: Design and Development for Future Atraumatic Insertion

Ahmad Itawi, Bacem Zribi, Prabhakar Sidambaram, Thanh Hang Tran, Guillaume Tourrel, Renato Torres, Sofiane Ghenna, Sébastien Grondel, Yann Nguyen, Cedric Plesse, et al.

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## ► To cite this version:

Ahmad Itawi, Bacem Zribi, Prabhakar Sidambaram, Thanh Hang Tran, Guillaume Tourrel, et al. Steerable Thin-Film Electrode Array for Cochlear Implantation: Design and Development for Future Atraumatic Insertion. Sensors and Actuators Reports, 2026, 11, pp.100447. 10.1016/j.snr.2026.100447. hal-05488634

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# Steerable thin-film electrode array for cochlear implantation: design and development for future atraumatic insertion

Ahmad Itawi<sup>a</sup>, Bacem Zribi<sup>a</sup>, Prabhakar Sidamaram<sup>b</sup>, Thanh Hang Tran<sup>a</sup>, Guillaume Tourrel<sup>b</sup>, Renato Torres<sup>d,e</sup>, Sofiane Ghenna<sup>a</sup>, Sébastien Grondel<sup>a</sup>, Yann Nguyen<sup>d</sup>, Cédric Plesse<sup>c</sup>, Eric Cattan<sup>a,*</sup>

## ARTICLE INFO

**Keywords:**

Electro active polymer

Cochlear implant

Thin film electrode array

Dynamic curvature

## ABSTRACT

During cochlear implant surgery, standard electrode arrays are inserted into the scala tympani to stimulate the spiral ganglion cells and rehabilitate hearing in deaf patients. However, conventional electrode arrays' stiffness and passive nature lead to potential trauma or incomplete insertion during the procedure. To overcome these limitations, an original steerable thin film electrode array (TFEA) has been developed. First, the twenty gold electrodes, distributed over a 25 mm length, with an average surface area of 0.16 mm², are significantly larger than those of existing TFEAs. These larger electrode surface areas enable safe neural stimulation within charge density limits below the Shannon threshold. By adjusting the material thicknesses, the proposed TFEA offers tunable stiffness, enabling safer and more flexible insertion. The microfabrication process, using SU-8 negative photoresist thin films, is both cost-effective and straightforward. In addition, the ability to dynamically adjust the curvature of the TFEA during insertion into a 3D printed cochlea model using low voltage conducting polymer based micro-actuator has been demonstrated. This marks the first instance of electrode array insertion with adaptive curvature, minimizing contact with cochlear walls. Successful insertion was achieved, with a curvature angle close to 360°. This active TFEA has the potential to improve insertion control and reduce the risk of trauma during cochlear implantation.

## 1. Introduction

According to World Health Organization<fnref n="1" /> statistics, over 5% of the world's population - or 466 million people - require rehabilitation to address disabling hearing loss. While hearing aids are adequate for individuals with mild to moderate hearing loss, those with severe to profound hearing loss require cochlear implants. A cochlear implant system comprises two main components: (i) an external sound processor equipped with a microphone, batteries, electronics, and a transmitting coil; and (ii) an implanted stimulator (IS) containing stimulation electronics, a receiving coil, and a standard electrode array (SEA). The sound processor captures the surrounding sound and transmits the information wirelessly to the IS, which converts it into electrical impulses. These

impulses are delivered via the SEA to stimulate the spiral ganglion cells of the auditory nerve, thus enabling patients to hear. SEA design has to deliver high frequencies at the proximal part and low frequencies at the distal part of the array, nevertheless any SEA can cover the whole cochlea, thus there is a mismatch between the frequency delivered by the electrode and its position in the cochlea without real impact on auditory performance. However, different SEA lengths are available to fit various cochlear dimensions and depths, providing comprehensive tonotopic coverage and targeting spiral ganglion cells along the cochlea's full median circumference. The risk of failed or incomplete insertion is due to the surgeon's lack of real-time feedback on SEA positioning within the scala tympani. Thus, during SEA insertion risks of translocation, damaging delicate intra-cochlear structures such as the basilar

<footnote><sup>a</sup> Univ. Polytechnique Hauts-de-France, CNRS, Univ. Lille, UMR 8520 - IEMN, Valenciennes, France</footnote>

<footnote><sup>b</sup> Oticon Medical, Research & Technology, Vallauris, France</footnote>

<footnote><sup>c</sup> CY Cergy Paris Université, LPPI, Neuville sur Oise, France</footnote>

<footnote><sup>d</sup> Université Paris Cité, Institut Pasteur, AP-HP, Inserm, Fondation Pour l'Audition, Institut de l'Audition, IHU reConnect, Paris, France</footnote>

<footnote><sup>e</sup> Service ORL et Chirurgie Maxillofaciale, Centre Hospitalier Sud Francilien, Corbeil-Essounes, France</footnote>

<footnote>* Corresponding author.</footnote>

<footnote>E-mail address: eric.cattan@uphf.fr (E. Cattan).</footnote>

<footnote n="1">¹ Deafness and hearing loss</footnote>
membrane or spiral ganglion cells, can trigger inflammation, fibrosis, and ultimately result in poor post-operative outcomes, such as residual hearing loss or impaired auditory perception [1].

The design of SEAs has been a longstanding concern for CI manufacturers and research teams. A comprehensive overview of current SEA designs is presented in [2]. All commercially available SEAs are currently hand-fabricated, assembled manually, and consist of a fine wire-bundle design with 12 to 24 platinum or platinum-iridium alloy electrodes. These arrays typically measure 16 mm to 30 mm in length, with electrode spacing of 0.75 mm to 2.4 mm, and their surface areas between 0.12 mm² to 1.5 mm² [2,3]. The electrodes are overmolded with a biocompatible silicone material required for SEA construction. Pre-curved electrode arrays are also commercially available [4]. However, these pre-curved modiolar hugging electrode arrays are shorter than lateral wall electrodes and typically cover the basal turn and the beginning of the second turn of the cochlea, while the spiral ganglion cells are distributed along Rosenthal's canal, which extends through approximately 1.75 turns of the cochlea [5]. In addition, translocation of the electrode array from the scala tympani to the scala vestibuli is an undesirable phenomenon that is seen with all types of electrode arrays, which increases the risk of damaging cochlear structures such as the osseous spiral lamina or spiral ligament. Such damage can lead to new bone formation, potentially affecting the hearing performance of cochlear implant users. The risk of trauma with pre-curved electrode arrays is increasing by their fixed curling geometry, which may not match the variable coiling of individual cochlea. Perimodiolar electrodes are straightened with either an internal stylet or an external sheath to enable their insertion into the cochlea. During insertion, the orientation of the straightened electrode contributes to the risk of electrode translocation. Electrode curvature is achieved by either removal of the stylet or advancement of the electrode from the sheath. For the straight SEAs, the number of electrodes in the array is constrained by both the available space and the need to minimize the mechanical stiffness. This limitation leads to friction between the SEA and the lateral wall of the scala tympani due to a significant mechanical mismatch between the stiffness of SEAs (e.g., Young's modulus $E_{SEA} \approx 210$ GPa) and the much softer basilar membrane (e.g. $E_{bm} \approx 50$ kPa) [6]. This mismatch, which can differ by up to seven orders of magnitude, poses a risk of causing irreversible tissue damage. Electrode stiffness, largely determined by the wire bundle, leaves little room for modifications to improve performance.

To achieve less traumatic implant insertion, reduce friction, and improve the cochlea-electrode interface, research is moving away from traditional manufacturing methods toward advanced thin-film electrodes manufactured using microfabrication techniques. This transition is largely driven by the success thin-film electrode for stimulation, recording, and monitoring of brain activity [7,8]. Notably, K. D. Wise and his team in 1970 were among the first to successfully fabricate biological microprobe arrays using microfabrication techniques, enabling the recording of neural activity from small populations of neurons in the brain [9,10]. Building on this success, in 2014, A. C. Johnson and K. D. Wise proposed an advanced cochlear thin-film electrode arrays (TFEA) that featured stacked layers of parylene and metal. This TFEA, comprising 32 stimulation sites with iridium oxide, offered as curved, position-controlled design along with integrated position sensing and a system-level application-specific integrated circuit [11–13]. Several other innovative methods of TFEAs microfabrication have also been proposed. In 2006, W. A. Johnson proposed a polyimide TFEA with reinforced base stiffness to avoid buckling [14]. In 2013, N. S. Lawand et al. designed and microfabricated a flexible polyimide device with a titanium nitride electrode, and accompanied by a compact programmable biphasic stimulator chip [15]. In 2019, Y. Xu et al. developed a parylene TFEA with 15 electrode sites along an 8-mm length, each site having a 75 µm diameter Pt disk. Their TFEA achieved an average impedance of 16 kΩ and a maximum current limit of 800 µA during monopolar stimulation in saline solution using a 1-kHz sinusoidal

current [16]. This simplified TFEA prototype marked significant progress toward automating the fabrication of cochlear implant electrodes, an important step to replace labor-intensive and costly manual assembly processes. This technology also offers smaller array dimensions, reproducibility of properties, and easy cost-effective batch fabrication. Additionally, more flexible electrode arrays are used for potentially atraumatic insertion into the scala tympani of the cochlea. Despite these advancements, none of the TFEAs have been integrated into commercial CI applications for human use. TFEAs, therefore, need to be improved so as not to exceed the Shannon limit, to reduce the risk of the bulking effect, to allow curvature and to better control insertion, so that in the future, insertion can be customized for each patient's cochlear anatomy.

In this context, we propose to fabricate an SU-8 TFEA with gold electrodes placed on a higher level than the power supply tracks, offering greater flexibility in electrode surface area modulation. One of the key benefits is that the gold tracks are well-protected and maintain electrical conductivity even when the TFEA undergoes curvature. Let's recall here that Shannon's relation, which defines safe stimulation thresholds by correlating several stimulation parameters, is often neglected in TFEA design. However, Y. Xu et al., whose studies show results beyond the Shannon limit [17,18], have addressed this issue by proposing to increase the exposed electrode surface in order to widen the dynamic range without compromising electrode density [16]. Thus, the first objective of our work is to increase the geometric surface area of stimulating electrodes through fabrication, enabling the design to remain within the Shannon limit.

Although various approaches are currently under investigation, it should be noted that the challenges posed by charge transfer on smaller electrodes are beyond the scope of this article. For instance, surface texturing electrode techniques such as laser roughening [19], chemical etching or metal electrodeposition [20,21] have been used to increase surface roughness, thereby expanding the electrochemically active area and reducing impedance. Additionally, metal electrodes have also been enhanced with conductive coatings, such as iridium oxide (IrOx) [22–24] and poly- (3,4)-ethylenedioxithiophene (PEDOT) [25–28], a widely used electronic conducting polymer (ECP). These coatings, particularly IrOx and PEDOT, are prevalent in neural engineering for their ability to increase the electrode electrochemical surface area and modify charge transfer mechanisms at the electrode-tissue interface. Such a solution could be integrated with our TFEA in future developments. Similarly, the long-term stability and aging of these electrodes must be evaluated, especially for applications requiring implants to remain functional the lifetime of the patient. Addressing these factors will be critical to ensuring the durability and reliability of next-generation cochlear implants.

A second objective of this work is to allow enough flexibility so as to be able to bend the TFEA and therefore to follow the natural shape of the cochlea by limiting the friction as possible. Currently, there are few methods explored to achieve perimodiolar positioning of SEAs or TFEAs with actuators [29]. Existing approaches rely usually on the use of external magnetic guiding [30,31], tubular manipulators [32], fluid-mechanical compliant actuator [33,34], shape memory materials [35–37], or micro-actuators to change EA curvature by hydrogel swelling, [38]. Nevertheless, these methods have no precise control over the bending. While ECPs have been widely studied to increase charge density, their potential to actively modulate the curvature of SEAs or TFEAs during insertion remains unexplored, despite a conceptual proposal by D. Zhou in 2003 [39]. To date, no concrete experimental results have demonstrated the use of ECPs for guided insertion of SEA or TFEA. ECPs can swell and contract reversibly under low voltage stimulation (< 1 V) due to electrochemically controlled ion exchange (insertion/expulsion) with a surrounding electrolyte. By associating an ECP layer with a passive thin film in a bilayer or trilayer configuration, significant bending deformation can be obtained in numerous electrolytes, including physiological ones. Moreover, these materials have demonstrated the ability to sense mechanical stimulation [40,41]. However,
their current limitation is the insufficient force produced to bend a SEA. As TFEAs become less rigid, this challenge may be overcome. Given that TFEAs are long (20 to 25 mm) and narrow, with actuator force inversely proportional to these dimensions, optimizations are necessary to achieve suitable curvature in the cochlea while minimizing friction.

The outline of this work is the following: Section 2 reports first presents the TFEA materials and actuators, followed by the micro-fabrication of an SU-8 TFEA with gold electrodes, and the electro-chemical characterization method. Then Section 3 describes the electrical and mechanical properties of the TFEA, with and without the ECP-based actuator. Finally, we demonstrate the effectiveness of complete insertion of the TFEA when coupled with the ECP-based actuator.

## 2. Design and fabrication

### 2.1. TFEA materials and design

The TFEA combines gold layers and SU-8. SU-8 was chosen because of the following favorable properties: it is a negative photoresist with a low curing temperature, and annealing precautions can be taken to limit internal stresses. SU-8 has a Young's modulus ranging from 0.86 to 2.33 GPa, which is a more flexible material than parylene (~3.2 GPa) or polyimide (~2.5 GPa). SU-8 has a fracture strength of approximately 89 MPa [42]. In addition, SU-8 is mostly immune to hydrolytic degradation, enabling long-term usage in the cochlear perilymph and is biocompatible [43]. Gold was selected over other high-performance materials (e. g., Pt, IrOx) due to its compatibility with our etching processes and its lower residual stress on the TFEA structure, ensuring better fabrication control and mechanical stability. It is easy to obtain thin film tracks using simple, inexpensive microfabrication techniques such as evaporation and chemical etching without damaging the SU-8 layers.

Moreover, gold exhibits biocompatibility and acceptable electro-chemical stability due to its relative corrosion resistance. Chromium (Cr) was chosen for its strong adhesion to SU-8 photoresist. The schematics and design of the TFEA, fabricated with twenty electrodes numbered Ei, are shown in Fig. 1, with the key dimensions provided in Table 1. Each electrode is rectangular with a surface size that decreases from the basal to the apical electrodes. The TFEA features a rectangular sharp side edges and a rounded tip at the apical end of the TFEA for easy insertion. The basal end contains terminal gold pads for electrical connections adapted to a zero-insertion force (ZIF) connector.

### 2.2. TFEA microfabrication

The process was performed on a temporary silicon wafer in a clean

room using standard microelectronic equipment. The temporary substrate was separated at the end of the process to obtain the flexible TFEA. Silicon oxide, used as the sacrificial layer, was deposited using plasma-enhanced chemical vapor deposition at 300 °C (Fig. 2.1). An 8-µm thick negative photoresist SU-8 2010 (MicroChem, USA) was spin-coated (3500 rpm/300 rpm s⁻¹/³⁰ s) with a standard spin coater (RC8, Karl-Süss, France), annealed by hot plate 95 °C during 3 min, exposed to UV light (10 mW·cm⁻², 365 nm) during 12.5 s, annealed by hot plate 95 °C during 3 min, and developed by SU-8 developer during 4 min (Fig. 2.2). The first metallic layer, Cr/Au with a thickness of 10/300 nm, was deposited through evaporation (Plassys MEB 550SL, France). The geometries were then defined using the positive photoresist S1828 at a thickness of 2 µm. Chemical etching with KI:I₂:H₂O (4:1:40) during 3 min and chrome etching during 20 s were used to obtain the Cr/Au metallic circuit. The wafer was immersed in acetone to remove residual photoresist (Fig. 2.3). Then, one layer of SU-8 2002 was deposited by spin-coating: the first one, 2 µm thick (3000 rpm/300 rpm s⁻¹/³⁰ s), was completely exposed to UV light during 9 s, followed by a second layer of SU-8 2010, 4 µm thick (4000 rpm/300 rpm s⁻¹/³⁰ s), was exposed during 10 s and developed by SU-8 developer during 60 s to define the electrical via contacts (10 µm x 350 µm) above the Au lines for the metal interconnects (Figs. 2.4 and 3b), soft bake and post exposure bake are the same for all SU-8 layers. The second layer of SU-8 served as a physical mask to open the first layer using oxygen plasma (300 W, 200 mTorr, 50 sccm) with a reactive ion etching (OxfordPlasmalab 80+, U. K.) and an end-of-etch detection. After this step, the combined thickness of these two layers was reduced to 2 µm (Fig. 2.5). The second metal layer, Cr/Au (40/800 nm), was deposited by evaporation and patterned using the same photoresist and wet etch process as described before (Fig. 2.6). Chromium was used rather than titanium as an adhesion layer for the gold layer to guarantee compatibility with the process. Titanium is etched in diluted hydrofluoric acid similarly to the sacrificial SiO₂ layer, which is not the case for chromium. Finally, the sacrificial layer was removed, and the foldable TFEAs were rinsed with deionized water (Fig. 2.7). The final thickness of the TFEA was approximately 10 µm [44].

### 2.3. Actuator material and design

The polymeric actuator composed of an NBR-PEO ion storage membrane sandwiched between two electroactive PEDOT:PSS-PEO layers was fabricated using the layer stacking method [40,45,46] (see Fig. SI-1). The trilayer was then micro-patterned as microbeams, using laser cutting, with the same geometry as the TFEA (Fig. 1) and with a thickness in the range of 50 µm to 60 µm. Before any actuation, the

<figure><img src="image_3.png" /><figcaption>Fig. 1. Design of the TFEA terminal gold pads adapted for a ZIF connector, with the ZIF connector added to the TFEA.</figcaption></figure>
<table><caption>**Table 1**
Most important dimensions of the TFEA.</caption><thead><tr><th>Length [mm]</th><th>Apical width [mm]</th><th>Basal width [mm]</th><th>Exposed electrode surface at basal [mm²]</th><th>Exposed electrode surface at apical [mm²]</th><th>Line size & spacing [µm]</th><th>Total thickness [µm]</th></tr></thead><tbody><tr><td>25</td><td>0.3</td><td>0.5</td><td>0.4 × 0.4</td><td>0.25 × 0.40</td><td>10</td><td>10</td></tr><tr><td>Geometric Area</td><td>E0</td><td>E4</td><td>E8</td><td>E9</td><td>E13</td><td>E17</td></tr><tr><td>[×10<sup>-3</sup>cm²]</td><td>0.76</td><td>1</td><td>1.24</td><td>1.3</td><td>1.54</td><td>1.6</td></tr></tbody></table>

<figure><img src="image_4.png" /><figcaption>**Fig. 2.** Illustrative schematic of the TFEA fabrication process.</figcaption></figure>

micro-actuator beams were swollen in ionic liquid (IL) – 1-ethyl-3-methylimidazolium bis-(trifluoromethylsulfonyl)imide (EMImTFSI) – for 72 h to incorporate the EMI⁺ and TFSI⁻ ions necessary for the actuation process.

## 2.4. 3D-printed cochlea model fabrication

The ST shape was generated thanks to Nautilus research software from Cochlear Ltd. Nautilus software relies on a database of more than a thousand patient CT scans. These scans were automatically segmented and fed a parametric shape model of cochlea. Nautilus software can generate and export in 3D file format any cochlea shape by choosing parameters [47]. The cochlea model was fabricated using a stereolithography printing process on DWS29J+printer. This technology consists in stacking 2D layers to build the 3D shape of the printed component. These layers are created in a photosensitive liquid resin tank. Layer by layer, a laser beam polymerizes the resin where solidified material is needed. The model was made of DS3000 resin, which provides good accuracy and transparency. The printing was performed with 10 µm thickness layers. The model was controlled with 3D µ-tomography (RX Solution, France) with a 12 µm voxel size resolution. The scanned cochlea model was reconstructed and compared to the reference file. The 3D printed model matched correctly to the reference within roughly ±40 µm, with no obvious default along the ST. We controlled the ST cross section every 90° and confirmed that the 3D printed model was highly accurate, with shape comprised most of time within ±0.05 mm, and sometimes-local default within ±0.1 mm.

## 2.5. TFEA electrochemical characterization method

The TFEAs were characterized using a Princeton Applied Research Parstat MC 2000 (Ametek Instruments) potentiostat with a three-electrode system in an artificial perilymph (AP) electrolyte set-up for stimulation safety investigations. The electrochemical characterization was carried out using Ag/AgCl (3 M potassium chloride (KCl)) as the reference electrode, Pt wire as the counter electrode, and gold electrodes (TFEA) as the working electrodes. The AP solution was prepared by dissolving 125 mM sodium chloride (NaCl), 3.5 mM KCl, 25 mM sodium bicarbonate (NaHCO₃), 1.2 mM magnesium chloride hexahydrate (MgCl₂), 1.3 mM calcium chloride (CaCl₂), and 0.75 mM monosodium phosphate (NaH₂PO₄) in distilled water, with the pH maintained at 7.4 [48,49]. All electrolytes were freshly prepared and used immediately with no degassing or modifications before performing the electrochemical experiments.

The charge storage capacity (CSC) and the charge injection capacity (CIC) of the gold electrode arrays with different geometric areas were calculated using cyclic voltammetry (CV) and chronopotentiometry, respectively [50] (see Fig. SI-2 c)). CV was performed in AP at a scan rate of 100 mV/s to determine the CSC. The CSC was calculated by integrating the current with respect to voltage during the cathodal and anodal phases, using the Formula (1):

$$CSC = \frac{1}{v \times A} \int_{V_1}^{V_2} i(V)dV \ (\text{where } i(V) \le 0) \quad (1)$$
<figure><img src="image_5.png" /><figcaption>Fig. 3. SU-8 TFEA after release from the silicon substrate, a) image of 25-mm long TFEA; b) Optical image of one 400 µm x 400 µm gold square electrode. In the image, the connection zone with the lower track is visible; this via is 10 µm wide and 350 µm long; c) Cross sectional SEM image of the interface between the first gold layer and the second SU-8 layer where a recess is observed and the slope of the SU-8 wall is slightly negative; d) SEM image after deposition of two SU-8 layers and reactive etching, the second layer is used as a physical mask. The wall profile is positive and the gold layer is continuous between the top and bottom electrodes.</figcaption></figure>

Where v is the scan rate (V/s), A is the geometric area of the electrode (cm²), V₁ and V₂ are the water reduction and oxidation potential, respectively, and i(V) is the current.

To determine the electrode's safe potential window within aqueous water window, cyclic voltammetry measurements were conducted.

For the stimulation process (CIC), chronopotentiometric measurements were performed using a stimulation pulse with a pulse amplitude of 0.1 mA and a pulse width of 10 µs (n = 5 pulses). The stimulation pulse used in this study was anodic-first, charge-balanced, and featured an asymmetric cathodic passive discharge. This was chosen for its exceptional capability to maintain charge balance, thus reducing the risk of electrode damage and ensuring safe and effective electrical stimulation [51].

The CIC was determined according to the total charge delivered/ injected during each cathodal and anodal phase of the stimulation pulse. The Formula (2) for CIC is given as:

$$CIC = i \times t \qquad (2)$$

Where CIC is the total injected charge (µC) during the stimulation pulse. The CIC is a key metric for determining whether an electrode can deliver charge safely without exceeding polarization limits or causing electrode damage.

Electrochemical impedance spectroscopy (EIS) was used to determine the impedance and phase change of the electrode arrays, before and after stimulation. The TFEA was placed within a three-electrode cell where the gold of the TFEA was placed as the working electrode, the Pt wire as the counter electrode, and the Ag/AgCl as the reference electrode, using AP (pH 7.4) as the electrolyte for ionic conduction. Impedance measurements were carried out within the frequency range of 1 to 10 kHz at 0 V vs. Ag/AgCl and with an amplitude of 10 mV.

## 2.6. Stiffness measurements

Actuator and TFEA+actuator were placed horizontally for stiffness measurements. The TFEA was clamped between two glass slides near the basal end where the width is 500 µm. A force sensor probe (FT-S micro-force sensing probe, FemtoTools AG) fixed to a micromanipulator xyz was precisely positioned using two micro-cameras. The force sensor F was brought into perpendicular contact with the TFEA surface. The micromanipulator was moved at 10-µm intervals to push on the TFEA and the blocking force was recorded at each new position. A bending force was applied at 2.5 mm away from the clamping position in the direction normal to the surface of the TFEA. The slope of the force recorded as function of the displacement d give the bending stiffness K following the Eq. (3):

$$F = Kd \qquad (3)$$

# 3. Results and discussion

## 3.1. TFEA microfabrication

Nine TFEAs were fabricated on a single 3-inch wafer due to the large size of the header required for the ZIF connector and the electrical measurements. By reducing this area, it is possible to manufacture around thirty prototypes on such a surface. No residual stress was observed; the TFEA remained flat when released. In Fig. 3a, the 25-mm SU-8-TFEA samples display 20 stimulation electrode tracks in the flared section. This section can be easily modified for future connection with the IS and the existing production control system. Fig. 3b presents an optical image of a gold square electrode measuring 400 µm x 400 µm near the basal end of the TFEA, revealing the via at the connection zone
with the buried electrode track.

To define the oblong via, two layers of SU-8 were used. Using only a single layer of negative photoresist, followed by exposure and development, results in a profile that does not adequately cover the gold on the lateral walls or establish contact with the track electrode that passes below. This profile exhibits a slightly negative slope, and a recess forms at the interface between the SU-8 layer and the gold track, as depicted in Fig. 3c. This recess prevents continuity of the gold layer, obstructing electrical contact between the track and the top electrode. To avoid this issue, two layers of SU-8 are employed. The first is fully exposed, and the second layer serves as a mask. The entire second layer is then removed using reactive O₂ plasma etching. The total thickness of the two SU-8 layers was reduced to 2 µm.

This method enhances the slope on the walls of the interconnection via, producing a smooth trapezoid structure to ensure the evaporated metal continuity between the top electrodes and the tracks (Fig. 3d). Consequently, this approach ensures that 100 % of the top electrodes are in contact with the lines. Impedance measurements were performed both on the flat TFEA after repeated bending the TFEA and under curvature with significant angulations, and no breaking electrical contacts between the base and the stimulation electrode were observed. The 10 µm wide lines of gold are placed approximately on the neutral axis and are encapsulated between two layers of SU-8, which limits the risk of track breakages.

This TFEA fabrication method can be compared to the state-of-the-art design by Johnson and Wise [12] and Xu et al. [16]. However, even though the number of lithography, deposition, and reactive ion etching steps is nearly the same, current TFEA microfabrication offers the following advantages. (i) Parylene C is replaced by a photoresist easy to pattern and does not require an expensive specialized chemical vapor deposition process. (ii) The gold square electrodes are on the surface of the TFEA and therefore will be closer to the areas to be stimulated. (iii) The size of the top electrode could be adapted (see Table 2), with a width close to that of the TFEA and a length that could be modified depending on the surface area that needs to be stimulated. The surface obtained is forty to one hundred times larger than the electrode size obtained in the studies [12,16]. This large surface has a strong impact on the electrochemical measurements, as we show later. (iv) Collective manufacturing, unlike SEA manufacturing with molding and manual assembly, provides more flexibility to vary the lengths of the TFEAs, the surface area of the electrodes, or the number of electrodes. Ultimately, this should make it possible to manufacture TFEAs tailored to each cochlea in terms of insertion depth and coverage of the middle turn of the cochlea, so as to encompass all spiral ganglion cells. It also allows several dozen TFEAs to be manufactured in a single batch, which should significantly reduce the price per TFEA and broaden the range of patients.

## 3.2. TFEA electrochemical characterizations in vitro

This study investigates the electrochemical properties of TFEAs with different geometric electrode areas and assesses their influence on safe stimulation strategies. The electrodes were characterized through cyclic voltammetry in artificial perilymph, highlighting their capacitive behavior, defining their potential window, and providing insights into the reaction mechanism at the electrode surface (see Fig. SI-2). The results reveal the full potential range and establish a safe operational potential window, ensuring that the electrode operates below the threshold for water electrolysis.

From the cyclic voltammetry measurements, it was evident that the gold electrodes in artificial perilymph began to oxidize at around +0.6 V (vs. Ag/AgCl), with water electrolysis occurring at significantly higher potentials. The CIC and CSC values for different electrode areas are tabulated in Table SI-1. Notably, the cathodic CIC and CSC are consistently higher than their anodic counterparts across all electrode areas. This can be attributed to the presence of dissolved oxygen in the

artificial perilymph, which enhances cathodic charge transfer through reduction reactions. During the cathodic phase, the reduction of oxygen molecules increases electron uptake, elevating the overall charge capacity. Additionally, the increased geometric electrode area, as shown in Fig. 4a, provides a larger surface for oxidation and reduction reactions at the electrode-electrolyte interface. This extended surface area allows for a higher density of reaction sites, thereby increasing the CSC and facilitating more efficient charge transfer during stimulation [52]. We attribute this to perimeter (edge) effects and to differences between geometric and electrochemically active area due to surface roughness, residue of photoresist and passivation-edge topography, as evidenced by the SEM in Fig. 3dThese features locally increase the effective area and double-layer capacitance, effects not captured by geometric normalization, thereby modestly increasing CSC during stimulation.

To assess the electrode behavior under realistic stimulation conditions, chronopotentiometric experiments were conducted [53]. These were not intended to determine the water window (which was characterized separately using cyclic voltammetry), but rather to evaluate the charge injection capacity (CIC) and electrode polarization during pulsed current stimulation. In this study, anodic-first, charge-balanced, asymmetric cathodic passive discharge stimulation [51,54] with a 0.1 mA pulse amplitude and a 10 µs pulse width was applied to the gold electrodes to mimic neural stimulation scenarios (see Fig. 4b).

In an ideal electrode, applying an anodic-first/oxidation pulse increases the electrode potential to a more positive level than the open circuit potential (OCP), and the subsequent cathodic/reduction pulse would return the potential back to the OCP without inducing water electrolysis. However, gold electrodes, being ideally non-polarizable, depend on Faradaic currents resulting from oxidation and reduction reactions. These reactions take place within a safe water window between -0.4 V and +0.9 V vs. Ag/AgCl, avoiding water electrolysis during stimulation. The increase in potential during the anodic phase results from the capacitance, hydride desorption, and metal oxidation. The changes in electrode potential (n = 5 pulses) were observed as shown in Fig. 4bAs the anodic potential increased, the cathodic potential decreased, a phenomenon known as the ratchet effect [53].

Fig. 4c shows that the electrode impedance contains both resistive and capacitive components in the frequency range, with a phase angle ranging from -78° to -2° and an impedance magnitude of <1 kΩ at 1 kHz. To better understand the effects of stimulation, impedance measurements were performed both before and after stimulation. It was observed that as the geometric area increases, impedance decreases for both pre- and post-stimulation measurements. While no significant differences were observed at higher frequencies, a slight decrease in impedance and phase value at lower frequencies was observed. This decrease is attributed to the electrode surface activation during stimulation (current sweeping) and the removal of surface impurities.

To establish safe stimulation limits, the Shannon plot was used, which defines a threshold beyond which tissue damage can occur [17]. The graph plots both the charge density D of the gold electrode array, calculated by dividing the charge per phase (stimulating pulse) by the electrode's geometric area of injected charge, represented in µC/phase/cm², and Q the total charge injected per phase (stimulating pulse @115 µs), represented in µC/phase (see Table SI-2). Shannon's plot is derived from Eq. (4) [18,52,55]:

$$ \log(D) = k - \log(Q) \quad (4) $$

where k is an empirically determined parameter.

This plot was originally derived from histological data obtained from cat brain studies using charge-balanced stimulation pulses [56]. The safe CSC and CIC of the electrodes were seen to be highly dependent on the electrode properties and stimulation conditions [50,57,58]. Understanding how electrode geometry influences electrochemical and electrophysiological responses is crucial for developing neural stimulation devices and ensuring safe stimulation practices. It is hypothesized that
<figure><img src="image_6.png" /><figcaption>**Fig. 4.** Electrochemical characterization: a) charge storage capacity (CSC) plotted against electrode area; b) Chronopotentiometric data for anodic-first asymmetric charge-balanced pulse; c) The E8 Bode plots before and after stimulation obtained in the frequency range of 1 Hz to 10 kHz at 0 V vs. Ag/AgCl by applying 10 mV (AC mode) in AP (pH 7.4); d) Charge densities at various electrode geometries for safe stimulation with Shannon limit, $k = 1.8$.</figcaption></figure>

the k-value represents the upper toxicity level of platinum/metal that can be tolerated by living cells in close proximity to the stimulating electrode [59]. Shannon suggested a band within which the damage limit could be found, $1.5 \le k \le 1.8$, however, the physical or physiological causes explaining the damage limit remain uncertain. A better understanding of underlying mechanisms and safe stimulation guidelines is needed. Shannon's relationship ignores some important parameters [60], including frequency, duty cycle [61], decreased electrode area and change geometry [59,62,63], and roughness [61,64], which can affect damage thresholds. This suggests that the damage limit should be handled cautiously, and further work is needed to adapt these limits for the microelectrode size.

In this study, using a k-value of 1.8, the safe limit on the Shannon plot was identified, and charge densities below this threshold were considered safe. Charge densities for the gold electrode array were within the safe stimulation range, below the 2 mA clinical current value (Association for the Advancement of Medical Instrumentation), except for the smallest geometric areas of 0.00076 cm² (E0) and 0.001 (E4) cm², as shown in Fig. 4dIncreasing the surface area of the last electrodes located at the tip of the TFEA (E0 to E7) can be easily achieved by modifying the mask used for the gold wet etching level (Fig. 2, step 6). The electrodes can be widened slightly and more particularly elongated. In conclusion, our TFEA ensures safe and effective neuronal stimulation thanks to the properties of the large gold electrode, stimulation conditions, and a thorough understanding of their influence on the charge density, as highlighted by the Shannon plot.

### 3.3. TFEA/actuator mechanical properties and insertion

For decrease trauma during implantation the objective is first to reduce the stiffness K of SEA [65]. This is obtained by reducing their thickness drastically (500 µm to around 10 µm). TFEA. The same measurements were conducted on the actuator, and on the TFEA coupled with the actuator (Fig. 5b), the actuator having the same shape as the TFEA. Coupling was achieved by adding a few very small spots of glue on the edge at the top, in the middle, and near the base. Bonding over the entire surface of the TFEA would considerably reduce the curvature by preventing the PEDOT/PSS layer in contact with the TFEA from contracting. The stiffness measured at 2.5 mm from the clamping point was as follows: for TFEA $K_{TFEA} = 0.015$ N/m, for the actuator $K_{Act} = 1.755$ N/m and for the TFEA coupled with the actuator $K_{TFEA+Act} = 4.017$ N/m (Fig. 5b). To confirm the value of the TFEA stiffness, a finite element simulation was performed on a SU-8/Au bilayer (with gold layer on full surface) with similar conditions regarding dimensions, material properties, and boundary conditions. The result of the simulation gave a stiffness of 0.042 N/m not so far from the value measured. The stiffness of the TFEA+Act was therefore twenty times lower than that measured for a SEA (EVO electrode array from Oticon Medical), and it is tunable by adjusting the thicknesses of each layer.

The same setup is used for the actuator force measurements: the force sensor is brought into contact with the actuator and a square voltage of 1 V at a frequency of 0.05 Hz is applied. The actuator bends and applies a force for four seconds to the sensor. This duration is considered
<figure><img src="image_7.png" /><figcaption>**Fig. 5.** a) Stiffness setup measurement, b) comparison of the stiffness of the TFEA, Actuator, and TFEA+Act measured at 2.5 mm from the clamping point, c) Zwick Roell traction/compression machine with 5-N force sensor, which records the displacement and insertion force simultaneously, used as an automated insertion test bench. d) Comparison of the insertion forces between SEA and TFEA in a 3D printed model of an L-sized cochlea with NaCl.</figcaption></figure>

acceptable for surgical TFEA implantation. The maximum forces measured at 2.5 mm, 8 mm, and 12 mm from the clamping point were 550 µN, 230 µN, and 90 µN respectively. These forces are sufficient to bend the TFEA several hundred micrometers inside a cochlea. Indeed, if we extrapolate the stiffness of the TFEA+Act to a distance of 12 mm, it would be 0.036 N/m and, therefore, with the force of 90 µN, a theoretical local axial displacement of 2.5 mm becomes possible, which corresponds to the targeted application since the free space available in a cochlea is of the order of a millimeter. Of course, this calculation does not take into account the viscous medium of the cochlea and the fact that the actuator coupled to the TFEA will cause compression of the TFEA and not just axial thrust.

The insertion forces between a SEA and a TFEA were compared on a test bench using a Zwick Roell traction/compression machine with a 5-N load cell (Fig. 5c). The speed was controlled by the machine and remained constant during the measurements. A transparent 3D printed model of an L-sized cochlea was attached to the base support of the machine and filled with saline solution (NaCl). The TFEA sample was fixed to the load cell and strictly centered on the aperture of the cochlear model. The frictional forces produced by the SEA or TFEA when in contact with the walls of the 3D printed model of the cochlea (3D-PMC) were measured (Fig. 5d). Insertion of the TFEA generates extremely low frictional forces compared to the SEA. The SEA quickly rubs against the outer wall and the friction surface was greater than that with the TFEA. During insertion, the TFEA tip first pressed against the curved outer wall of the cochlea. This friction was reduced by creating a rounded shape at the tip (see Fig. 1, which shows the rounding in top view). However, the

TFEA is very thin in thickness and thus the edges are "sharp", unlike the SEA, which has a cylindrical shape. Later than SEA, these edges rubs against the outer wall, which is quite oval, resulting in sufficient friction to prevent the TFEA from moving forward (maximum insertion 16.5 mm). To reduce the friction caused by the edges along the cochlea rim, some research teams rounded off the backside of the TFEA, the part in contact with the outer surface of the cochlea, as described in [12]. This addition makes microfabrication more complex, whereas manufacturers aim to minimize the number of manufacturing steps. As W. A. Johnson [14] pointed out, the stiffness of the SEA is reduced to limit the friction forces during insertion, but the slightest local friction is enough to prevent progression, and the SEA can rapidly buckle at the entrance to the 3D-PMC. This effect was observed during the insertion of our TFEA when an angle of approximately 200° was reached, and deeper insertion into the cochlea was no longer possible without forcing. Therefore, an actuation system coupled with this TFEA would minimize contact with the walls of the cochlea by maintaining the TFEA as much as possible along the central axis of the cochlea Fig. 6.

The surgical operation to insert an implant should only take a few minutes. Conducting polymer-based actuators are generally considered slow actuators. The speed at which the actuator's radius of curvature decreases was therefore evaluated in air when a 1 V DC electrical voltage was applied for an actuator coupled or not with a TFEA. For a single actuator, the maximum radius of curvature was reached in approximately 10 s, but 90 % of the maximum curvature was reached in 5 s. As expected, when the actuator was coupled with the TFEA, the radius of curvature decreased more slowly: 15 s were needed and 90 % of the
<figure><img src="image_8.png" /><figcaption>**Fig. 6.** Evolution of the actuator's radius of curvature as a function of time in air when a DC voltage of 1 V is applied. The red line is for the actuator alone and the blue line is for the actuator coupled to a TFEA.</figcaption></figure>

maximum curvature was reached in 10 s. The speed to decrease the actuator's radius of curvature is compatible with the duration of the electrode array insertion (around 1 min). Higher curvature and curvature speed could be reached by increasing the voltage up to 3 V. In addition, tunable curvature during insertion was not necessarily based on a single DC voltage step but rather on successive voltage steps or voltage ramps. The smallest radius of curvature was 4 mm for the actuator alone and 9 mm for an actuator coupled with a TFEA and compressed the TFEA. When the TFEA was stretched by the actuator, the maximum radius of curvature obtained was 14 mm, indicating that it is easier to compress the TFEA than stretch it. For a favorable orientation of the electrodes, the TFEA must be compressed so that they are in the direction of the modiolus wall. Optimization of the TFEA stiffness and actuator force could be considered to improve the radii of curvature. Reducing the thickness of the actuator would increase the strain to the detriment of its actuation force.

For the insertion tests, a manual micro-positioning XYZ stage was used and the 3D-PCM was filled with IL, which has a viscosity between that of NaCl and glycerin. However, the slightly oily nature of the IL can cause a suction effect, altering the initial trajectory and causing the actuator to move toward the wall of the modiolus. After progressing 8

mm inside the 3D-PCM, the tip of the actuator alone meets the outer wall (Fig. 7a). A DC voltage of 1.5 V is then applied; the actuator tip is pressed against the wall of the modiolus (Fig. 7b).

The actuator is pushed again while maintaining the voltage applied. At the entrance to the 3D-PCM, the actuator moves toward the outer wall (Fig. 7c), and by continuing to push, the basal part of the actuator also presses against the outer wall (Fig. 7d). By increasing the voltage to 3 V, only the central part of the actuator remains in contact with the outer wall (Fig. 7e), allowing the actuator to advance within the 3D-PCM. Fig. 8f shows the final position of the actuator when the applied voltage is stopped and an insertion angle of 360° is achieved. If a voltage of 3 V is reapplied after a few hours of rest, the actuator's strain is improved, enabling it to penetrate deeper into the 3D-PCM with an angle of 430° (Fig. 7g). Finally, when the voltage is turned off again, the actuator remains stuck in its previous position, with its internal surface close to the wall of the modiolus inside the 3D-PCM (Fig. 7h).

The TFEA+Act insertion experiment was repeated several times. Fig. 8 presents a fairly representative example of the results observed. Initially, the TFEA+Act is positioned at the entrance to the 3D-PCM, ensuring it does not come into contact with the walls near the modiolus. It is moved, without any voltage applied, until it touches the outer wall (Fig. 8a). A small voltage of 1.5 V is then applied to keep the TFEA+Act as close as possible to the wall of the modiolus until the tip reaches the outer wall at an angle of around 180° (Fig. 8b). Increasing the voltage to 3 V allows the tip of the TFEA+Act to curve further and move closer to the wall of the modiolus in the 3D-PCM. The TFEA+Act is pushed further along the wall of the modiolus until an angle of approximately 270° is reached (Fig. 8c). The electrical voltage is maintained and the TFEA+Act is pressed closer to the outer wall until the tip remains stuck near 360° (Fig. 8d). At the entrance, a buckling effect is observed, indicating that the insertion of the TFEA+Act is less effective than the actuator alone. However, reducing the thickness of the TFEA would reduce its stiffness and enhance the curvature of the TFEA+Act. Further improvements and optimizations should be considered to increase the actuator force and the radius of curvature for a smaller cochlea.

This assembly offers several promising developments. (i) Once the global behavior of the TFEA+Act is well identified under different electrical voltage levels or profiles, a model and control system could be set up. This could be facilitated by switching the actuator to sensor mode [40]. After identifying the specific anatomy of the cochlea requiring treatment, a robotic arm can be programmed to follow insertion instructions, considering the entry angle, insertion speed, as well as the

<figure><img src="image_9.png" /><figcaption>**Fig. 7.** Micro-actuator alone inside a 3D-PCM. A DC voltage (1–3 V) is applied.</figcaption></figure>
<figure><img src="image_10.png" /><figcaption>Fig. 8. SU-8 TFEA coupled with a micro-actuator. A DC voltage (1–3 V) is applied to detach the TFEA from the outer wall and avoid friction inside the 3D-PCM.</figcaption></figure>

actuation of the TFEA and its interactions (contact and friction) with the cochlea. (ii) Figs. 7 and 8 illustrate that when the voltage is switched off, the actuator will release its flexion very slowly and the TFEA+Act returns to the outer wall of the 3D-PCM. As its stiffness is low, the forces applied to the lateral cochlear wall will be much lower than those produced by the SEAs. Similar to the standard pre-curved modiolar electrode arrays, pre-curving the TFEA+Act could be considered, so it would stay quite close to the wall of the modiolus when the voltage is off. Unlike standard pre-curved modiolar electrode arrays, which are generally only available to cover the basal turn of the cochlea, the TFEA+Act could more effectively cover the spiral ganglion cells distributed in Rosenthal's canal, which extends 1.75 turns into the cochlea. (iii) As the lifetime of any CI is typically less than twenty years, many CI recipients will require explantation and re-implantation during their lifetime. The actuator can then be reactivated to facilitate explantation.

## 4. Conclusion

This study introduces a novel flexible TFEA designed for cochlear implants, presenting significant advantages over standard electrode arrays. The TFEA concept was developed to address key challenges in cochlear implant technology, including low resolution, high fabrication costs (manually assembled), and friction encountered during insertion. Through micro-fabrication using SU-8 photoresist, the TFEA was designed with the following key features: (i) a large rectangular electrode surface (0.16-mm²) fabricated using gold evaporation and wet etching, (ii) an electrode density of 20 sites along a 20 mm length, achieved with a reduced number, cost-effective fabrication steps. (iii) wider electrodes than previous designs, achieving impedance values below 1 kΩ at 1 kHz.

This design represents a significant improvement, ensuring that charge densities remain within safe limits to avoid cellular toxicity in the ganglion cells of the auditory nerve near the stimulating electrodes. Using a Shannon plot with respect to a k-value of 1.8, it was confirmed that the charge densities for the gold electrode array remained below the safe stimulation threshold of 2 mA.

The simplicity of the design and the stability of the fabrication process offer a robust foundation for future optimization. By adapting materials and surface treatments, the electrochemical properties of the electrodes can be further enhanced. Previous studies on TFEAs demonstrated a significant reduction in stiffness compared to SEAs, with tunability easily achieved by adjusting the thickness of the structural material. However, despite the reduced stiffness, challenges such as buckling effect at the cochlear entrance and incomplete insertions for longer electrodes (15–25 mm) remain.

To overcome this, ECP micro-actuators were integrated into the TFEA. These actuators adapt the TFEA's curvature during insertion, minimizing the risk of buckling. By optimizing the thickness of the actuator, the forces generated were sufficient to bend the TFEA as needed. Testing in an artificial cochlea showed promising results, with the TFEA successfully detaching from the outer wall and moving closer to the modiolus. Under an applied voltage of 1.5 V to 3 V, the TFEA

achieved nearly 360° insertion, demonstrating full insertion of its 20 mm length with minimal friction and only limited contact with the outer wall.

These findings are encouraging, opening the possibility for more controlled and customized TFEA insertions based on the individual patient's cochlear anatomy.

Future work will focus on developing predictive models to identify the optimal timing for voltage application, guiding the TFEA along the best insertion path to further minimize contact with the cochlea.

In addition, the ECP micro-actuators, which also function as sensors, could enable a closed-loop system to further enhance control. Optimization of both the actuator and the TFEA will be necessary to ensure the long-term stability of the system in liquid environments under electrical stimulation. Testing in temporal bones will also be conducted to confirm the atraumatic properties of the TFEA during insertion. Finally, the biocompatibility of the system will be enhanced by using perilymph as electrolyte. Indeed, actuation tests were conducted in a 3D-printed cochlear model using EMImTFSI ionic liquid, chosen for its high electrochemical stability and ability to enable controlled bending of the actuator in ambient conditions. While EMImTFSI is not suitable for *in vivo* use, the actuator will be adapted in future work to operate in physiological media like artificial perilymph.

### CRediT authorship contribution statement

**Ahmad Itawi:** Writing – original draft, Validation, Supervision, Investigation, Formal analysis, Conceptualization. **Bacem Zribi:** Writing – original draft, Validation, Methodology, Investigation, Formal analysis, Conceptualization. **Prabhakar Sidamaram:** Writing – original draft, Formal analysis, Conceptualization. **Thanh Hang Tran:** Writing – original draft, Validation, Methodology. **Guillaume Tourrel:** Writing – original draft, Validation, Supervision, Investigation, Formal analysis, Conceptualization. **Renato Torres:** Writing – original draft, Supervision. **Sofiane Ghenna:** Writing – original draft, Investigation, Conceptualization. **Sébastien Grondel:** Writing – original draft, Validation, Investigation, Conceptualization. **Yann Nguyen:** Writing – original draft, Methodology, Conceptualization. **Cédric Plesse:** Writing – original draft, Validation, Investigation, Formal analysis, Conceptualization. **Eric Cattan:** Writing – review & editing, Writing – original draft, Validation, Supervision, Project administration, Methodology, Formal analysis, Conceptualization.

### Declaration of competing interest

The authors declare that they have no known competing financial interests or personal relationships that could have appeared to influence the work reported in this paper.

### Acknowledgements

This work is financially supported by the French National Research Agency through the ROBOCOP (ANR PRCE 19 CE190026), ACCESS (ANR-24-CE19-1137) projects and this work was supported by the
LAAS-CNRS MultiFAB platform and the French RENATECH network.

## Supplementary materials

Supplementary material associated with this article can be found, in the online version, at doi:10.1016/j.snr.2026.100447.

## Data availability

Data will be made available on request.

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