**4. Light-based hearing restoration**

Based on the literature review presented in the previous section, the main bottlenecks of CIs are low accuracy and low precision of nerve stimulation methods, bandwidth scarcity and constraint capacity of RF communication techniques, and high energy consumption of both. To this end, we present two architectures capable of mitigating the effect of these limitations and even eliminating them [11, 17].

## **4.1 Optical wireless cochlear implant**

The utilization of optical wireless communications in order to develop CI transdermal optical links has been recently investigated [11], where the authors proposed a novel system architecture, termed optical wireless cochlear implant (OWCI), that improves the power and spectral efficiency as well as the reliability of the transdermal optical link. Moreover, in the same contribution [11], the capabilities and feasibility of the OWCI are evaluated and design guidelines are provided. The main comparison points between OWCIs and conventional CIs are illustrated in **Table 1**. In addition, the presented advances in the communications of CIs are in line with optical neural stimulation advances on the acoustic nerve [21, 63–67].

The unique technical contributions of the OWCI entail the establishment of a novel system model for transdermal optical links that incorporates the various design variables such as the stochastic misalignment between the receiver (RX) and the transmitter (TX), the scale of the optical components, the skin thickness, and the

*Hearing Restoration through Optical Wireless Cochlear Implants DOI: http://dx.doi.org/10.5772/intechopen.104622*


#### **Table 1.**

*OWCIs versus conventional CIs (bold fonts demonstrate the advantages).*

transmission power. The external component of the OWCI is comprised of a microphone, the TX, and a digital signal processing (DSP) unit, while the implanted one contains the RX as well as a stimulation and a DSP unit. The external DSP unit is responsible for digitizing and compressing the sound signal from the microphone into coded signals, which are then forwarded from the TX to the RX over the transdermal optical link. In the implanted component, the DSP and stimulation units transform the received signal into a series of electrical pulses that will stimulate the auditory nerve (**Figure 2**). Based on this system model, the performance of the OWCI was evaluated with regard to the SNR, channel capacity, outage probability, and spectral efficiency. The results not only validated the feasibility of the proposed architecture and provided meaningful insights that can be used as design guidelines, but also revealed the superior effectiveness and reliability of the OWCI compared to the conventional CI.

In the aforementioned architecture, the transmitted signal, *x*, is conveyed over the wireless channel, *h*, with additive noise *n*. Thus, the received signal can be written as [68–70].

$$y\_1 = Rh\mathbf{x} + \mathbf{n} \tag{1}$$

with *η* denoting the quantum efficiency of the photodiode, *R* the RX's responsivity, *v* the frequency of the photons, *q* the electron charge, and *p* the Planck's constant. It is highlighted that the channel coefficient can be expressed as *h* ¼ *hl hp*, where *hl* represents the deterministic channel coefficient caused by propagation loss, while *hp* denotes the collected power fraction due to the geometric spread from the origin of the detector and is caused from the TX-RX misalignment (**Figure 3**).

The CI channel's deterministic term can be expressed as in ([71], Eq. (10.1))

$$h\_l = \exp\left(-(\mu\_a(\lambda) + \mu\_s(\lambda))\delta\right) \tag{2}$$

where *λ* is the transmission wavelength, *δ* is the skin thickness, *μα*ð Þ*λ* is the skin attenuation coefficient, and *μs*ð Þ*λ* is the skin scattering coefficient, which can be acquired from a plethora of experimental results [72–76]. In this analysis, the term skin refers to the biological structure that consists from the stratum corneum, the

#### **Figure 2.**

*Diagrammatic illustration of the architecture of the OWCI. The OWCI captures the sound information via the microphone located outside of the human body. Afterward, it utilizes optical wireless communications to transfer it toward the receiver fixed on the cranial bone. Finally, the implanted unit stimulates the acoustic nerve by delivering the appropriate signals via the stimulation electrode.*

#### **Figure 3.**

*The effect of misalignment with regard to skin thickness. As the RX (photodiode) moves away from the TX under the same severity of misalignment, its distance from the perfect alignment conditions becomes enclosed in the TX's beam width. This phenomenon creates an equilibrium between the optimal TX-RX distance and the TX beam width under fixed misalignment conditions.*

epidermis, and the dermis [71], while both the RX and the TX touch with the inner (adipose) and outer (epidermal) side of the skin [24], and thus, the TX-RX distance is regarded equivalent to skin thickness.

The misalignment between the TX and the RX can be modeled as the stochastic channel coefficient, which expresses the collected power due to geometric spread with radial displacement *r* from the origin of the detector and can be written as

$$h\_p \approx A \exp\left(-\frac{2r^2}{w\_e^2}\right),\tag{3}$$

*Hearing Restoration through Optical Wireless Cochlear Implants DOI: http://dx.doi.org/10.5772/intechopen.104622*

which is based on the assumption that at distance *d* from the TX, the circular aperture of the transmitted beam has a radius of *R* and the spatial intensity on the plane of the RX is *wd*. In addition, *we* represents the equivalent beam waist radius and *A* expresses the collected power under perfect alignment. This approximation has been utilized in various previous works for modeling stochastic pointing errors [77, 78].

Based on this model, if we assume independently and identically Gaussian distributed horizontal and vertical displacement, it has been proven that *r* follows a Rayleigh distribution [79]. As a result, the probability density function (PDF) of the stochastic term of the channel coefficient can be written as

$$f\_{\,h\_p}(\mathbf{x}) = \frac{\gamma}{A^{\gamma}} \mathbf{x}^{\gamma - 1}, \qquad 0 \le \mathbf{x} \le A,\tag{4}$$

where

$$A = \left(\text{erf}(\beta)\right)^2, \quad \beta = \frac{\sqrt{\pi}R}{\sqrt{2}w\_d}, \quad \gamma = \frac{w\_\epsilon^2}{4\sigma^2}, \quad w\_\epsilon^2 = w\_d^2 \frac{\sqrt{\pi}\text{erf}(\beta)}{2\beta\exp\left(-\beta^2\right)},\tag{5}$$

while *σ*<sup>2</sup> denotes the variance of the misalignment.

#### **4.2 All-optical cochlear implant**

The CI implementations presented so far counterbalance either the RF scarcity that plagues the communications part of the system or the nerve stimulation limitations. To this end, the AOCI has been proposed as an architecture that converts the audio captured from the microphone into a light signal inside the external component for propagation to the cochlea [17] (**Figure 4**). This way, the AOCI counterbalances the aforementioned challenges and, at the same time, eliminates the need for an energyconsuming DSP unit in the implanted component. The AOCI not only builds upon the fruitful characteristics of the OWCI but also proposes breakthrough alterations such as the fact that it consists of only passive components, and thus, the implanted component has no power demands, which eliminates the requirement of complex power transfer policies and boosts energy efficiency. Furthermore, the AOCI utilizes optical neural stimulation, which is characterized by higher fidelity than electrical neural stimulation due to the lower spread of optical signals in human tissues. The technical advancements include the introduction of the AOCI architecture, its main building blocks, and the end-to-end system model. The AOCI takes into account channel, building block, and biological particularities [17]. Moreover, a novel tractable expression is derived for the instantaneous coupling efficiency in scenarios with misalignment fading. The feasibility of the proposed architecture is proven through the theoretical framework, which also evaluates its performance with regard to the power efficiency, the photon flux, and a plethora of design parameters that greatly influence the success or failure of the system.

Much like OWCI, the architecture of AOCI consists of the implanted and the external component, with the former located on the skull and the latter on the external surface of the skin. The external component captures the acoustic signal with a microphone, performs the necessary DSP, and converts it into the appropriate optical signal capable of generating the desired action potentials on the targeted spiral ganglion neurons. This signal is transmitted from the TX, which is a laser source, to the implanted component, where the guiding lens, the microelectromechanical device,

#### **Figure 4.**

*Illustration of the architecture of the AOCI. The all-optical nature of the AOCI resides in the combined utilization of optical wireless communications and optogenetics for stimulating the auditory nerve. Initially, the auditory neurons are sensitized to optical radiation with optogenetic techniques. Next, the sound captured from the external microphone is converted into an optical signal capable of stimulating the light-sensitive nerves, which is then forwarded to the cochlea.*

the coupling lens, and the optical fiber ensure its delivery to the appropriate place in the cochlea. Specifically, the guiding lens guides the light toward the microelectromechanical device to maximize the power of the received optical signal. Afterward, the microelectromechanical acts as a mirror that mitigates the misalignment to a degree by steering the light beam to the center of the coupling lens in order to be coupled into the optical fiber. Finally, the latter delivers the light into specific points along the cochlea based on their spectral content.

#### *4.2.1 Microelectromechanical device*

Microelectromechanical devices have been the subject of much hype during the past decade due to their adaptability as well as low cost, low weight, and small size [80–82]. In the case of the AOCI, the microelectromechanical device is required in order to account for the individuality of each patient. In particular, the AOCI is required to adapt to the particularities of the patient, such as different skin thickness and color or slightly varied orientation of biological tissues, in order to ensure uninterrupted hearing restoration. Moreover, imperfections during the implantation process can cause slight variations to the final placement of the implant. To this end, the microelectromechanical device provides an externally operated light control system by enabling the steering of the optical beam toward the coupling lens. Finally, the microelectromechanical device adjusts its optical properties and, thus, steers the beam after receiving the appropriate electrical charge that can be applied during implantation, while in normal operation, the need for adjustment is eliminated, and therefore, the microelectromechanical device operates passively [83, 84].

The signal received by the guiding lens presented in (1) is forwarded to the microelectromechanical device, which introduces a collimation gain [85].

$$\mathbf{G}\_{\mathbf{c}} = \frac{1}{\sqrt{\left(\mathbf{1} - d\_{\rm in}/f\right)^2 + z\_0^2/f^2}}. \tag{6}$$

Therefore, the updated received signal at the output of the microelectromechanical device can be expressed as

$$y\_2 = G\_c h\_l h\_p \text{x} + n.\tag{7}$$

#### *4.2.2 Coupling lens*

The coupling lens receives the optical beam from the microelectromechanical device and focuses it in the center of the optical fiber. The fact that incident light on the end of the optical fiber that arrives at a greater angle than the acceptable angle of the optical fiber is not coupled highlights the detrimental impact it plays on the maximum achievable coupling efficiency of the system. Moreover, the coupling efficiency is also affected by the dimensions of the coupling lens and the diameter of the optical fiber with its maximum value being in the order of 80% [86].

The coupling lens captures the optical signal that is reflected by the microelectromechanical device and couples it into the optical fiber. The signal that successfully enters the optical fiber can be written as

$$y\_3 = \eta G\_c h\_l h\_p \text{x} + n \tag{8}$$

with the coupling efficiency given by

$$\eta = \left(\frac{3.83\sqrt{2}Da\_0}{1.22\lambda F} \exp\left(-\frac{r^2}{a\_0^2}\right) \Psi\_2\left(\mathbf{1}; 2, \mathbf{1}; -\frac{3.83^2 D^2 a\_0^2}{1.22^2 \lambda^2 F^2}, \frac{r^2}{a\_0^2}\right)\right)^2. \tag{9}$$

In Eq. (9), *ω*0, *F*, *D*, and *ρ* denote the optical fiber mode field radius, the focal length, the focusing lens diameter, and the radial distance on the focal plane, respectively, while it becomes obvious that the achievable coupling efficiency is dependent on the optical fiber's mode field radius, the coupling lens's focal length and diameter, as well as the intensity of misalignment and the transmission wavelength.

#### *4.2.3 Optical fiber*

The optical fiber of the AOCI takes the place of the electrode array of the conventional CI. The incident optical signal must be delivered to specific locations alongside the cochlea in order to generate action potentials at the targeted spiral ganglion neurons that are responsible for the appropriate sound frequency. To achieve the required performance, the optical fiber proposed in the AOCI architecture propagates the optical signal through its single-mode core with a Gaussian beam profile in the output [87, 88]. Furthermore, despite the fact that state-of-the-art conventional CIs can be equipped with a

maximum of 20 electrodes, due to the limited spatial resolution of electrical neural stimulation, the sound perceived by the patient has the fidelity of eight functional electrodes [40]. In addition, to achieve speech and music perception under suboptimal noise constraints, CI must house approximately 32 electrodes, which is also the goal of the AOCI [89, 90]. Therefore, tilted fiber Bragg gratings (FBGs) were introduced in the AOCI architecture that enable light delivery in various locations alongside the optical fiber [91, 92]. These FBGs are located in the core of the optical fiber, along the propagation direction, with a periodic variation of the refractive index. These components have low insertion loss, low complexity structures, and high wavelength selectivity. Specifically, tilted FBGs allow a limited number of wavelengths to penetrate them by filtering the incident optical signal based on its spectral content and, at the time, redirecting it based to their angle [93, 94].

When the optical signal travels through the optical fiber, it attenuates due to the curvature of the optical fiber and the existence of FBGs, and therefore, the emitted signal can be expressed as

$$\mathcal{Y}\_4 = k \eta G\_c h\_l h\_p \text{x} + n,\tag{10}$$

where *k* denotes the propagation efficiency, which is limited to 0*:*14 dB*=*90<sup>∘</sup> by the strong optical confinement of microfiber, even for increased bending radius or index values [95]. In addition, *k* incorporates the signal attenuation due to the existence of FBGs, which has been proven to be in the order of 10% [96].

### **5. The road ahead**

From the presented analysis, it is evident that, despite their extensive applications, electrical neural stimulation techniques suffer from insufficient coding of spectral information, low power efficiency, low stimulation precision, accuracy, and frequency, as well as questionable safety. To this end, promising optical neural stimulation methods that surpass these limitations have been proposed. In an effort to establish these methods, the scientific community has pushed toward proving their feasibility as well as theoretically modeling and augmenting them. The state of the art of optical neural stimulation techniques offers great promise toward realizing nextgeneration biomedical systems.

One of the main offerings of optical neural stimulation is the outstanding stimulation precision it offers compared to electrical neural stimulation. In more detail, the increased precision can be translated into higher customization of the produced neural activity in two respects. First, the increased stimulation frequency that comes with optical neural stimulation leads to higher accuracy of excitation due to the fact that action potentials are delivered faster to the target spiral ganglion neurons and, therefore, to the brain, thus limiting the time between sound acquisition and perception. Second, optical neural stimulation depends on the optical particularities of light sensitive opsins with each one being expressed in a specific type of cell. Therefore, this offers another layer of light selectivity that can be leveraged by optical neural stimulation techniques [97]. The combination of these two aspects equips optical neural stimulation with the necessary tools to achieve unprecedented performance not only in the field of hearing restoration but also in other biomedical application such as retinal implants that would utilize this advantage to provide higher perceived image fidelity.

#### *Hearing Restoration through Optical Wireless Cochlear Implants DOI: http://dx.doi.org/10.5772/intechopen.104622*

Another aspect that boosts the performance of optical neural stimulation is the exceptional spectral coding of the information carried by the optical signal. On the contrary to electrical neural stimulation techniques that are characterized by wide current spread from the electrode contacts, optical radiation attenuates with a greater rate when it propagates inside human tissue, and therefore, the applied optical stimulations are more spatially confined than electrical ones. The importance of this phenomenon is highlighted even more by the fact that human sound perception requires at least 32 stimulation channels in order to recognize music or sound in noisy environments [89, 90]. As a result, the superior spectral coding of optical neural stimulation enables support for stimulation units that can house significantly more channels.

Contrary to previous detrimental improvements offered by optical neural stimulation methods, their performance in terms of power efficiency is comparable to the one of electrical neural stimulation. In more detail, optimization is required for optical neural stimulation stimulation policies in order to achieve similar power consumption as electrical neural stimulation [58]. Therefore, the optimization of optical neural stimulation techniques in terms of their power demands is one of the key requirements for their successful application in future biomedical applications. Similarly, the safety and ethical concerns of optical neural stimulation pose another controversial aspect. On the one hand, the optical power that is required for the reliable activation of light-sensitive spiral ganglion neurons is below the limits defined in various standardization protocols [98]; on the other hand, the modification of the targeted spiral ganglion neurons in order to acquire light sensitivity poses ethical concerns.

From a purely biological perspective, action potentials generated from electrical stimulation signals resemble the morphology and waveform of the membrane potential. As a result, these electrical signals are superimposed on each other and become almost indistinguishable, which hinders hearing restoration [58]. However, owing to its core functionality, optical neural stimulation triggers action potentials that differ significantly from membrane potential based on the stimulation protocol and the type of the excited cell. Specifically, not only the waveform of the generated action potential is affected by the amplitude and the duration of the stimulation, but also the instant release of ions when opsins are illuminated, which causes the membrane to react immediately. In addition, each opsin-cell-type combination is characterized by a distinct morphology of transmembrane potential and in conjunction with the wide variety of opsins available; they ensure the generation of a distinct action potential.

Finally, from an engineering point of view, the plethora of opsins that have been developed can highly impact the performance of optical neural stimulation biomedical applications. All future research in this field should take into careful consideration the selection of the applied opsin, as suboptimal ones may result in low stimulation precision and reliability, which, in turn, can determine whether the application is successful or not. The most important design choices include the compatibility with the target cell type, the amplitude and morphology of the resulting action potential, and the nature and the direction of the released ions.
