1. Introduction
The use of implantable medical devices is a common and indispensable part of medical care for both diagnostic and therapeutic purposes [
1,
2]. However, implantation of medical devices often leads to the occurrence of difficult-to-treat infections, because of the colonization of their abiotic surfaces by biofilm-growing microorganisms, which are increasingly resistant to antimicrobial therapies [
3]. A promising strategy to combat device-related infections is based on anti-infective biomaterials that either repel microbes so they cannot attach to the device surfaces, or kill them in the surrounding contact areas [
1,
2,
4,
5].
With the use of different medical devices, it is critical to ensure that the active/functional areas remain free of microbial contamination, and the surface functionalisation of the materials plays a major role. Healing efficiency may be improved through the manipulation of the material surface. For the development of new generations of highly effective medical devices, several requirements need to be fulfilled in regard to the material’s surface that interacts with the biological environment. These include biocompatibility, hydrophilicity/hydrophobicity, mechanical resistance, antimicrobial activity, as well as antioxidant and antifouling (anti-biofilm attachment) properties [
1,
6]. Smart coatings and surface morphology manipulations are, nowadays, the driving force for creating such active surfaces of medical devices. An active release strategy and advanced natural-bio-based principles are priorities [
2].
Amongst the most interesting basic materials for medical devices is silicone [
7]. The possibilities for its expanded use in medical applications are bright, considering its already successful utilization in orthopedics, as part of various catheters, drains and shunts, contact lenses, artificial organs, as components in kidney dialysis, heart-bypass machines, blood-oxygenators, and many others [
8,
9,
10].
Silicone is a commonly used material for tympanostomy tubes, whose insertion is a leading surgical procedure performed on children worldwide [
11]. A tympanostomy tube is implanted surgically—through an incision in the eardrum (myringotomy)—to allow drainage of fluid from the middle ear in patients with
otitis media (a group of inflammatory diseases of the middle ear). Despite this, persistent otorrhea (ear drainage) is the most common complication following tympanostomy tube insertion and can lead to tube occlusion and discomfort [
6].
General acute otitis media is, commonly, caused by microorganisms, including
Streptococcus pneumonia, Haemophilus influenza and
Moraxella catarrhalis, while
Staphylococcus aureus and
Pseudomonas aeruginosa are typically implicated microorganisms in ottorhea and are likely to have entered the middle ear via the auditory canal through the tympanostomy tube [
12].
Antibiotics have been shown to be effective for children with otitis, where the biofilm has not yet been formed. For children with chronic diseases, the medications do not work well enough. In this case, the more effective procedure is the myringotomy. Furthermore, otitis media with effusion and recurrent acute
otitis media are common problems in children, with a cumulative incidence of up to 80% by the age of 4 years [
12].
In addition, tympanostomy tubes are often used to relieve pathologic ear conditions such as Meniere’s disease and serous
otitis media [
13,
14,
15,
16]. When these tubes are inserted into the ear, complications include tympanic membrane perforation, extrusion, tympanosclerosis, post-tympanostomy otorrhea, and further infections, where the treatment is either a topical antibiotic or steroid treatment, or oral antibiotic [
12,
14,
15].
It has been shown that conventional routes of drug administration, such as oral or parenteral routes, are largely ineffective, mainly due to the blood-labyrinth barrier, which is a highly selective semipermeable border that separates the circulating blood from the brain and extracellular fluid in the central nervous system, and, as such, limits (or even prevents) the contact of drugs in the blood stream with the inner ear. Therefore, local (non-systemic) delivery of drugs can be significantly more efficient, and might even be essential for a successful treatment [
17,
18].
Since myringotomy still prevails as one of the most common surgical interventions for the pediatric population (also important for other patient populations) [
12], a major challenge of current research in this field is to find suitable types of active tympanostomy tubes’ functionalisation. In addition to the mechanical function to allow fluid flow and enabling unhindered ventilation of the ear, the functionalisation needs to provide an antimicrobial activity through an effective and controlled drug delivery. Antimicrobial functionalisation of the tube may indirectly reduce inflammation processes (through neutralization of the infection, which causes inflammation) as well as the incidence of biofilm formation on the surface of these tubes. An efficient drug delivery system would improve healing synergistically. A literature review on this topic revealed that only a very limited number of related studies were performed [
18], thus providing a large manoeuvrable space for research and development of novel related solutions, e.g., a smart coating for tympanostomy-silicone tubes acting as an antimicrobial surface with instantaneous drug delivery system as proposed in this study.
Tympanostomy tubes with a phosphoryl choline coating have been tested [
19]. During the treatment of the inflammatory process in coated and non-coated tympanostomy tubes, no significant differences were observed in the follow-up of 21 and 24 months. Although their study did not find any statistically significant differences between standard tubes and coated tubes, they indicate that their sample size (n = 70) may not have been large enough to allow efficient assessment of the coating and decide whether it leads to the desired improvement [
19].
The human serum albumin (HSA) was also used as a coating on standard tympanostomy tubes of various materials [
20,
21]. Fibronectin, a typical serum protein, which is one of the most adhesive glycoproteins, was used as a microorganism-blocking agent for tympanostomy. HSA coated tubes inhibited fibronectin binding from 59 to 85%, depending on the type of tube used. The study showed the potential of HSA coating in preventing the binding of pus and other undesired secretions in tympanostomy tubes [
20,
21].
Tympanostomy tubes from solidified polymer melts (Elvax and polyurethane) and [
1] three antimicrobial agents (ciprofloxacin, usnic acid and polyhexamethylene biguanide) for insertion into the tympanostomy membrane (ear drum) were fabricated and tested in terms of drug release kinetics as well as antimicrobial activity against
Pseudomonas aeruginosa,
Staphylococcus aureus,
Haemophilus influenzae and
Streptococcus pneumoniae. The release kinetic curves revealed two stages of antibiotic discharge, with released drug concentrations being above the minimal inhibitory concentrations (MICs) in the first six days of elution [
22].
Moreover, three brominated furanones (furanone-1, furanone-2 and furanone-3) applied to the surface of PVC as a tympanostomy tube material did not seem to be effective against
Staphylococcus aureus biofilm [
23] but one of the three, furanone-3, inhibited
Escherichia coli biofilm formation [
23].
Wang et al. investigated the effectiveness of an organoselenium coating on Donaldson tympanostomy tubes based on silicone material for inhibiting biofilm formation using an in vitro study.
Staphylococcus aureus and nontypable
Haemophilus influenzae formed considerable biomass on uncoated tympanostomy tubes, while the organoselenium tubes inhibited biofilm formation drastically, thus showing potential as a long lasting inhibitory agent [
24].
Ojano-Dirain et al. determined if biofilm formation on silicone tympanostomy tubes could be prevented by commercially available polyvinylpyrrolidone (PVP) or/and silver oxide coatings. The success of their approach was tested while exposing the prepared coatings to human plasma and to cultured
Pseudomonas aeruginosa or
Staphylococcus aureus. Biofilm formation after 4 days was assessed by quantitative bacterial counts and SEM, concluding that PVP and silver coatings reduce
Pseudomonas aeruginosa biofilm formation (PVP was superior to silver), while combining the PVP and silver coatings does not improve biofilm resistance further [
25].
Furanones were also loaded into microparticles [
26] and nanoparticles [
27]. Specifically, Cheng et al. [
27] used the halogenated 4-bromo-5-(bromomethylene)-2(5H)-furanone in combination with biodegradable poly-DL-lactic acid nanoparticles to fabricate a novel antibacterial coating for titanium (Ti) implants. Ti material may also be used for tympanostomy tubes. The mentioned antibacterial coating also exhibited an inhibition against
Staphylococcus aureus throughout a 60-day study period, which is considered a period long enough to prevent implant-related infections in the early and intermediate stages of device implantation. The antibacterial rate was about 100% in the first 10 days and 90% in the following 20 days [
28].
Although researchers have attempted to prevent biofilm formation using tympanostomy tubes coated with various antimicrobial compounds (mostly using antibiotics, synthetic polymers and/or inorganic substances), there is still, currently, no type of tympanostomy tube to which bacteria will not adhere [
6]. Furthermore, most of the existing solutions are either fully or partially cytotoxic, due to which they may cause an allergenic reaction by contact during use. Therefore, the use of natural, biodegradable and functional polymers is of high interest for scientific and industrial communities.
Chitosan, a natural amino polysaccharide, has been explored widely for biomedical applications like tissue engineering, gene therapy, wound healing and drug delivery. It is particularly attractive in the form of nanoparticles or nanocapsules, as they can provide some advantages over a chitosan coating alone. These include improved antimicrobial efficiency due to their small size, resulting in a high active surface area to volume ratio, as well as their ability to control the release of active agents, and their contribution to improved mucoadhesive properties [
29,
30].
In our previous work it was shown that chitosan nanoparticles themselves may be especially attractive as a surface coating for different medical materials such as cellulose fibres, silicone catheters, polyethylene terephthalate (PET) vascular grafts, etc. [
10,
29,
31,
32]. Moreover, chitosan nanoparticles were also used successfully as a delivery system for iodine and drugs previously attached onto cellulose medical textiles [
33,
34].
In this paper, chitosan nanoparticles were prepared by ionic gelation, which is a simple, non-time-consuming preparation procedure that can be performed at mild reaction conditions and is additionally free of any toxic reagents. These nanoparticles were used to encapsulate the co-amoxiclav (CoAM) drug mixture, and their consequent adsorption onto O
2 plasma-activated tympanostomy-silicone tubes was checked. Plasma activation was used to turn the hydrophobic surface character of the tubes into a more hydrophilic one, thus improving the chitosan nanoparticles’ attachment. In our previous work, it has been already shown and discussed that advanced and environmentally friendly cold O
2 plasma treatment can be used to enhance the adhesion of chitosan onto inert polymer materials [
17,
35]. The functionalised silicone material was analysed regarding the surface elemental composition, morphology, coating stability, drug release performance and antimicrobial activity. It has been shown that, either the chitosan-nanoparticle-based coating alone, or as part of a drug delivery system, acts as a promising functional layer for tympanostomy-silicone based tubes.
2. Materials and Methods
2.1. Materials
Chitosan (low molecular weight) and Sodium tripolyphosphate (TPP, purum p.a., ≥ 98.0%), were purchased from Sigma-Aldrich (Taufkirchen, Germany). Acetic acid was purchased from Honeywell (Seelze, Germany). Co-amoxiclav was obtained from Lek (Ljubljana, Slovenia) and normal saline (0.90% w/v NaCl) from B. Braun (Melsungen, Germany). A silicone elastomer base and silicone curing agent kit SYLGARD® 184 was purchased from DOW (Wiesbaden, Germany), tryptic soy agar (TSA) plates from Millipore (Wien, Austria) and polyethylene terephthalate mesh from Beti (Metlika, Slovenia). All chemicals and materials were used as received, without any further purification. Ultrapure water (with a resistivity of 18.2 MΩ cm, obtained from Milli-Q, Millipore Corporation, MA, USA) was used throughout the experiments.
2.2. Preparation of Chitosan, TPP and CoAM Solutions
An appropriate amount of chitosan was dissolved in Milli-Q water to prepare a 1% (w/v) solution. The pH of the solution was adjusted to 3.8 initially with acetic acid, followed by constant overnight stirring. Acetic acid was then added to adjust the pH to 3.5 and the total volume of the solution was filled to the required amount.
TPP was suspended in Milli-Q water in order to prepare a 0.2% w/v solution.
Co-amoxiclav was used as a model drug, composed of amoxicillin (amoxicillin sodium) and clavulanic acid (clavulanate potassium) in a 5:1 w/w ratio. 20 mL of normal saline was added into the vial with the drug (1000 mg/200 mg), and shaken for a further 10 min.
2.3. Preparation of Bare and Loaded Chitosan Nanoparticles
Chitosan nanoparticles were prepared by the ionic gelation technique. Simultaneously 0.2% (
w/
v) of TPP solution was added to a fixed volume of 1% (
w/
v) chitosan solution, in order to obtain a 5:1 chitosan to TPP weight ratio. This ratio was chosen according to the previously published work, reporting it as an optimal ratio for obtaining the desired antimicrobial activity of nanoparticles’ dispersion [
36]. Particles were formed spontaneously under magnetic stirring for 1 h at room temperature. The final pH of the chitosan nanoparticles dispersions was adjusted to 4.0 by the addition of concentrated acetic acid.
Co-amoxiclav embedded chitosan nanoparticles were prepared as follows. 10 mL of chitosan solution was added to a 50 mL beaker, following constant mixing. Further, 10 mL of TPP was added, and, at the same time, 10 mL of co-amoxiclav previously dissolved in saline. Nanoparticles with encapsulated drug formed spontaneously and the dispersion was stirred for 30 min.
2.4. Preparation of Silicone Material and O2 Plasma Treatment
Polydimethylsiloxane (PDMS), also known as dimethylpolysiloxane or dimethicone, which belongs to a group of polymeric organosilicon compounds, was used as a representative silicone material. In a 200 mL plastic crucible, 90 g of a silicone elastomer was weighed, to which 10 g of a silicone curing agent was added. The solution was mixed, followed by pouring the resulting solution into plastic molds. These were dried in a vacuum oven for 24 h at 80 °C. A uniform distribution of the polymer solution was required to ensure that silicone plate samples were as comparable as possible. These thin silicone plates simulated the tympanostomy tubes and due to their dimensional and geometrical characteristics, serve as ideal model platforms, allowing for an easier physicochemical characterisation of functional silicone-based materials.
Silicone plates were cleaned, dried and cut to a required size for further activation with O2 microwave plasma in a surfatron mode. The forward power was set to 2 kW (I = 0.3 A). O2 pressure in the treatment chamber, made of quartz glass was set to 30 Pa. The samples were exposed to O2 plasma treatment for the following times: 1 min, 2 min, 3 min and 5 min, respectively. The experiments were done at the Institute “Jožef Stefan”, Ljubljana, Slovenia.
2.5. Application of Chitosan Nanoparticles onto Silicone Material
Dispersions were applied to the top side of the silicone plates using an airbrush (SP-575, Sparmax, Taiwan). The nanoparticles’ dispersions were applied to inactivated and plasma-activated samples. The appropriate pressure in the airbrush tubes for application of the sample was regulated through a connection with a nitrogen cylinder. Functionalised silicone samples were stored in a refrigerator at a temperature of 5 °C, due to the drug’s sensitivity to high temperature. Sample description and notation is given in
Table 1.
2.6. Characterisation of Dispersions
2.6.1. Evaluation of Hydrodynamic Diameter, PDI and ζ-Potential
The average particle hydrodynamic diameter (dh) and polydispersity index (PDI) of the prepared nanoparticle dispersions (CN and CN-CoAM) were determined by dynamic light scattering and through the ζ-potential (ZP) measurement using a Zetasizer Nano ZS (Malvern Instruments, Worcestershire, UK). Samples were injected into disposable cuvettes (DTS0012, Malvern Instruments, UK) for dh and PDI measurements or folded capillary cells (DTS1070, Malvern Instruments, UK) for ZP determination. Prior to analysis, the samples were stirred for 15 min and their pH adjusted to 4 using acetic acid (0.1 M), if necessary.
2.6.2. Drug Encapsulation Efficiency
The drug encapsulation efficiency (EE) was determined indirectly after separation of CN-CoAM nanoparticles from medium, containing non-encapsulated co-amoxiclav using a centrifugation-based technique. The percentage of encapsulation efficiency of co-amoxiclav in the nanoparticles was measured using UV-Vis spectrophotometry (275 nm, Cary 60 UV-Vis spectrophotometer, Agilent Technologies, Santa Clara, CA, USA) and calculated as follows:
where
C0 and
Cs are total drug concentration used to prepare the particles and the concentration of co-amoxiclav present in the supernatant after centrifugation, respectively. Standard co-amoxiclav calculation curve was plotted beforehand using Milli-Q water.
2.7. Surface Characterisation
2.7.1. X-ray Photoelectron Spectroscopy
In order to assess the surface of the sample (functionalised silicone material), XPS spectra were recorded (PHI TFA XPS Physical Electronics, Chanhassen, MN, USA). The base pressure in the XPS analysis chamber was 10−8 Pa. The samples were excited with X-rays using monochromatic Al Kα1.2 radiation (1486.6 eV) operating at 200 W. Photoelectrons were detected with a hemispherical analyser, positioned at an angle of 45° with respect to the normal to the sample surface. The energy resolution was about 0.6 eV. Spectra were recorded from at least two locations on each sample, using a 400 μm analysis area. Surface elemental composition were calculated from the survey-scan spectra using the Multipak software (ULVAC-PHI, Inc., Chigasaki, Japan)
2.7.2. Scanning Electron Microscopy
Surface morphology was evaluated using scanning electron microscopy (SEM). Prior to imaging, the samples of silicone materials were prepared by cutting the foils into small, approximately 50 mm2 × 50 mm2 square pieces, which were attached to aluminium sample holders using an adhesive carbon tape to ensure conductivity. Imaging was performed using an SEM (FEI Quanta 200 3D, Hillsboro, OR, USA, whilst the samples were analysed at an accelerating voltage of 1 kV, and at a variable working distance (4–5 mm) using different-sized apertures.
2.8. Evaluation of Antimicrobial Activity
To perform microbiological evaluation, the bacterial growth after exposure to respective samples was used to determine the effectiveness of the prepared coated silicone materials. Microbiological testing of the antimicrobial activity of the as-prepared coated silicone plates was carried out using a direct (bacteria attachment to the samples) and indirect (exposure of the bacteria culture to sample extracts) method. For this purpose, a
Staphylococcus aureus test strain (DSM 799) was used, where the tests were performed according to the internal protocols of the Department of Microbiological Research, Centre for Medical Microbiology of the National Laboratory for Health, Environment and Food in Maribor, i.e. No. P96 Biofilm production on various materials—
Staphylococcus aureus’. Briefly, sample antimicrobial activity was tested by determining the number of colony-forming units (CFU) in the supernatant after exposure to respective samples. Variously functionalised silicone materials in the size of 10 mm × 10 mm were exposed to the standardized medium inoculated with
Staphylococcus aureus and adjusted to 0.5 on the McFarland scale [
37]. After 4 h of incubation at a temperature of 37 °C and slight shaking, the obtained sample was diluted appropriately and inoculated to a TSA plate. Colonies were counted after incubation on agar plates at appropriate conditions after 24 h and after one month. The effect of the functionalised silicone material was determined as a reduction in growth, where the number of bacteria after exposure to a non-functionalised silicone material was compared to the functionalised one.
2.9. In Vitro Drug Release Testing
In vitro drug release studies were performed using an Automated Transdermal Diffusion Cells Sampling System (Logan System 912-6, Somerset, KY, USA). The drug-loaded samples PDMSCN-CoAM, PDMSPA1, CN-CoAM, PDMSPA2,CN-CoAM, PDMSPA3, CN-CoAM, PDMSPA5, CN-CoAM were cut into 10 mm × 10 mm squares and placed on the top of a PET mesh. The receptor compartment was filled with Phosphate Buffered Saline (PBS, purchased from Sigma-Aldrich, Germany) with a pH value of 7.4 and its temperature was maintained at 37 °C. During the dissolution testing the medium was stirred continuously with a magnetic bar. Samples were collected over a period of 24 h at different time intervals (1, 5, 10, 20, 30, 60, 120, 180, 240, 300, 360 and 1440 min), while the released/dissolved CoAM concentration in the receptor medium was determined by a UV-Vis spectrophotometer (Cary 60 UV-Visible Spectrophotometer, Agilent, Germany) by quantification of the absorption band at 276 nm. The withdrawn sample volumes were replaced by fresh PBS with a stable temperature of 37 °C of the same volume. Sink conditions were assured due to sample withdrawal, followed by sample dilution through media replacement. In calculation of concentrations using the Beer–Lambert Law, this dilution was accounted for. All release studies were performed in triplicates and are reported as average value with standard errors.