# Merging bioelectrochemical transducer and antenna functions for continuous monitoring in biological tissues: an innovative volume-saving strategy

Luigi Di Trocchio, Pablo Pérez-Nicoli, Sabrina Bichon, Vincent Raimbault, Jean-Luc Lachaud, Sébastien Gounel, Corinne Dejous, Alexander Kuhn, Claudine Boiziau, Nicolas Mano, et al.

<figure><img src="image_1.png" /></figure>

## To cite this version:

Luigi Di Trocchio, Pablo Pérez-Nicoli, Sabrina Bichon, Vincent Raimbault, Jean-Luc Lachaud, et al. "Merging bioelectrochemical transducer and antenna functions for continuous monitoring in biological tissues: an innovative volume-saving strategy." Biosensors and Bioelectronics, 2025, 290, pp.117970. 10.1016/j.bios.2025.117970. hal-05248693

HAL Id: hal-05248693

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Submitted on 10 Sep 2025

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# Merging bioelectrochemical transducer and antenna functions for continuous monitoring in biological tissues: an innovative volume-saving strategy

Luigi Di Trocchio<sup>a</sup>, Pablo Pérez-Nicoli<sup>a,b</sup>, Sabrina Bichon<sup>c</sup>, Vincent Raimbault<sup>d</sup>, Jean-Luc Lachaud<sup>a</sup>, Sébastien Gounel<sup>c</sup>, Corinne Dejous<sup>a,●</sup>, Alexander Kuhn<sup>e</sup>, Claudine Boiziau<sup>f</sup>, Nicolas Mano<sup>c</sup>, Simon Hemour<sup>a,*</sup>

## ARTICLE INFO

**Keywords:**

Implantable biosensor

Miniaturization

Wireless communication

Biofuel cell

Continuous monitoring

Zeropower

## ABSTRACT

The continuous monitoring of physiological parameters can provide crucial information to better understand and treat health disorders more efficiently. However, because the size of implantable sensors remains generally too large, numerous parts of the body can still not be accessed. We propose here to combine in a synergistic way the two most volumetric functions of a typical sensing device, i.e. the energy supply and the communication unit, to reduce the overall device footprint. The electrode of a biofuel cell (BFC), whose role is to deliver DC current, is merged with the antenna electrode, whose function is to provide radiofrequency alternating current, transforming *de facto* the BFC electrode into an antenna. This first proof-of-concept BFC, acting at the same time as an energy source and an antenna, enables wireless communication of glucose concentration over macroscopic distances. A novel dielectric screening strategy is also proposed as a trade-off between the analyte diffusion (BFC function), and electric field preservation (wireless communication function) to optimize the overall efficiency. Finally, we demonstrate that the integrity of the proposed device can be maintained after subcutaneous implantation in rats. These results pave the way for wireless miniaturized implantable devices for continuous monitoring in biological tissues.

## 1. Introduction

The rapid technological advances in continuous monitoring of bio-physical and biochemical parameters are part of the fourth healthcare revolution, known as Healthcare 4.0 [Naghdi, 2023; Gupta, 2023]. A wide range of medical applications benefit from the recent progress in implantable biosensors, such as cardiovascular diseases, neurological monitoring, intraocular pressure monitoring, metabolic diseases or cancer [Yogev, 2023]. Implantable biosensors may be designed to provide biophysical information (e.g. local temperature increase associated with infection, cardiac motion, neuronal activity, arterial restenosis) or

to detect biomarkers (e.g. blood CO₂ concentration, local hypoxia in the microenvironment of tumors, glucose concentration in the context of diabetes).

Beyond the measurement challenges, the issue of biocompatibility of implantable devices is crucial. Indeed, the foreign body reaction (FBR) begins as soon as the device is in contact with the host tissue, and is another obstacle to overcome. It takes place in successive overlapping phases, as described in many reviews [Anderson, 2008; Klopfleisch and Jung, 2017]. The main parameters are the wound due to the surgical procedure and device implantation, as well as the device's size, weight, surface properties and stiffness. In the field of implantable biosensors for

This article is part of a special issue entitled: Continuous Monitoring Biosensors published in Biosensors and Bioelectronics.

* Corresponding author.

E-mail addresses: corinne.dejous@ims-bordeaux.fr (C. Dejous), simon.hemour@u-bordeaux.fr (S. Hemour).

https://doi.org/10.1016/j.bios.2025.117970

Received 20 May 2025; Received in revised form 5 August 2025; Accepted 5 September 2025

Available online 5 September 2025

0956-5663/© 2025 The Authors. Published by Elsevier B.V. This is an open access article under the CC BY license (http://creativecommons.org/licenses/by/4.0/).

<footnote><sup>a</sup> Univ. Bordeaux, Bordeaux INP, CNRS, IMS, UMR 5218, 351 Cours de la Liberation, Talence, 33400, France</footnote>

<footnote><sup>b</sup> Instituto de Ingeniería Eléctrica, Facultad de Ingeniería, Universidad de la República, Montevideo, Uruguay</footnote>

<footnote><sup>c</sup> Centre de Recherche Paul Pascal (CRPP), CNRS UMR 5031, Univ. Bordeaux, 115 Avenue du Docteur Schweitzer, 33600, Pessac, France</footnote>

<footnote><sup>d</sup> LAAS CNRS, CNRS, UPR 8001, 7 avenue du Colonel Roche, 31400, Toulouse, France</footnote>

<footnote><sup>e</sup> Univ. Bordeaux, CNRS, Bordeaux INP, ISM, UMR 5255, Site ENSMAC, 351 Cours de la Liberation, F-33400, Talence, France</footnote>

<footnote><sup>f</sup> Univ. Bordeaux, INSERM, BioTis, U1026, 33000, Bordeaux, France</footnote>
continuous monitoring, fibrous encapsulation limits the interaction with the surrounding tissue (e.g. inhibition of analyte and ion diffusion), resulting in an overall loss of sensitivity. In addition to an optimization of the biomaterials used for sensing devices [Khan, 2024], biosensor miniaturization can contribute to minimize FBR [Wang, 2015].

The other main challenge is related to the electrical functions, especially in terms of energy and data acquisition. Today's wireless implantable biosensing systems are build, both conceptually and physically, as a combination of elementary building blocks, i.e. sensing electrode, controller, communication antenna, energy storage and conversion unit. While electronic circuits can be miniaturized, two critical functions have specific sizes that are problematic for implantation and FBR as they cannot be shrunked due to physical limitations: 1) The antenna, on which the implant relies for communicating through the body, is a device that converts high frequency alternating electric current into radio waves. It cannot be scaled down beyond a size related to its wavelength, otherwise the communication range is compromised [Kim, 2004]. Additionally, an implantable antenna is typically protected against electromagnetic tissue losses by encapsulating it in a dielectric layer (usually silicone), which supports the implantable medical device operations (sensor signal acquisition, processing, high frequency current generation ...) and usually occupies the major part of the volume. 2) Energy reservoirs, as batteries, can be replaced by enzymatic biofuel cells (BFC). These energy conversion small devices can provide sustained power converted from compounds naturally present in body fluids without the need to carry electrolyte [Mano, 2018; Xiao, 2019]. However, after 40 years of research and engineering, we have reached physical limits of miniaturization: reducing further the size of these two

blocks, would reduce the performance of the implantable system. Until now, BFC have only been used as energy sources to power sensors or develop self-powered biosensors [Mano, 2018; Xiao, 2019]. In all these cases, communication has occurred via wires [Zebda, 2011], Bluetooth [Shitanda, 2021], radio transmittance [MacVittie, 2015], or RFID technology [Di Trocchio, 2021].

To our knowledge, this work reports the first example of a BFC acting at the same time as an energy source and an antenna, enabling communication of glucose concentration. This combination of BFC and antenna in the same physical object will reduce the volumetric footprint of the implant, thus limiting the FBR. There are various sensing applications that could benefit from this joint implementation (e.g. glucose monitoring, lactate sensing, neurotransmitter detection). Indeed, the very process of electricity generation in the BFC is closely linked to the concentration of various analytes present in the tissue. The sensing information could then be communicated outside the body through the same electrodes used for energy harvesting.

Practically, the proposed strategy implies merging the metal of the BFC electrodes and the metal of the antenna, transforming *de facto* the BFC electrode into an antenna. This strategy is however inducing a new challenge: as the electrodes need a direct contact with the tissue containing the enzymatic substrates (glucose, oxygen), it will affect the wireless transmission range. Indeed, if not encapsulated, the antenna efficiency will be reduced because its electric field lines are not prevented from penetrating the tissue, thereby inducing dielectric losses.

In this article, we describe two complementary techniques that aim at reducing the total volume of an implantable biosensing system (Fig. 1) while maintaining its performance. Since gold solenoids can act as

<figure><img src="image_3.png" /><figcaption>**Fig. 1.** Principle of the hybrid implanted transducer for continuous wireless monitoring in biological tissues. (a) A BFC uses an Anode and a Cathode to generate DC current from the biofuel present in the body (e.g. glucose and oxygen). (b) An implantable BFC is usually coupled to a sensor, a controller and an antenna to build a wireless sensor. (c) The first novelty presented here is the transformation of the electrode into an antenna, so that the implantable system could gain in volume. (d) The second novelty is to propose an internal core to limit dielectric losses, and therefore provides gain in antenna efficiency. (e) This core allows both radiofrequency operation and bioelectrochemical operation.</figcaption></figure>
antennas, this topology is implemented to validate our strategy. The first effort described hereafter is to reduce the electromagnetic tissue losses by associating the solenoid with a solid core, made out of dielectric material. For this purpose, a new screening strategy is described, serving as a trade-off between the analyte diffusion (BFC function), and electric field preservation (wireless communication function); measurements confirm the strategy effectiveness. The second step demonstrates the BFC antenna concept: solenoids are wrapped around an optimized dielectric core and functionalized with a combination of enzymes and redox polymer. The resulting device is acting as an anode (with mediated glucose oxidase activity) and cathode (with mediated bilirubin oxidase operation), thus demonstrating the possibility to merge the functions of a BFC electrode with those of an antenna. In addition, we show that the dielectric core preserves the device integrity when subcutaneously implanted in rats. In the future, the electronic components will be installed inside the dielectric core, thus contributing to reduce the overall size of the device by half.

# 2. Materials and methods

## 2.1. Electromagnetic high frequency simulations

The general shape of the electrode-antenna is inspired from past BFC works that were using fiber-type electrodes [Di Trocchio, 2021]. For the present work, a similar 250 µm thick gold wire has been winded into a solenoid shape with a diameter of 3 mm, compatible with rat implantation. The shape was chosen for wireless transmission via magnetic coupling, enhancing efficiency by reducing electric field penetration in tissues found in radiative-type transmission [Pérez-Nicoli 2021]. The electromagnetic behavior of the coil structure has been simulated with a 3D Finite Element Method engine (HFSS, Ansoft, Pittsburg, PA). The tissue environment was modeled as a cube extending over the protected stripline area, used to further retrieve the signal from the immersed coils (SM1, Fig. SM1-2 (a-b)). The electric field map was computed on a section through the axis of the coil.

**Material properties.** The parameters of body tissues (body fluid, fat, muscle, grey matter) were taken from [Andreuccetti, 1997]. Standard HFSS gold properties were used. A silicone rubber (Dow Corning 3145 RTV MIL-A46146, $e_r = 2.83$, $\tan\delta = 0.0001$) has been used to waterproof the connection of the immersed coils to the FR4-based stripline (SM1, Fig. SM1-2 (c)). The dielectric properties of the core material have been measured over a range from 100 kHz to 100 MHz. For this measurement, a sample was 3D-printed with the DL260 resin (DigitalWax®), then characterized in a capacitive measurement cell, showing slowly decreasing real and imaginary parts of the relative permittivity over the range 20–100 MHz (SM2), with respective values of $3.383 \pm 0.017$ and $0.0890 \pm 0.0011$ over the range $50 \pm 2$ MHz.

**Quality factor and Figure of Merit estimation.** To quantify the capability for efficient wireless transmission, while separately studying the transmitting (TX) and receiving (RX) structures, the Figure of Merit (FoM) defined in [Cheng, 2017] is used, in addition to the quality factor Q (SM3). The TX FoM is defined as:

$$F_{TX} = k_{TX}^2 Q = \frac{(\mu_{eff} S_{TX})^2}{L} Q \quad (1)$$

where $\mu_{eff}$ is the effective magnetic permeability, that can be set equal to $\mu_0$ in tissue environment. L is the self-inductance of the solenoid, $S_{TX}$ is the effective area where the magnetic flux is calculated, that can be approximated as the product of the coil section with the number of turns, thus $S_{TX} \approx N\pi r_c^2$. The electrical matching between the coil and the load circuit can be discarded, being not part of this study. The magnetic permeability can also be removed since it is constant in all the designs. The FoM can then be rewritten as:

$$F_{OM} = \frac{(N\pi r_c^2)^2 Q}{L} \quad (2)$$

expressed in m⁴/H. The quality factor is computed as: $Q=Im(Z)/Re(Z)$, where Z is the coil input impedance. By using the low-frequency impedance value, the apparent inductance can be calculated as $L = Im(Z)/(2\pi f)$.

## 2.2. Dielectric setup and measurements

After characterizing the DL260 resin and designing the dielectric core shape, the proposed structure has been printed by stereolithography using a DWS Systems 029J + printer (SM4, Fig. SM4-1, left).

Then, the 250 µm gold wire has been winded inside the groove of the core to form the solenoid-shaped electrode (SM4, Fig. SM4-1, right). Two such electrodes were then connected each to a stripline, and waterproofed with a silicone sealing layer, leaving the coil part in contact with the external medium. For the sake of comparison, a similar core-less solenoid has also been fabricated.

Preliminary measurements were performed using a Vector Network Analyzer (VNA, Keysight E5061B), with the structures dipped inside a Phosphate Buffered Saline (PBS) solution which mimics the body fluid (SM5). To mimic muscle tissue, the structures were later inserted inside bovine minced meat, at room temperature (21–23 °C) (SM6, Fig. SM6-1).

## 2.3. Implantation of devices in rats

All animal studies were carried out in accredited animal facilities at the University of Bordeaux (accreditation number: A33-063-917) and were approved by the French Ministry of Higher Education, Research and Education (APAFIS 16443-2018072016398155 v4). Details of implantation of linear (3 rats) or solenoid (5 rats) electrodes and their integrity monitoring by X-ray imaging are described in SM7.

## 2.4. BFC-antenna simultaneous operation setup

### 2.4.1. Antenna-to-antenna wireless link

The wireless link was set up through the glass of a beaker. On the inner side, one of the two solenoid-shaped electrodes (3 mm diameter) was used as a transmitting antenna (TX). Note that those two electrodes were immersed in PBS solution and were simultaneously operating as BFC Cathode and Anode. On the outer side of the beaker, a resonating coil (4 cm diameter) with an operating frequency of 28.25 MHz was installed to catch the magnetic field produced by the immersed electrode-antenna. The distance between the resonating coil and the antenna was typically 0.5–2 cm.

The transmitting antenna (immersed electrode) was connected to port 1 of the VNA, while the receiving antenna (external, large coil) was connected to port 2 to measure the wireless transmission losses. A bias tee has been used to route the DC current towards the DC load and to route the high frequency currents from the VNA to the antenna electrode. The bias tee internal DC resistance of 0.36 Ω is negligible if compared to the load of the circuit.

### 2.4.2. RF signal measurement

The VNA has been implemented to test the high frequency wireless transmission with -10dBm signals from 5 to 50 MHz. DC voltage measurements were done during the transmission to validate the dual operation (wireless link and DC power production). The BFC DC power measurements were performed by means of a potentiostat, with the VNA disconnected to avoid measurement artefacts resulting from ground loop effects.

The wireless communication link (WCL) performance was estimated from the scattering parameters $S_{xy}$ measured with the VNA:
$$WCL = \frac{P_r}{P_t} = \frac{|S_{21}|^2}{1 - |S_{11}|^2} \quad (3)$$

where $P_r$ and $P_t$ are, respectively, the received and transmitted power. This method discards the mismatch between the VNA port and the small coils, not object of this study.

### 2.4.3. Enzyme-modified electrodes

The two solenoid-shaped electrodes were used as working electrodes. The preparation of the cathodic and the anodic composite hydrogels, based on osmium based redox polymers, and anodic (glucose oxidase “GOx” from A. niger) and cathodic (Bilirubin oxidase “BOD” from M. oryzae) enzymes, is described in SM5. Each coiled gold electrode was immersed overnight in a 50 µl drop of its respective bio-electrocatalyst at 4 °C in a 100 % humidity atmosphere. Before measurement, electrodes were let dry for 3h at room temperature.

### 2.4.4. BFC characterization

After enzyme-modification, the two coils/electrodes were immersed in PBS (20 mM of NaPi, 140 mM NaCl) to characterize both, the EM and electrochemical behavior at 37 °C. Electrodes were first characterized under Argon at 50 mV/s. Bioelectrocatalysis was measured at 5 mV/s after addition of 40 mM of glucose to the $O_2$ saturated PBS buffer.

## 3. Results and discussion

Theoretical background and equations are given in SM3.

### 3.1. Simulation of a new dielectric core-based electrode

#### 3.1.1. Dielectric core-based structure design

To communicate through the body, low frequency (MHz) magnetic fields are known to be more efficient than traditional microwave frequencies (GHz). The most effective implantable antenna are therefore loops, spirals or solenoids. However, at resonance, there is an exchange of energy in the vicinity of such “coil” antenna, between magnetic energy, necessary to transmit the signal, and the electric field, which acts as a flywheel for the antenna operation. This electric field is maximum inside the solenoidal structure and between its turns [Chute, 1981] (Fig. 2a). Therefore, the quality factor and WCL performance of the antenna may be compromised when the volumetric electrical energy density is deployed in a lossy dielectric environment, such as body tissues. This is why conventional implanted antennas are always coated to prevent the electric field from penetrating the lossy tissues (Fig. 2b).

However, our implanted electrode-antenna has to be in contact with the tissues for its electrochemical operation. It is therefore proposed to carefully shape a low-loss dielectric material to fill the critical volume and avoid the penetration of E field into the tissues, while keeping the contact between the electrode/coil and the tissue (Fig. 2c).

#### 3.1.2. Parametric analysis of the dielectric core

The threaded-furrow geometry has an outer radius $r_s$ and a height $h_s$, and an inner radius and a height $r_c$ and $h_c$ respectively (Fig. 2c, top). The simulated electric field map evidences a large proportion of the field confined inside the structure, almost as much as with the fully coated coil case (Fig. 2b and c, bottom).

The role of the key dielectric core geometric parameters is analyzed by electromagnetic simulations before optimization (while maintaining constant coil dimension $r_c = 1.5$ mm, $h_c = 3$ mm, $N = 3$ and $r_w = 125$ µm, constant dielectric core height and losses, respectively $h_s = 4$ mm and $\tan\delta = 0.001$). The surrounding medium is also fixed with $\epsilon = 75$ and $\sigma = 1.5$ S/m over the whole frequency band, simulating a high-dielectric lossy medium such as the human tissues.

As expected, an increase of the core radius $r_s$ and thus the isolation thickness $\Delta d$ led to an increase of Q (Fig. 2d). This is because the electric

field was less distributed in the lossy volume, reducing thus the volumetric losses in the tissue. Moreover, the increase of the core dielectric permittivity $\epsilon_r$ resulted in an increase of the parasitic capacitance, and thus of the electric energy stored, leading to a lower Q. This effect was more important as $r_s$ increased. Therefore, this showed that the dielectric core should have losses and dielectric constants as low as possible, in order to maximize the performances of the structure.

The configuration with $\Delta d = 300$ µm, $\epsilon_r = 1$ and $\tan\delta = 0.001$, was then compared with the same bare coil and the coated one (Fig. 2e). As expected, the quality factor of the structure is reduced when the electric field is fully exposed to the lossy tissues (topology of Fig. 2a). Conversely, the quality factor increases with full silicone encapsulation (case of Fig. 2c). Moreover, the frequency corresponding to the maximum value of Q shifted to higher values, indicating a decrease of the parasitic capacitance as the electric field leakage into the tissues was reduced (confined inside the low-loss dielectric structure).

While varying $\Delta d$ for the three structures and computing for each point the maximum value of Q, the previous behavior was confirmed, with the dielectric core-based coil between the bare and the silicone coated ones (Fig. 2f). It is interesting to observe the linear behavior of the dielectric core curve. More simulations related to the stored energies and dielectric losses are presented in SM8.

#### 3.1.3. Simulation validation

Preliminary measurements were conducted with a device based on a dielectric core mounted on a specific fixture immersed in PBS (Fig. 2g and h). Comparable behaviors of Q for the simulated and measured full structures can be observed, validating thus the simulation process and confirming the performance increase when the dielectric core was introduced despite a slight decrease attributed to the test fixture (Fig. 2i). The slight difference in the maximum values could be explained by the difference between the media used for the measurements and the simulation. In particular, the PBS is more lossy than the body fluid, resulting in more energy dissipated and thus, in a lower Q value. This is confirmed by the frequency shift that affected both the bare coil and the dielectric core-based coil results.

#### 3.2. Optimization of the dielectric core-based antenna at constant volume

To further push the study, it was important to consider the entire implanted volume during the optimization, unlike most of the work in the literature. Indeed, this parameter is important depending on the application, and because tissue displacement impacts the inflammatory response of the body. This section focuses on the TX solenoid design towards this aim.

Based on the previous design, the whole volume constraints were set to $r_s^{max} = 3.5$ mm and $h_s = 4$ mm. $h_c$ and $r_w$ were also kept constant, respectively 3 mm and 125 µm. Geometries with different values of $\Delta d$ were considered (Fig. 3a). For negative values of $\Delta d$, the coil radius was set to the maximum value to fit the whole constant volume ($r_c = r_s^{max} - r_w$) and decreasing the dielectric core radius. As a consequence, as $\Delta d$ increased, the coil was more confined inside the full-size dielectric core, thus also the electric field, leading to an improved quality factor. On the other hand, this gain was expected to be counterbalanced by the coil radius reduction, decreasing the effective area. Similarly, increasing $N$, thus the self-inductance, was expected to decrease the Q factor, but at the same time to increase the effective area. Therefore, the presence of optimum values of $\Delta d$ and $N$ that maximize the FoM was expected.

#### 3.2.1. Simulation: influence of the design parameters and tissue environment

The structure was simulated by sweeping $\Delta d$ and $N$ from -1000 µm to 1250 µm and from 2 to 8, respectively. For each design, the maximum value of FoM as a function of the frequency was computed, aiming at finding the optimum trade-off.

The simulation results showed that the FoM was more affected by the
<figure><img src="image_4.png" /><figcaption>**Fig. 2.** **Proposed dielectric core:** design (top) and simulated electric field cross-section map (down) of the bare coil (a), the conventional coil with silicone coating (b) and the new coil with the dielectric core (c). Simulation results: evolution of the maximum attainable quality factor as a function of isolation thickness $\Delta d = r_s - r_c$ (solid line corresponding to core material of relative permittivity of 1 while the dashed line corresponds to a relative permittivity of 120) (d), comparison between the bare coil, the silicone coated one and with the dielectric core at $\Delta d = 300$ µm (e), and comparison of the structures as a function of $\Delta d$, at the frequency of the respective maximum (f); 3D view of the coil mounted on its fixture (g), prototype measurement in PBS (h) and preliminary comparison of simulated (in body fluid) and measured (in PBS) quality factor of the coil mounted on its test fixture with and without the dielectric core ($\Delta d = 130$ µm) (i).</figcaption></figure>
<figure><img src="image_5.png" /></figure>

<figure><img src="image_6.png" /></figure>

<figure><img src="image_7.png" /></figure>

<figure><img src="image_8.png" /><figcaption>**Fig. 3.** The coil associated with the dielectric core structure with different $\Delta d$ at constant volume. Simulations: Sketch of the different structures, the other parameters are: $r_s^{max}$ = 3.5 mm, $h_s$ = 4 mm, $h_c$ = 3 mm, $N$ = 4, $r_w$ = 125 µm (a); Simulation results of a coil associated with a dielectric core structure at constant external volume: FoM (maximum value) in muscle tissue as a function of $\Delta d$ and $N$ (b), FoM and $Q$ at constant volume in fat with $N$ = 5, muscle, grey matter and body fluid tissue environment with $N$ = 4 (c). Experimental validation: Quality factor of the simulated and experimental structures in bovine minced meat, modeled as a cube of muscle tissue extending over the protected stripline area, as a function of frequency (d) and synthesis of maximum $Q$ and FoM as a function of $\Delta d$ (e) (SM6).</figcaption></figure>
variations of $\Delta d$ than by the variations of $N$ (Fig. 3b). When the coil radius was constant ($\Delta d < 125$ µm) and only $N$ was changed, a maximum value of the FoM could be observed at $N = 4$ for muscle tissue. This could be explained by the initial increase of the effective area proportional to $N$, but, beyond an optimum value of $N$, this increase did not compensate the decrease of $Q$ and the simultaneous increase of the self-inductance.

Four different tissue environments, subject to implantation, were taken into account: fat, muscle, grey matter and body fluid, from a low-epsilon low-losses best case scenario (fat), to the high-epsilon high-losses worst case (body fluid). Similar FoM behaviors were obtained (Fig. 3c), with maximum values at $N = 4$ for muscle, grey matter and body fluid, $N = 5$ for fat with lower losses which slowed down the decrease of $Q$ as $N$ was increased. For all tissues, with $\Delta d < 0$, the FoM and $Q$ slightly increased with dielectric structure dimension. A maximum of the FoM was observed at $\Delta d = r_w$, since in this case the coil was completely covered while the effective surface was maximized. For $\Delta d$ beyond $r_w$, $Q$ increases faster as expected, but, at the same time, $r_c$ and thus the effective surface $S_{TX}$ of the coil start to decrease. Since the FoM has a quadratic dependence of $S_{TX}$, the increase of $Q$ with $\Delta d$ is not high enough to compensate the decrease. Consequently, the FoM decreased rapidly with the decrease of $r_c$. As a result, for all the tissues, the best geometry was that with the coil covered and the maximum effective area ($r_c = r_s^{max} - r_w$).

### 3.2.2. Measurements: electromagnetic characterization

Structures based on a set of 3D-printed dielectric structures, with $N = 4$ and different values of $\Delta d$: -125, 125, 500 and 1000 µm, placed in minced bovine meat close to muscle tissue, were characterized, as well as a bare coil with the same external dimensions for appropriate comparison (SM6).

When considering the full frequency range for each structure, a very good agreement between measurements and simulation was observed (Fig. 3d). For $\Delta d = 125$ µm, the maximum measured $Q$ factor was slightly lower than the simulated one. This could be attributed to resolution limits of the 3D-printer for the thread definition, resulting in incomplete covering of the coil wire, thus exposing the electric field to the lossy tissue. This was consistent with very similar measured $Q$ factors for $\Delta d$ of -125 and 125 µm, in contrast to simulation. On the other hand, a small increase of $Q$ was witnessed at $\Delta d$ of 500 and 1000 µm, probably because, even if accurately squeezed, the meat was not completely entering inside the dielectric structure thread. Furthermore, the frequency of the maximum $Q$ was expectedly increased as $r_c$ was reduced, as it could be observed for these last two $\Delta d$ values.

From these results, the $Q$ maximum values are reported on Fig. 3e, as well as the FoM. The value $\Delta d = -1625$ µm corresponds to the bare coil case. As described earlier, the measured and simulated $Q$ were in good agreement, and increased with the dielectric core, compared to the bare coil case. The resulting FoM also exhibited similar behaviors, though the measured values were slightly underneath the simulated ones. This was attributed to measured apparent inductances slightly greater than the simulated ones. The difference, less than 1 nH, got smaller at high $\Delta d$ values, where the two curves were almost overlapping.

To conclude, the measurements confirmed that the best shape to ensure good WCL performance, together with good contact between the coil/electrode and the tissues, was the core with the same diameter as the outer one of the solenoid, while keeping the coil as large as possible within an acceptable implanted volume. With such an optimized structure, a higher sensitivity of the electrodes is expected.

### 3.3. Implantability and in vivo structure stability

One major condition for using this sensor in continuous medical monitoring is its integrity within the patient's body. In a preliminary in vivo study performed on rats [Di Trocchio, 2021], linear electrodes, functionalized with the GOx, were subcutaneously implanted. The transcutaneous connection with an antenna immobilized on the rat back

allowed a continuous monitoring of the open circuit voltage during approximately 20 h. But as seen in Fig. 4a, X-ray imaging showed that the rat movements caused significant bending and electrode breakage for all electrodes after only one or two days. In the present study, we hypothesized that the design of the coil immobilized around the dielectric core might preserve the coiled structure and the electrode integrity. 10 devices were subcutaneously implanted on 5 rats (2 per rat) over a 21-day period. As shown on Fig. 4b, the gold electrodes were clearly detectable by X-ray imaging and neither bending nor breakage were observed. This demonstrated that the addition of an internal dielectric core to increase the WCL performance was also stabilizing the electrode structure. After euthanasia and dissection at Day 21, histology analysis confirmed that the gold electrodes remained tightly associated with the dielectric core (data not shown), which supports the hypothesis that such coiled structure could be maintained for longer.

### 3.4. Ability of the BFC to behave as an antenna

Having validated the core structure to improve the efficiency of wireless communication and stabilize the gold electrode after implantation, we tested the ability of this device to act as both, a BFC and an antenna. The experimental setup is represented in Fig. 5a, with the two solenoid-shaped electrodes immersed in PBS solution, one of them being used as antenna transmitter. The receiving coil is positioned outside the beaker, aligned with the transmitting electrode-antenna. High frequency measurements are simultaneously performed with the VNA while the DC power is flowing in the decade resistance box placed on the VNA and serving as adjustable load resistance.

The individual characterization of each electrode-antenna has been undertaken. The CVs in the absence of substrate (Fig. 5b) are characteristic of polymer-bound osmium complexes with respectively an apparent redox potential of +350 mV vs. Ag/AgCl for the cathode (blue curve) and 10 mV vs. Ag/AgCl for the anode (red curve). Fig. 5c shows the catalysis upon addition of 40 mM glucose to the oxygenated PBS buffer, reaching 60 µA at the anode and 25 µA at the cathode (in vivo, glycemia can vary between extreme values of 1.0 and 45 mM, with normal fasting blood glucose concentrations between 3.9 and 5.6 mM). The open circuit voltage (OCV) of the BFC was 0.52 V, which is defined as the difference between the open circuit potentials (OCPs) of the anode and cathode, determined by the redox potential of the corresponding enzymes at zero current [Mano, 2018; Xiao, 2019]. The resulting power curve reaches 5 µW at 0.24 V (Fig. 5d). Such a power is sufficient for widely used systems like pacemakers and cardioverter-defibrillators [Chandrakasan, 2008]. It also meets the requirements of body-area monitoring applications with low duty cycles, where the system operates in sleep mode most of the time, maintaining only a low-power clock (<1 µW), and periodically wakes up to sense relevant physiological variables [Chandrakasan, 2008].

The high frequency instruments and bias tee wiring are shown in Fig. 5e along with their associated WCL results (Fig. 5f). The wireless transmission losses have been quantified as 12 dB and 21.5 dB, respectively at 0.5 cm and 2 cm TX/RX distances, increasing with the distance due to the nature of the magnetic coupling link. Such losses at a 2 cm distance are similar to those reported for standard commercial RFID systems [Ahson, 2017] at around 10 cm. Considering the small size of the BFC-antenna (3 mm × 4 mm) and its direct contact with tissue, the obtained losses are completely reasonable. Furthermore, typical acceptable transmission losses in practical communication scenarios can exceed 100 dB, thus demonstrating the feasibility of establishing a reliable communication link. Moreover, PBS buffer has larger high frequency losses than muscle tissues. It is therefore expected that under implantation conditions, the device should allow better communication distances.
<figure><img src="image_9.png" /><figcaption>(b) Electrodes associated with the dielectric core</figcaption></figure>

<figure><img src="image_10.png" /><figcaption>**Fig. 4.** X-ray imaging of gold electrodes (diameter 250 µm) at different times after subcutaneous implantation. (a) Linear electrodes were implanted in rats (arrows). Due to the rat's movements, the electrodes might be bent (lightning bolt) as soon as after 1 day (Rat 1) or even broken (star) (Rat 2). Representative pictures of 3 rats implanted with linear electrodes. (b) Coiled electrodes associated with the dielectric core were implanted (2 per rat, arrows). Their integrity was assessed over a 21-day period.</figcaption></figure>

**4. Conclusion**

The feasibility of merging a bioelectrochemical transducer and an antenna, by shaping the transducing electrodes as coils, was

experimentally validated for continuous monitoring in PBS. Then, a novel structure based on a dielectric core was developed in order to improve the wireless transmission capability of the bare coil, while keeping the contact between the tissue and the conductor for sensing or
<figure><img src="image_11.png" /><figcaption>**Fig. 5.** WCL performance improvement of the BFC-antenna with the dielectric core. Image of the antenna-electrode experimental setup (a). Electrochemical results: Cyclic voltammograms (CVs) of the mediated bilirubin oxidase cathode (blue curve) and mediated glucose oxidase anode (red curve) in the absence of substrate, under Argon in a PBS buffer, 37 °C, 50 mV s⁻¹ (b) or in the presence of substrate: 40 mM glucose, 37 °C, under 1 atm O₂ in a PBS buffer at 5 mV s⁻¹ (c) and curves of the resulting electrical current and power delivered from the BFC (d). Scheme of the equivalent electrical circuit (e). Transmittance of the RF link (anode/receiving coil) with respect to frequency at distances 0.5 mm and 2 mm (f). (For interpretation of the references to colour in this figure legend, the reader is referred to the Web version of this article.)</figcaption></figure>

excitation function. This could avoid adding wired external electrodes and minimizes the implant volume, together with enhanced mechanical properties for implantation. The structure was designed and simulated in various tissue environments. An optimization design routine, aiming to fit to a defined available whole volume, was presented. In particular, a FoM was used to maximize the TX characteristics, without taking into account the RX placed outside the body. The best results were obtained when the coil was flush with the edge of the dielectric core, with 5 turns for fat and 4 turns for the other tissues. In all cases an improvement of Q and of the FoM with respect to the bare coil case was observed. A set of prototypes was fabricated and measurements in minced meat showed a good agreement with simulations, with Q factor improvement compared

to the bare coil case. We demonstrated that such a coiled geometry was an asset to maintain mechanical integrity upon implantation. Finally, as a proof of concept, we used such mechanically stable coiled electrodes to report, for the first time, an enzymatic BFC acting as both, electrode and antenna.

Not detailed here but exposed in (SM9), such a structure also allowed the wireless sensing and the discrimination of a healthy muscular tissue surface and one having a higher concentration of collagen, as produced during fibrosis development. This opens the possibility to design a wireless impedance monitoring system, capable to monitor the inflammatory response of the body to an implant, by designing a dedicated electronic circuit.
However, this work has some limitations, that can be addressed in the future to improve the overall device properties for in vivo implementation. Bare gold electrodes could be replaced by macroporous electrodes to increase the effective surface area [Zhang, 2019], thus increasing the signal emitted by the antenna for an interrogation distance beyond a few centimeters. For the cathode (functionalized with BOD), oxygen releasing biomaterials developed for tissue engineering [Du, 2025] could provide oxygen to the enzyme. Furthermore, the device could be embedded in a hydrogel mitigating biofouling and FBR for an extended life-time of the BFC [Vaddiraju, 2012; Lee, 2020], which should not degrade the antenna function compared to the body tissue environment. Lastly, new designs with alternative electrode materials such as carbon-based ones, e.g. carbon nanotubes with high electroactive surface area [Zebda, 2011; Chakraborty, 2024; Mishra, 2021], could be also investigated, considering cost-efficiency and performance on both the BFC and the antenna side, provided biocompatibility is still ensured.

To conclude, the proposed coil-on-dielectric core structure has evidenced the capability to merge an electrode-based biosensor and an antenna function for continuous monitoring in biological tissues, paving the way for wireless miniaturized implanted devices. Some improvements are studied within undergoing project (https://anr.fr/Project-ANR-21-CE19-0015), and the approach can be also transposed in terms of merged functions by teams involved in any types of BFCs, enzymatic or based on other functional materials. Therefore, it is reasonable to imagine that in the future, devices such as miniaturized pacemakers could be developed that are powered (BFC), communicate externally, and even stimulate, all through the same two electrodes.

### CRediT authorship contribution statement

**Luigi Di Trocchio:** Writing – original draft, Visualization, Validation, Methodology, Investigation, Formal analysis. **Pablo Pérez-Nicoli:** Writing – review & editing, Writing – original draft, Visualization, Validation, Investigation, Formal analysis. **Sabrina Bichon:** Investigation. **Vincent Raimbault:** Resources, Investigation. **Jean-Luc Lachaud:** Resources, Investigation. **Sébastien Gounel:** Visualization. **Corinne Dejous:** Writing – review & editing, Writing – original draft, Visualization, Validation, Supervision, Resources, Project administration, Methodology, Funding acquisition, Formal analysis, Conceptualization. **Alexander Kuhn:** Writing – review & editing, Supervision, Resources, Project administration, Funding acquisition, Conceptualization. **Claudine Boiziau:** Writing – review & editing, Writing – original draft, Visualization, Validation, Resources, Investigation, Funding acquisition, Formal analysis. **Nicolas Mano:** Writing – review & editing, Writing – original draft, Validation, Resources, Funding acquisition, Formal analysis. **Simon Hemour:** Writing – review & editing, Writing – original draft, Visualization, Validation, Supervision, Resources, Project administration, Methodology, Investigation, Funding acquisition, Formal analysis, Conceptualization.

### Funding

This work has received financial support from the French Government in the frame of the "Investments for the future" Programme IdEx Bordeaux (ANR-10-IDEX-03-02), and from the ANR (WEB, ANR-21-CE19-0015-01). LDT received a doctoral fellowship from LabEx AMADEUS (ANR-10-LABX-0042-AMADEUS) and Univ. Bordeaux. This work was supported by the French RENATECH network and the MultiFAB project funded by FEDER European Regional Funds and French Région Occitanie (16007407/MP0011594).

### 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

We would like to thank the animal house A1 facility (University of Bordeaux, France), and the people who work there for their help and advice.

## Appendix A. Supplementary data

Supplementary material to this article can be found online at https://doi.org/10.1016/j.bios.2025.117970.

### Data availability

Data will be made available on request.

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