**1. Introduction**

Cochlear implants (CI) are currently the most effective treatment options for severe to profound hearing loss. During cochlear implantation, an electrode array consisting of different numbers of platinum contacts on a silicone carrier is inserted into the scala tympani of a cochlea. Cochlear nerve cells, the spiral ganglion neurons (SGN), can then be electrically stimulated by application of pulses of constant current. Clinical results with CI are typically good; for example, most patients can communicate via the telephone again [1]. Nevertheless, there are several known limitations. First, after hearing loss, SGN also start to degenerate [2]. Second, for insertion of a CI electrode, the cochlea has to be opened and the electrode array is positioned in the scala tympani. This causes some additional trauma which is considered to be a risk for surviving SGN [3]. As a reaction of the human body to this trauma, but also to the implanted foreign body, fibrous tissue is formed around the electrode array [4]. As shown in postmortem studies, the amount of tissue formation can be variable from a few cells to the formation of new bone [5]. The increase in electrical impedance at the stimulating contacts, as reported for the first two to three weeks after implantation [6], was shown to be correlated with the tissue response after implantation [7].

**Citation:** Wulf, K.; Goblet, M.; Raggl, S.; Teske, M.; Eickner, T.; Lenarz, T.; Grabow, N.; Paasche, G. PLLA Coating of Active Implants for Dual Drug Release. *Molecules* **2022**, *27*, 1417. https://doi.org/10.3390/ molecules27041417

 Academic Editors: Marek Brzezi ´nski and Małgorzata Ba´sko

Received: 3 February 2022 Accepted: 17 February 2022 Published: 19 February 2022

**Publisher's Note:** MDPI stays neutral with regard to jurisdictional claims in published maps and institutional affiliations.

**Copyright:** © 2022 by the authors. Licensee MDPI, Basel, Switzerland. This article is an open access article distributed under the terms and conditions of the Creative Commons Attribution (CC BY) license (https:// creativecommons.org/licenses/by/ 4.0/).

Furthermore, when the tissue formation is not uniform along the electrode array, it might also affect the specificity of the electrical stimulation.

Currently, there are several approaches under investigation to reduce trauma and the formation of fibrous tissue after cochlear implantation. Amongst them are surface patterning of the electrode array [8], application of drugs via pumps [9], cells [10], coatings [11], or from a reservoir, such as the silicone of the electrode array [12,13], as well as intraoperative deposition of steroids either directly [6] or by using a catheter [14]. To the best of our knowledge, besides one report on three patients receiving mononuclear cells obtained from bone marrow with the cochlear implant [10], only intraoperative deposition of steroids and steroid elution from the silicone of CI electrodes have been used clinically so far [12,14] and were shown to reduce or delay the impedance increase after implantation. As elution from the silicone results in a slow release [15], combination with a faster release from a surface coating might be a promising way to effectively address the tissue reaction right after implantation and in the long term.

The release of active substances can basically be divided into two types: diffusioncontrolled drug release and chemically controlled drug release [16]. Diffusion-controlled release is further divided into membrane-associated and matrix-associated release. No matter which of these two is considered, both behave according to Fick's first law of diffusion [16]. Characteristically, the release depends on the concentration gradient. At the beginning, the drug-release system is fully loaded whereas the tissue environment does not contain any drug. This results in a so-called initial burst release, a strong increase in concentration of the active ingredient in the tissue. In the further course of time, the release continues to level off.

In contrast, chemically controlled drug release requires steps that occur before the actual release. In so-called "swelling-controlled systems", the active ingredient is distributed in a polymer matrix but cannot diffuse out of the material, e.g., due to a small pore size. After a solvent is added the polymer swells, causing the pores to enlarge in such a way that diffusion is no longer inhibited, and the active ingredient is released. In degradation-controlled systems, bonds have to be cleaved before the active ingredient is released. These bonds belong to the polymer in which the active ingredient is incorporated. Furthermore, there are systems in which the active ingredient is covalently bound to the polymer, e.g., as side chains. Hence, there is also a bond that has to be cleaved, before the active ingredient can diffuse out [17–19].

For dual drug release—which means the release of two active ingredients, each exhibiting a different release behavior—it is advantageous to use different mechanisms of release control. Therefore, drugs incorporated in polymer matrices can be used together with, e.g., drugs that are covalently immobilized at the surface of the dedicated polymer. Different polymers with different diffusion-controlled properties, e.g., different pore sizes, may also be used.

This principle was used to combine vascular endothelial growth factor (VEGF) and paclitaxel for application to the cardiovascular system [20,21]. There, paclitaxel was incorporated into a PDLLA polymer coating and VEGF was covalently attached to the surface of the polymer on either films or nanofiber non-woven. The resulting dual drug release improved endothelial cell viability in vitro, even in the presence of paclitaxel, which alone resulted in significantly decreased viability.

Furthermore, polymers with different diffusion properties can be used to obtain a dual drug release. For example, a silicone matrix can be loaded with an active ingredient that shows a relatively slow release. A faster-releasing coating on the silicone body then leads to either a faster release of the same or a different drug [15].

Dexamethasone (DMS) was already incorporated in coatings intended for CI [11,22]. Growth factors and other substances applied to the cochlea and being released from a coating were IGF1, HGF [23], BDNF [24], NT-3 [25] and Ara-C [22]. Most of these examples were intended to enhance the survival of SGN but not for reduction of fibrous tissue formation. A further search for additional substances that might reduce the inflammatory reaction and are already approved for other applications revealed diclofenac (DCF) [26,27] and enalapril [28,29] as possible candidates. DCF is a nonsteroidal anti-inflammatory drug. Its anti-inflammatory action can be explained by the inhibition of the cyclooxygenase in vitro and in vivo [30,31]. Enalapril is an angiotensin-converting-enzyme inhibitor and can reduce local inflammation after myocardial infarction [32].

Therefore, the aim of the current study was to investigate DCF and enalapril for their safety when applied to cells from the inner ear. In a second step, these substances should be included in a surface PLLA coating for a fast initial release and this PLLA coating shall be combined to DMS-loaded silicone of the electrode array for a slower long-term release of DMS. Release characteristics and the influence of the coatings on electrode contact impedances were investigated.
