The artificial steroid dexamethasone is known for its potent glucocorticoid effects:
anti-inflammatory, immunosuppressant and causing remarkable inhibition of fibroblast growth (RAMALINGAM et al. 1997). However, many studies on steroidal action on the proliferation of fibroblasts in vitro have led to contradictory results with either an inhibition of the proliferation or a stimulatory effect. Durant et al. found that most discrepancies may be due to the use of different experimental models, schedules, culture conditions, to the various methods of proliferation monitoring and to the choice of different cell lines (DURANT et al. 1986). Still, most in vivo studies or such with primary cultures have led to similar conclusions of glucocorticoids inhibiting both fibroblast proliferation and collagen synthesis. Thus, the effect seems to be proved for the application conditions of cochlear implants. Also, many findings in human studies imply that steroids like Dex are considerable agents for optimization of cochlear implant applications (PAASCHE et al. 2006; PAASCHE et al. 2009).
However, most application methods, such as one-shot injections via syringe or cannula, are based on short term release without the possibility of affecting the long-term tissue reactions like fibrosis. Even polymer or hydrogel formulations providing a prolonged drug release cannot last longer than several days (ENDO et al. 2005; LEE et al. 2007; DINH et al. 2008; PAULSON et al. 2008; SALT et al. 2011). Applied intratympanically, the rapid loss of the drug through the Eustachian tube and the high concentration gradient between basal and apical cochlear turns are further limitations (PLONTKE et al. 2008; SALT u. PLONTKE 2009). Currently available long term release methods like mini-osmotic pumps, microcatheters and the MicroWickTM give direct access to the inner ear tissues for active drug application with lower concentration gradients, but also for no more than several weeks (KOPKE et al.
2001; RICHARDSON et al. 2006; SWAN et al. 2008). Additionally, since the drug reservoir has to be refilled frequently, it increases the risk of infection of the inner ear.
The invasiveness of the therapy, the required time, and the costs should also come to consideration.
Using the cochlear implant itself as a drug delivery device, no additional surgery is required and direct access to the cochlear structures is granted without the limitations of the basal-apical concentration gradient of round window applications (RICHARDSON et al. 2008). Several research groups developed a modified cochlear implant with incorporated drug delivery channels (PAASCHE et al. 2003a; permeability for lipophilic agents (particularly steroids) (BAKER 1987). Crystalline Dex enclosed in the PDMS-matrix was shown to be released over a time period of 3 months with the potential for even longer time periods referring to the diffusion kinetics, which revealed the release of only 4.4% (non-coated) and 3.36% (hydrogel-coated) of the embedded drug amount within 3 months. Our results are consistent with the findings of Ghavi et al., who studied Dex release from silicone rubber CI coatings with different weight percentages up to 2% w/w (FARAHMAND GHAVI et al.
2010). A release similar to our findings (5% w/w) was revealed over a time period of 21 months, indicating an analogous long term release for our system. Assuming a volume of the human cochlear perilymph of 160µl (BUCKINGHAM u. VALVASSORI 2001) and a delivery system weight of 1mg, from a delivered dose of 5.87ng per day results a concentration of 36.7µg/ml achieved at the target location directly after implantation. But this value only applies to a closed fluid system like the cell culture is. Considering the complex mechanisms of drug clearance from the cochlear fluids as well as the perilymph loss due to implantation, further computer simulations (SALT 2005) and in vivo experiments should be performed.
For the second biochemical functionalization to prevent unspecific protein adsorption, encoating of the silicone matrix with sPEG was chosen. It constitutes only a small
diffusion barrier for Dex. However, the coating reduces the release to about 23% by changing the diffusion coefficient (Figure 1). As it is also the case in other implantation surgeries, the acute wound caused by the insertion of CI electrode into the inner ear is suggested to undergo a healing process with a time course of maximally 30 days, divided into four time dependent stages (coagulation/haemostasis, inflammation, proliferation, wound remodelling)(VELNAR et al. 2009). A high release rate during the first weeks after implantation is therefore beneficial for the initial healing process and although the initial release from coated PDMS is lower and more constant, it is still recognizably increased (Figure 2). The retardation of release in the coated filaments provides a slower and longer drug delivery with sufficient concentration levels (for exceeding 0.03-0.04µg/ml (SALT et al. 2011), weight percentage of Dex or size of the delivery system could be adjusted) for long term therapy, with positive effects over several years (WISH et al. 1990;
ANDERSON et al. 1991; MOND u. STOKES 1996; PEETERS et al. 1998; DE CEULAER et al. 2003; PAASCHE et al. 2006; PAASCHE et al. 2009). The protein repelling effect is due to increased hydrophilicity and degree of hydration (OSTUNI 2001) and could be proved through adsorption of TAMRA marked Bovine Serum Albumin on untreated and hydrogel coated silicone filaments (Figure 4). Compared with non-coated filaments both filaments with and without Dex equipped with a hydrogel layer showed no unspecific protein adsorption onto their surfaces. This is an indication that the incorporation of Dex inside the PDMS matrix does not influence the protein repelling properties of the hydrogel layer.
Electron microscope pictures show a homogeneous distribution of the drug throughout the matrix (Figure 7), while simultaneously flexibility and mechanical stability of the material are not restricted. Nano-roughness of the filament surfaces detected by AFM was significantly affected by different loadings and coatings compared to untreated PDMS (Figure 5). Dex-loaded filaments showed the highest nano-roughness as well as the most uneven surface with small cavities. This may be due to the enclosed Dex crystals, while the crystals in PDMS+sPEG+Dex samples are covered with hydrogel, which therefore smoothens the surface but is still rougher than PDMS alone. The fine fissures in the hydrogel-coated filaments (Figure 6) could
be caused by desiccation or mechanical manipulation. The manufacturing process is assumed to be the reason for the linear grooves found in all filament types.
Our in vitro model examined eGFP fibroblast growth in contact with Dex-loaded, unloaded, sPEG-coated and uncoated PDMS filaments. Within experimental setting I, growth of fibroblasts on the surfaces of the filaments and of the bottom of the wells was evaluated (Figure 8). Since only in the wells with Dex and Dex+sPEG filaments the cell growth on the filament surfaces as well as in the vincinity of the filaments (well bottom) was affected (reduced by 70%), we proved the release and diffusion of Dex from PDMS as well as from the sPEG coated PDMS. Hydrogel itself has an antiproliferative effect (GROLL et al. 2004; GROLL et al. 2005a; GROLL et al. 2005b;
GROLL et al. 2005c), however, without contact to the cells on the well bottom it had no influence on them. Through diffusion into the cell culture medium, only Dex accounts for the reduced cell growth on the well bottom.
Transferred to the circumstances in vivo, the diffusion of the drug into the perilymph would not only affect the cells in direct contact with the electrode, but also the connective tissue cells surrounding the electrode inside the scala tympani.
Additionally, this benefit also applies to the platinum electrode-nerve interfaces, which are not coated by PDMS. The slightly higher cell number in wells with PDMS and sPEG-coated PDMS than in the control wells could be explained through the additional growth area on the filaments presented to the cells. The surface of the well bottom including the filaments’ surface was increased by nearly 17% (300µm diameter) and 29% (500µm diameter) compared to the well bottom surface alone.
Considering variations in cell distribution and different positions of the filaments, the increase of cell number by 12-17% matches these circumstances.
Considering only the surfaces of the filaments (setting II), we found a cell number reduction of over 90% on all treated filaments in comparison to the untreated PDMS (Figure 10). There, not only the Dex-loaded filaments were nearly cell free, but also the sPEG-coating alone showed a remarkable cell repelling effect. Still, highest reduction (up to 99%) of cell proliferation was achieved by combining Dex and sPEG.